Keywords: drug delivery, nanocarrier degradation, NMR spectroscopy, metal-organic framework, polymer-drug conjugate, in situ NMR

Controlled drug delivery technology has improved medical treatment. A drug delivery system (DDS) should transport and deliver the active molecules progressively with time and at the determined location in the patient body. Nanoparticle-based DDSs have promising abilities to overcome the physiochemical properties of the drug and bypass biological barriers, yielding complete control of the drug delivery process. Nonetheless, the degradation and reactions of nanomaterials in biological media are still burning questions hampering the clinical translation of nanoparticles (NPs). Here, we investigate the degradation process in biological-mimicking media of two types of nanocarriers made of polymer and metal-organic framework. Various characterization techniques are used with a particular focus on nuclear magnetic resonance (NMR) spectroscopy.

First, we described a new approach to preparing polymer-drug conjugate NPs and studied their degradation. Following base-catalyzed ester hydrolysis and bulk erosion, poly(lactic acid)-itaconate PLA-ITA NPs at physiological pH degraded into different oligomers of lactate and itaconate. Liquid NMR spectroscopy and mass spectrometry determined the nature and quantity of water-soluble reaction products and residual polymer. Electron microscopy gave information on the size and morphology of NPs during their degradation. The discovery of slow degradation, hence slow drug release of PLA-ITA NPs, gives suggestions for better design of pH-responsive DDSs.

Second, we conducted an in-depth investigation on the degradation of two nanoMOFs (metal-organic framework NPs), nanoMIL-100(Al) and nanoMIL-100(Fe), in phosphate buffer at pH 7.4. Using a multiscale analysis, we found a stepwise degradation mechanism in which phosphate successively substitutes water and trimesate ligand of nanoMOF, leading to the formation of aluminum/iron phosphate and the release of free trimesate ligands. The role of phosphate was proved by liquid NMR and X-ray photon electron spectroscopy. The degraded nanoMIL-100 particles were analyzed by solid NMR and X-ray absorption near edge (XANES) spectroscopy. Furthermore, the size, morphology, and surface roughness of MIL-100 NPs were monitored during degradation by small-angle X-ray scattering (SAXS) and electron microscopy (TEM and SEM), supporting information on different erosion processes of aluminum-and iron-based MOFs. The molecular motions of water in the nanoMOF aqueous suspension were studied by NMR relaxometry to probe how the degradation changes the interior and exterior surface of nanoMOF. Additionally, the degradation of drug-loaded and surface-coated nanoMOF with phosphate-bearing molecules was demonstrated as a function of metal-ligand complexing strength. The obtained knowledge about nanoMIL-100 degradation contributes to nanoMOF chemistry and lays the basis for further degradation study in complex scenarios.

Finally, an in situ NMR spectroscopy method was set up to track the degradation of NPs without separating them from their suspension. It delivered both qualitative and quantitative information about the nanoMIL-100(Al) degradation. The kinetics of liquid-phase trimesate release and solid-phase trimesate loss were simultaneously determined. Coordination of aluminum with phosphate was also detected, consistent with the proposed degradation mechanism. The established method was then applied to the paramagnetic nanoMIL-100(Fe) and the drug-loaded nanoMIL-100(Al). The presented in situ NMR method offers an option for accurate characterization of NPs' dynamic processes without altering the reaction medium.

Overall, the blend of analytical methods used in this thesis contributed some knowledge on the degradation mechanisms of polymer and metal-organic framework materials, along with examples of using NMR spectroscopy to characterize DDSs. Titre : Mécanismes de dégradation des nanovecteurs à base de polymères conjugués aux molécules actives et de MOFs Mots clés : administration des médicaments, dégradation de nanovecteur, spectroscopie RMN, solide hybride poreux métal-organique, conjuguées polymère-médicament, RMN in situ Résumé : Les technologies d'administration contrôlée de médicament ont révolutionné le traitement de nombreuses maladies. Un système d'administration de médicaments ("drug delivery system", DDS) doit transporter et délivrer les molécules actives avec un contrôle à la fois dans l'espace (site malade) et dans le temps. Les DDS à l'échelle nanométrique ont la capacité prometteuse de pallier les inconvénients liés aux propriétés physiochimiques défavorables du médicament et de contourner les barrières biologiques, ce qui permet un contrôle complet du processus d'administration du médicament. Néanmoins, la dégradation et les réactions des nanomatériaux dans les milieux biologiques restent une question brûlante qui entrave la traduction clinique des nanoparticules (NPs). Nous étudions ici le processus de dégradation de deux types de nanovecteurs à base de polymères et de réseaux métalloorganiques ("metal-organic framework" ou MOF). Diverses techniques de caractérisation sont utilisées avec un intérêt particulier pour la spectroscopie par résonance magnétique nucléaire (RMN).

Tout d'abord, nous avons décrit une nouvelle approche pour préparer des NPs de conjuguées polymère-médicament et étudié leur dégradation et leur libération de médicament. Suite à l'hydrolyse de l'ester, les NPs de poly(acide lactique)-itaconate PLA-ITA à pH physiologique se dégradent en oligomères mixtes d'acide lactique et d'acide itaconique. L'identification et la quantification des produits de dégradation ont été réalisées à l'aide de la spectroscopie RMN liquide et la spectrométrie de masse (ESI et MALDI-TOF). La taille et de la morphologie des NPs ont été caractérisées par microscopie électronique. La dégradation s'est avéré lente, ce qui entraîne une libération lente du medicament.

Deuxièmement, nous avons mené une étude approfondie sur la dégradation de deux nanoMOFs de type nanoMIL-100(Al) et nanoMIL-100(Fe) dans un tampon phosphate à pH 7.4. En utilisant une analyse multi-échelle, le mécanisme de dégradation par des phosphates a été dévoilé. Les phosphates remplacent les deux ligands natifs des nanoMOF, suivi par la formation de phosphate d'aluminium/fer et la libération de ligands trimésate. Le rôle du phosphate a été mis en évidence par la RMN liquide et solide, la spectrométrie photoélectronique X (XPS), et la spectroscopie d'absorption au niveau du seuil X (XANES). En outre, la taille, la morphologie et la rugosité de surface des nanoparticules ont été suivies pendant la dégradation par diffusion des rayons X aux petits angles (SAXS) et par microscopie électronique (MET et MEB), ce qui a permis d'obtenir des informations sur les différents processus d'érosion des MOFs. Pour sonder les mouvements moléculaires de l'eau et notamment aux abords des surfaces, la relaxométrie RMN a été employée. Cette dernière renseigne sur les changements de l'état de surface interne et externe des nanoMOFs. De plus, la dégradation des nanoMOFs chargés et recouverts de molécules phosphatées a été démontrée en fonction de la force de complexation métalphosphate.

Enfin, une méthode de spectroscopie RMN in situ a été mise en place pour suivre la dégradation des nanoMIL-100(Al) sans séparer les particules de la suspension. Les cinétiques de libération du trimésate en phase liquide et de perte du trimésate en phase solide ont été déterminées simultanément. La coordination de l'aluminium avec le phosphate a également été observée, renforçant le mécanisme de dégradation proposé. La méthode a été ensuite appliquée au le nanoMIL-100(Fe) paramagnétique et le nanoMIL-100(Al) chargé de médicament. La méthode de RMN in situ présentée ici permet la caractérisation précise des processus dynamiques des NPs dans leur milieu naturel sans altération.

La panoplie des méthodes analytiques employées dans cette thèse a élucidé le mécanisme de dégradation des nanvecteurs de médicaments sous forme de polymères et MOF.

Remember to have FUN!!! \m/\m/\m/

Acknowledgement

When writing my Bachelor's and Master's thesis, the acknowledgment was one of the easiest parts. However, for this PhD thesis, it is one of the most challenging parts to write as I am not good at expressing my emotions. Three years is long enough to get attached to something, yet somehow still short to get a deep connection with someone with plain words like me. And so, I am grateful for everyone with whom I have gotten the chance to be connected. The kindness and patience of people have never ceased to surprise me. This acknowledgment is by no means an exhaustive list. If you do not see your name here, it does not mean you are not important to me and my thesis; it is just that my mind was not perfect when I was writing down these thoughts, and I apologize. I will keep all the memories with everyone involved with my PhD.

I have learned many things about science and life during this PhD study. Most of the lessons must come from my supervisors. I would like to express my gratitude to my supervisors, Mohamed Haouas, Ruxandra Gref, Charlotte Martineau-Corcos, and Anne-Laure Rollet, for guiding me through my PhD. They have given me priceless freedom, which has given me the confidence and the courage to speak out opinions. They are always welcoming and patient with my long list of questions, curiosities, and issues.

It is my honor, and it has been interesting to get to know and work with four different experts and personalities. I will remember the 3-4-hour meetings that we had from time to time. I thank all four of my supervisors for making me a better researcher and scientist in their own ways.

Mohamed is the one from whom I got to learn and practice NMR spectroscopy, especially liquid-state NMR. He helped me set up experiments and operate the NMR instruments. As a supervisor, Mohamed gives me a lot of freedom and authority to make decisions on my project and still gives me advice on each step. I feel comfortable talking with him about everything without hiding. Thank you for always being available for all my little questions every day! Moreover, I consider him the savior of my Ph.D., and I am not exaggerating. I cannot thank him enough for taking over the project when Charlotte decided to leave the university, even though, at that point, he knew nothing about me and drug delivery. I also learned a lot about communication from the many times I tried to explain what I wanted to do, and he tried to explain what I could not do. He made me slow down and listen to other people. Another thing I will remember about Mohamed is how he believes in my capabilities. He has continuously provided compliments, encouragement, and criticism for me and my work. He was determined to help me with my crisis regarding the meaning of my PhD, especially during the final year. We might not choose each other to be PhD student and supervisor at the beginning, but we achieved a great deal in the end! Ruxandra is the person who introduced me to the world of drug delivery. Four years ago, she agreed to take me as a Master intern and, after a misadventure, proposed a collaboration with Charlotte. She spends enormous amounts of time speaking to each of her many students. She has many visions and is always enthusiastic and encouraging toward my data and ideas. Her comments and questions motivate me to do better research. I think I inherited from Ruxandra her ambition and curiosity to do multiple projects and techniques, as shown in my thesis. She also forgave my mistakes and was willing to give me chances to improve myself. Through my mistakes, she taught me to deal with liberty and autonomy with responsibility and to be a better team player. On a side note, I will surely miss her gifts every Christmas.

Charlotte is the one who taught me everything I know about solid-state NMR since day one of the project Master 2 internship and PhD. Charlotte is quick-witted in experiments, discussions, and decision-making, which makes me strive to catch up with her. I wish I had had more time working with her in the lab. Thank you for trusting me with your proposal and letting me steer the original project the way I like! Thank you for keeping calm when I freaked out about losing the expensive NMR rotors and breaking the spectrometer! I also like her frank comments and opinions because I never have to speculate anything. Although she could not follow me closely for the whole three years as in our original plan, she was always present for me at critical moments.

Anne-Laure is an expert on high and low frequency and more. Her carefulness and organization towards doing science are admirable. Even though we worked together only for a short time due to three lockdowns and apparatus issues, we managed to obtain important data. Her comments undoubtedly helped me to improve my writing skills and my manuscript.

I would like to thank the jury of my defense, Bruno Alonso, Guillaume Bastiat, Nicolas Giraud, and Caroline Mello-Draznieks, for taking the time to read and evaluate my thesis. Their comments and discussions are helpful for my research and give me the pleasure of being a writer with readers.

In addition, I want to thank Nicolas Giraud for being my comité de suivi. He gave me assurance and advice at the right time when I felt doubt about myself and this whole PhD study.

It was an opportunity and a challenge to work in three laboratories in three different cities. I would like to thank Emmanual Magnier, the director of ILV, Thomas Pino and Bernard Bourguignon, the directors of ISMO, and Laurent Michot, the director of PHENIX, for their welcome to the lab. Without their help, I could not conduct research in proper condition. I would like to thank the administration teams who facilitated my integration and research in the three laboratories: Lise Michot, Sandra Jelbi at ILV, Nathalie Baudry, Sylvie Chabot, Madhan at ISMO, Gérard Guillard at PHENIX. I also thank Anne-Sophie Soares Capela at UVSQ for helping me to deal with the contract and all HR-related administration and Suzanne Thuron for assisting me with doctoral school-related administration.

I would like to acknowledge the funding from the RESPORE network. It allowed me to do my PhD study and the 3-month extension critical for writing my thesis and completing my research projects.

Academia has the reputation of "lone ranger" projects where a PhD student is a project manager responsible for all decisions. It is partly true, but I can never finish all the work alone. I am lucky to collaborate with several dedicated researchers who contributed significantly to my PhD.

Dr. Didier Desmaële, it is my pleasure to collaborate with you. Your broad chemistry knowledge and meticulous attitude set a high standard for me.

Dr. Marianna Porcino, thank you for sharing advice about NMR and doing the experiments in Orleans. The obtained data are key to my very first article as the first author.

Dr. Benedikt Lassalle-Kaiser, thank you for agreeing to my proposal to do the XANES study. It was the first time that I successfully found a research collaboration. Thanks to you, I enjoyed the beam time in Synchrotron SOLEIL and learning about XANES.

Dr. Mathieu Frégnaux, thank you for proposing the XPS study. The XPS data have brought added value to my projects. I was delighted to explore a new technique in the final year of my PhD.

Thanks to my teams and colleagues in ILV, ISMO, and PHENIX, my PhD life has never had a dull day. I rarely failed to find someone with whom to have coffee, take a break, and have both random and deep talks. The discussions with other PhD students and researchers in these three laboratories made my project better.

To my lab mates and teammates in ILV: Ali, Soumaya, Maxence, Jonathan, Subha, Juan, Dmitrii, Khaled, Sanchari, Sa, Cédric, Sergiu, Nour.

To my lab mates and teammates in ISMO: Ioanna, Seray, Tom, Mengli, Wenbo, Jesus, Killian, Lisa, Jingwen, Céline, Mario.

Thank you for sharing your moments and journeys with me! All the best for your future! Ali, our minds constantly clicked when we first met. I do not remember how it happened, but since then, we have been able to talk about lots of things for many hours. I thank Yuliia, the intern student. She stayed for only six months, but her work had a lasting impact on this PhD thesis. I appreciate her autonomy, and I got to practice teaching and training while working with her.

Arcadie and Cendra were the few people who shared the office with me in ILV and asked me to go to lunch. Fay was the first to realize that I was isolating myself in my office. Thank you for talking me into going to the Lavoisier building to see people! I want to thank Flavien Bourdreux for measuring 100x, if not 1000x, NMR tubes with the autosampler. Thanks to him, I was able to perform the quantitative NMR study critical for my research.

I want to thank Emmanuel Cadot, the head of the MIM group. He was determined to find a way for me to stay working in ILV, and he continued paying attention to my work during my time there.

Clément Falaise was also a supporter of my work. He often reassured me, "I am sure you will do a great job!".

Ioanna, you are a special person to me because we continuously worked together, first at ISMO, where you did your PhD, and then at ILV, where you did a post-doc. Since we first met at ISMO when you supervised me during my internship, I have admired your spirit. You work hard and play hard. Thank you again for showing me an example of a chemist by day and a lady by night! Seray, the second lady whom I am a fan of, I enjoyed working and discussing with you. We also made a great team. It was always nice to hang out and party with you! Charles, I could not ask for a nicer and funnier office mate. It was so easy to speak French to you! I would like to thank Farah Savina for helping me with all the issues related to safety and instruments in ISMO. My accident with acid in ISMO was quickly resolved with her help.

To the PhD seminar team in ISMO, Jean-Nico, Samrit, and Jonathan, thank you for being part of the organizing team! The time I spent organizing PhD seminar gave me more energy to do my main job.

Although I spent only a little time at PHENIX, I enjoyed the lunch, the bar, and the chat with the students and researchers. I like the friendly atmosphere and the activities the team creates to bring everyone together.

I also want to thank Pierre Levitz and Guillaume Mériguet for helping me with the data analysis and interpretation for SAXS and relaxometry, notably Pierre, for showing me the beauty behind the SAXS curve.

I am bad at keeping connections, so I thank my friends for not forgetting me and keeping me sane through the dark times. To you, all my dear friends, we do not talk regularly, but we will always be available for each other.

To my Vietnamese friends in Paris, thank you for being the reason that makes me get out of bed and go out.

Chi, for all the gossip and cheering on all frontiers, from work to family to love. Thank you for traveling all the way to Paris to support and help me with the defense! Thúy, for listening to me and helping me no matter what. I am blessed to have a friend like you.

Kinanti, my classmate in SERP+ Master, helped me to analyze SAXS data and navigate through the Igor software.

A special acknowledgment goes to my boyfriend, Thành, also my best friend, chef, travel buddy, study buddy, collaborator, private consultant, chess teacher, and more.

Thank you for making me sleep more (I admit that I always underestimate how long I need to sleep)! He is always present for me and surrounds me with warmth, (dry) humor, care, and love. I cannot describe how much he has done for me, especially while I was writing my thesis and he was writing his. Without him, I would not have been able to put that much effort into writing.

You will very soon be Dr.! I cannot wait to witness the bright future ahead of you and what you can do. Look around, and you shall see many beautiful and good things (in addition to cats on the roads).

We made it together! I want to say my final words to my family. You are far, but you are always in my heart. For all the years I lived abroad, I missed all the chances to see my two sisters grow up and help my parents when they needed me. I always owe the love, support, and pride of my family. They are the only ones who care about me unconditionally.

What a journey a PhD thesis is! It has been unpredictably exciting and scary at the same time. I cherish everything I have learned along the way and every person accompanying me.

See you on the next adventures! CHEERS!

2023-01-02 3.3.3 Chemical environment and atomic composition at nanoMOF surface ......................... 3.3.4 Discussion on the ligand exchange process............................................................................. 3.3.5 Evolution of metal coordination sphere in nanoMIL-100 (Al) ............................................ 3.3.6 Evolution of metal coordination sphere in nanoMIL-100 (Fe) ............................................ 3.3.7 Discussion on the degradation mechanism of nanoMIL-100 in phosphate and possible degradation products ................................................................................................................ 3.4 Degradation and drug release mechanisms of modified nanoMIL-100(Al) and (Fe) ..... 3.4.1 Optimization of drug loading formulation for ATP/AMP drug and nanoMIL-100(Fe)/(Al) carriers .................................................................................................................................... 3.4.2 Drug-carrier 3.5.3 Effect of drug loading and surface coating on morphological evolution of degraded MIL-100 nanoparticles ............................................................................................................................... 3.5.4 Interaction between nanoMIL-100 and water in the nanoparticle suspension ........... 3.5.5 Behavior of water in degraded nanoMIL-100(Al) and (Fe) ................................................. 3.6 3.8.1 Chemicals ............................................................................................................................................ 3.8.2 Synthesis of nanoMIL-100 ............................................................................................................. 3.8.3 Drug encapsulation.......................................................................................................................... 3.8.4

Synthèse en français

L'administration de médicament a pour but de formuler et administrer des médicaments visent à obtenir la meilleure performance et efficacité en toute sécurité. Les technologies d'administration contrôlée de médicament ont révolutionné le traitement de nombreuses maladies (décrit au chapitre 1.1.). Un système d'administration de médicaments ("drug delivery system", DDS) agit comme une interface entre les médicaments et l'environnement biologique. Un véhicule doit transporter et délivrer les ingrédients actives avec un contrôle à la fois dans l'espace (site malade) et dans le temps (durée de traitement). Le médicament, ou substance/ingrédient actif, peut être incorporé dans véhicule par des interactions physiques ou chimiques. La surface de véhicule peut être modifiée pour assurer une meilleure stabilité du système et faciliter le transport dans l'environnement biologique. Un DDS doit être optimisé et adapté au type spécifique de maladie et à l'effet thérapeutique recherché.

Les DDSs à l'échelle nanométrique ont la capacité de pallier les inconvénients liés aux propriétés physicochimiques indésirables du médicament et de contourner les barrières biologiques, ce qui permet un contrôle complet du processus d'administration du médicament. Néanmoins, la dégradation des nanomatériaux dans les milieux biologiques reste une question brûlante qui entrave l'application clinique des nanoparticules (NPs) comme DDDs. La connaissance de la dégradation des nanoparticules est donc critique pour contrôler la libération de médicaments et limiter la toxicité des NPs et de leurs produits de dégradation.

Dans cette thèse, nous sommes curieux sur les réactions des NPs dans le milieu mimétique biologique. Les NPs sont incubées dans un tampon phosphate salin (PBS). Après la réaction, les propriétés des NPs, telles que la morphologie, la stabilité, la structure chimique, la composition, sont évaluées par diverses techniques de caractérisation. Nous avons un intérêt particulier pour la spectroscopie par résonance magnétique nucléaire (RMN). La RMN est une technique flexible qui peut fournir des informations structurales au niveau moléculaire et des analyses quantitatives sur les solides, les liquides et les gaz. La RMN a été utilisée ici pour déterminer la structure de DDS, quantifier la libération du médicament et suivre la dégradation du véhicule.

Nous étudions ici le processus de dégradation de deux types de nanovecteurs à base de polymères et de réseaux métallo-organiques ("metal-organic framework" ou MOF). Les NPs polymères (revues au chapitre 1.2.2.) sont des matériaux organiques, basés sur l'assemblage de centaines aux millions d'unités monomères identiques reliées par des liaisons covalentes. Les NPs de MOF, ou nanoMOFs (revues au chapitre 1.2.3.) sont des matériaux hybrides inorganiques-organiques formant un réseau de coordination entre des métaux et des ligands organiques avec des cages volumineuses. Alors que les polymères sont bien connus pour leur caractéristique de nanovecteurs, les MOFs restent des nouveaux nanovecteurs moins étudiés.

Tout d'abord, nous discutons au chapitre 2 d'une nouvelle approche pour préparer des NPs conjuguées polymère-médicament et étudier leur dégradation et leur libération de médicament. L'objectif ultime est de développer un nouveau nanovecteur pour lutter contre la tuberculose, une maladie infectieuse très mortelle. Les traitements actuels sont longs et compliqués à mettre en oeuvre, surtout pour des patients dans les pays en développement qui n'ont pas l'accès à l'assistance médical. Nous voulons utiliser l'acide itaconique (ITA), qui a une activité antibactérienne potentielle. Pour administrer cette nouvelle molécule active, nous utilisons les nanovecteurs bien connue à base de poly(acide lactique), ou PLA, dont la biocompatibilité est largement reconnu. Une méthode chimique doit être établie pour incorporer l'ITA hydrophile dans le PLA hydrophobe afin d'éviter la libération prématurée de l'ITA. Cette méthode consiste à synthétiser dans un premier temps un polymèremédicament conjugué à base d'acide lactique et d'acide itaconique, puis à le formuler en NPs. Les NPs doivent être dégradables et libérer l'ITA, donc le processus de dégradation dans le PBS est un focus de recherche.

Dans notre nouvelle approche, le polymère conjugué PLA-ITA a été synthétisé par la polymérisation du PLA dans deux directions, suivie par la condensation de ITA dans les deux groups terminaux du polymère. Le polymère a été caractérisé par les spectroscopies RMN et IR, et par spectrométrie de masse. Le polymère PLA-ITA a ensuite été formulé en NPs en utilisant la méthode de nanoprécipitation. Après, nous avons utilisé la RMN liquide et la spectrométrie de masse (ESI et MALDI-TOF) pour identifier la structure moléculaire des produits de dégradation. Nous avons trouvé qu'ils contiennent divers monomères et oligomères mixtes d'acide lactique et d'acide itaconique. La RMN quantitative a montré que le processus de dégradation de PLA-ITA est plus favorable à pH 7.4 qu'à pH 5.3. Il est probable que l'hydrolyse des groupes esters dans le PLA-ITA suit un mécanisme catalysé par une base, où les liaisons ester sont coupées de manière aléatoire. La lente dégradation explique la lente libération de molécules d'ITA. Aussi, la taille et la morphologie des NPs ont été caractérisées par microscopie électronique MET. Les résultats suggèrent que l'érosion des NPs de PLA-ITA se produit de l'intérieur sans changement de leur morphologie et taille.

En résumé, nous avons présenté une méthode de préparation de polymère conjugué pour encapsuler des molécules hydrophiles qui ne conviennent pas au chargement physique de médicament. La RMN liquide offre une méthode pour vérifier la libération avant les évaluations biologiques. La vitesse de libération de ITA peut être amélioré par utiliser des autres polymères qui se dégradent plus rapidement ou par fabriquer des nanoparticules plus petites.

Dans le chapitre 3, nous avons mené une étude approfondie sur la dégradation de deux nanoMOFs de type nanoMIL-100(Al) et nanoMIL-100(Fe) dans le PBS. Les nanoMIL-100 sont constitués de cations métalliques trivalents, e.g., Al(III) et Fe(III) et de ligands tricarboxylates de benzène, i.e., trimesates. Chaque centre métallique possède 6 liaisons de coordination avec les trimesates et l'eau. Les cages sont assez grandes pour encapsuler des molécules actives volumineuses. Dans la littérature, le nanoMIL-100(Fe) est le nanoMOF le plus étudié pour l'administration de médicaments pour sa grand capacité d'encapsulation une variété de médicaments. Le nanoMIL-100(Al), l'analogue structural du nanoMIL-100(Fe), a également des capacités prometteuses pour encapsuler et libérer des médicaments. Les deux nanoMIL-100(Al) et nanoMIL-100(Fe) se caractérisent par une forte affinité vis-à-vis des groupements phosphates, que l'on peut retrouver dans les médicaments ou les molécules de revêtements de surface. Il existe des recherches sur la dégradation de nanoMIL-100(Fe) dans les milieux phosphates. La question de recherche qui nous intéresse ici est de savoir si les deux nanoMIL-100 ayant des structures similaires mais des métaux différents, présentaient les mêmes mécanismes de dégradation.

En utilisant une analyse multi-échelle, le mécanisme de dégradation par des phosphates a été dévoilé. Les phosphates remplacent les deux ligands natifs des nanoMOFs, suivis par la formation de phosphate d'aluminium/fer et la libération des ligands trimesates. Le rôle du phosphate a été mis en évidence par la RMN à l'état liquide et solide, la spectrométrie photoélectronique X (XPS), et la spectroscopie d'absorption au niveau du seuil X (XANES). La RMN liquide détecte le trimesate libre en solution et suit la cinétique de formation du trimesate libre. Le XPS détecte la perte du ligand trimesate dans le MIL-100 et l'incorporation du phosphate dans le réseau solide MIL-100. Les données quantitatives de RMN liquide et de XPS sont corrélées, mettant en évidence l'échange de ligands entre les trimesates sortants et les phosphates entrants. Plus le rapport molaire P/M (phosphate/métal) est élevé, plus la quantité de trimesate libérée est importante. Ensuite, la formation de phosphate d'aluminium et de fer dans MIL-100 dégradé est démontré par RMN à l'état solide et XANES respectivement. Cependant, nous avons constaté que le nanoMIL-100(Al) est plus réactif vis-à-vis de phosphate que son analogue de Fe car le Fe(III) a une affinité plus grande que l'aluminium envers les phosphates et les carboxylates.

En outre, la taille, la morphologie et la rugosité de surface des nanoparticules ont été suivies pendant la dégradation par diffusion des rayons X aux petits angles (SAXS) et par microscopie électronique (MET et MEB), ce qui a permis d'obtenir des informations sur les processus d'érosion différents des MOFs à l'échelle nanoscopique. D'une part, les NPs MIL-100(Fe) semblent réagir comme des particules individuelles et conservent leur taille et leur morphologie globales. A partir d'un certain rapport P/Fe qui provoque un haut degré de dégradation, le nanoMIL-100(Fe) dégradé subit une rugosité de surface progressive due à la formation de phosphate de fer. D'autre part, les NPs MIL-100(Al) sont transformées aux agrégats des petites particules de phosphate d'aluminium.

Pour sonder les mouvements moléculaires de l'eau et notamment autour des surfaces, la relaxométrie RMN a été utilisé. Cette dernière renseigne sur les changements de l'état de surface interne et externe des nanoMOFs. La spectroscopie RMN à basse température nous a permis de distinguer l'eau confinée dans les cages du MOF ou près de la surface interne de l'eau à l'extérieur du MOF. La relaxation RMN de l'eau dans le nanoMIL-100(Fe) est fortement affectée par le paramagnétisme du fer(III). L'étude de la relaxométrie par RMN de MOFs dégradés suggère une mobilité réduite de l'eau à l'intérieur des pores du MOF et une accessibilité altérée aux sites métalliques de surface.

De plus, la dégradation des nanoMOFs chargés et recouverts de molécules phosphatées s'est avérée dépendre de la force de la complexation métal-phosphate. Le triphosphate ATP interagit fortement avec nanoMIL-100, beaucoup plus que son dérivé équivalent monophosphate AMP. Cette interaction entre en compétition avec les ions phosphates du milieu PBS, altérant le processus de dégradation. Le revêtement de surface avec des cyclodextrines phosphatées (CD-P) améliore la stabilité colloïdale du nanoMOF dans l'eau mais n'affecte pas sa dégradation, probablement en raison d'une revêtement de surface partiel.

En sommaire, les deux nanoMOFs analogues nanoMIL-100(Al) et nanoMIL-100(Fe) ont montré le même mécanisme de dégradation à l'échelle moléculaire, mais les évolutions différentes de leurs particules à l'échelle nanoscopique. Ces nouvellles connaissances contribuent à une meilleure compréhension de la chimie des MOFs et peuvent aider à l'optimisation des nanoMOFs en tant que DDSs. L'approche multi échelle peut être appliqué à d'autres études sur des MOFs.

Enfin, le chapitre 4 est consacré à la mise en oeuvre d'une méthode de spectroscopie RMN in situ pour suivre la dégradation des nanoMIL-100(Al) sans séparer les particules de la suspension. La méthode consiste à exploiter simultanément les signaux RMN du solide et du liquide, bien que leurs réponses soient intrinsèquement différentes en raison du degré de liberté contrasté dans le solide et liquide. La rotation à l'angle magique (MAS) est nécessaire afin d'obtenir une résolution des signaux RMN de la phase solide suffisante pour les différentier de ceux de la liquide. En utilisant des inserts scellés et une séquentiel configuration de RMN de 1 H puis 27 Al, des suspension ont pu être mesurées en conditions MAS permettant un suivi cinétique simultané de libération de trimesate en phase liquide et de la perte du trimesate en phase solide. La coordination de l'aluminium avec le phosphate a également été observée, renforçant le mécanisme de dégradation proposé. La méthode a été ensuite appliquée au nanoMIL-100(Fe) paramagnétique et au nanoMIL-100(Al) chargé de médicament. Pour le nanoMIL-100(Fe), seulement le trimesate formé en liquide sont détectés. Pour le nanoMIL-100(Al) chargé de médicament, l'influence de l'ATP et l'AMP sur la dégradation du nanoMOF et la libération des espèces actives est comparée.

La méthode de RMN in situ présentée ici permet la caractérisation précise des processus dynamiques des NPs dans leur environnement naturel sans séparation de phases. Il convient de tenter d'appliquer les méthodes de RMN in situ à d'autres processus liés à l'administration de médicaments, tels que le chargement de médicaments et le revêtement de surface. Différentes molécules actives et différents noyaux RMN comme le 19 F et le 31 P peuvent être envisagés.

En conclusions, au cours de cette thèse, nous avons appris les différents processus de dégradation de deux types différents de nanovecteurs de médicaments. Le polymère conjuguée PLA-ITA composé d'un enchainement de liaisons ester se dégrade par l'hydrolyse. Les nanoMOFs MIL-100(Al) et MIL-100(Fe) constitués de réseaux de liaisons de coordination métal-carboxylate réagissent avec les anions phosphates. La panoplie des méthodes analytiques employées a permis une description fine des processus mis en jeu lors de dégradation des nanovecteurs et d'élucider leurs mécanismes.

Chaque DDS composé de matériaux uniques nécessite sa propre analyse sur le mécanisme de dégradation et de libération du médicament. Nous espérons que les approches et méthodologiques développées dans cette thèse pourront aider les chercheurs à la conception et la caractérisation de nanovecteurs ou à l'utilisation de la spectroscopie RMN pour comprendre la chimie de nanovecteurs.

General introduction

Drug delivery has been one of the pillars of pharmaceutical technology, to administer the drug to patients in the best state and quality. Controlled drug delivery technology, seeded from the research in the 1960s 1,2 , has sought to regulate the time and location of drug delivery by designing a controlled drug delivery system (DDS), an interface between the drug and biological environment 3,4 . An ideal DDS carries the drug through biological barriers and protects its activity until the DDS arrives at the target location and releases the drug in the desired duration. The drug, or the active substance, can be incorporated with the carrier by physical or chemical methods. The carrier's surface can be modified to improve the system's stability and regulate the impact of the external environment on the carrier. Furthermore, the design of a DDS involves understanding the mechanisms of drug delivery processes, such as drug release and carrier degradation, as well as the underlying parameters behind a successful (or failed) formulation of drug loading and surface modification.

In the diverse library of potential materials for drug carriers, nanomaterials have prompted attention from the research community owing to their nanosize-dependent physiochemical properties, potential ability to bypass biological barriers, and versatility to build a multifunctional platform for disease therapy and diagnostic 5,6 . Nevertheless, the clinical translation of nanoparticles (NPs) for drug delivery applications has not matched their expectation and effort spent [7][8][9][10] . One of the reasons is the incomplete understanding of the interaction between NPs and biological environments, thereby hampering the ability to control their drug delivery performance and raising concerns about their toxicity.

In this context, we want to contribute to address these issues by studying the chemical reactions of nano-scaled DDS with biological mimicking media, especially during the carrier degradation and drug release processes. That information is essential to understand the relationship between the nanoparticle's design and its drug delivery performance so that one can optimize the suitable parameters. In-depth characterization of structure, physiochemical properties, and reactivity of degradation products should be done. To this end, nuclear magnetic resonance (NMR) spectroscopy is a promising tool that affords both qualitative and quantitative analysis at the atomic level. Liquid-state and solid-state NMR spectroscopy have often been employed to determine the DDS structure, follow the drug release, and quantify the released species [11][12][13] . Also, NMR experiments can be performed in situ, which is suitable for studying dynamic processes. Therefore, it will be beneficial to explore and extend the capabilities of NMR spectroscopy for the characterization of DDSs.

Capitalizing on the expertise in the field of drug delivery and formulation as well as material characterization and NMR spectroscopy in the host laboratories, this PhD project aims to investigate the degradation mechanisms of some nanoparticle-based DDSs, and to exploit different uses of NMR spectroscopy to do that. How does the degradation happen? What are the degradation products? How fast is the process? Do drug loading and surface modification play any role in carrier degradation? Those questions motivate the experimental work and the direction of this thesis, in particular the choice of two case studies of NPs based on polymer and metal-organic framework and the attempt to develop in situ NMR spectroscopy method for NP suspension.

The present thesis has four main chapters, as summarized in the scheme below.

Chapter 1 gives the scientific background and literature on the application of controlled drug delivery, the carrier's materials polymer and metal-organic framework nanoparticles, and the characterization techniques for DDSs with attention on NMR spectroscopy.

The following two chapters are about the degradation mechanisms of two distinctive nanoscale DDSs, made of organic and hybrid inorganic-organic material, the polymer and metal-organic framework NPs, respectively. While polymer has a long history in controlled drug delivery, metal-organic framework is a relatively new material for drug carriers.

In chapter 2, we study the degradation process of polymer-drug conjugate NPs made of itaconic acid (ITA), an active molecule, and lactic acid (LAC), a polymer monomer. The poly(lactic acid)-itaconate, PLA-ITA, was synthesized by a new approach and formulated as NPs. The potential of a pH-dependent release of ITA through the degradation of PLA-ITA was investigated. The nature and quantity of PLA-ITA NPs degradation products at pH 7.4 and 5.3 were determined using mainly liquid NMR spectroscopy. Their chemical structures were proposed. The degradation behavior of PLA-ITA NPs was rationalized by ester hydrolysis and bulk erosion mechanism. This chapter underlines the role of studying polymer degradation in the design of polymer-drug conjugate DDSs.

Chapter 3 is dedicated to the degradation of metal-organic framework NPs, shortly nanoMOFs, made of MIL-100(Al) and MIL-100(Fe). The nanoMIL-100 (MIL stands for Matériaux de l'Institut Lavoisier) is a MOF family famous for its drug delivery application. The degradation of nanoMIL-100 in phosphate-containing media was thoroughly examined at both molecular and nanoscopic scales. Also, the impact of metallic MOF's components on the drug delivery chemistry was studied through the two aluminum-and iron-based structural MIL-100 analogs. By combining multiple characterization techniques, mechanisms for degradation reaction with phosphate and erosion of nanoMIL-100 in stimulated physiological medium, with or without encapsulated drug and surface coating, were proposed. The phosphate-driven degradation phenomenon was evidenced by quantitative NMR spectroscopy and semi-quantitative XPS. The identity of degraded MIL-100 was determined by solid-state NMR and X-ray absorption near edge structure (XANES) spectroscopy. Complementary information about the NPs' morphological evolution was provided by small-angle X-ray scattering (SAXS), electronic microscopy (TEM and SEM), and NMR relaxometry. The findings of chapter 3 may help further studies on nanoMOFs' behavior in more complex media, which is important for making biodegradable nanoMOFs for drug delivery applications.

Chapter 4 describes the development of an in situ NMR spectroscopy method to study reactions of NPs suspension as they are without phase separation. Using nanoMIL-100(Al) and nanoMIL-100(Fe) as examples, we set up an NMR method to simultaneously detect degradation species in both liquid and solid phases. It offers both qualitative and quantitative information about the nanoMIL-100 degradation empty and loaded with drugs, in agreement with the results in chapter 3. The presented in situ NMR technique can become an alternative time-and labor-saving way to study the kinetics of heterogeneous reactions.

Introduction -Literature review

This chapter introduces the context and the scientific literature on the topic of this thesis, a combination of the drug delivery applications, the carrier materials, and the associated characterization techniques. It is not meant to be an exhaustive review but a view of where our work is in the big picture of the field. There are three main parts:

1) Overview of drug delivery, its development status, and the drug delivery processes that are dictated by the physical chemistry of drug delivery systems 2) State-of-the-art of nanocarriers and two specific types that were studied in this thesis: polymer and metal-organic framework nanoparticles 3) Generalities about the characterization techniques for drug delivery systems, with attention on NMR spectroscopy and in situ techniques 1.1.1 Basics of drug delivery ...................................................................................................................... 1.1.2 Physical chemistry of drug delivery .............................................................................................. 1.2 ............................................................. 1.2.2 Polymer nanoparticles ...................................................................................................................... 1.2.3 Metal-organic framework nanoparticles .................................................................................... 1.3

Drug delivery

Medicine and drug have always been a significant part of human history and are indispensable in modern human lives. Drugs cure diseases that used to be death sentences in the past. Among other drugs of historical significance 1 , penicillin, discovered by Fleming in 1928 and tested by Florey and Chain in the 1940s 2,3 , is the first antibiotic for bacterial infection. Another example is insulin, discovered and purified by Banting, Best, Collip, and Macleod in the 1920s 4 is the first hormone therapy for diabetes.

Nevertheless, a drug is not inherently effective. Its therapeutic benefit depends on how it enters and acts inside the human body. For instance, the drug can lose its activity in the acidic stomach in the body. Those behaviors are dictated by drug delivery method, how and in what form the drug is introduced into the body 5 . Simply, drug delivery "makes drug work better" 6 .

Basics of drug delivery

Drug delivery has always accompanied drugs throughout history. Anywhere, anytime, where there is a drug, there is drug delivery. In ancient times, the medicines in China were formulations of herbal plants, and the pill was described in Egypt in 1500 BC 7,8 . Fast forward to the 19 th century, tablet and capsule were invented in France 7,9 .

The scientific approach to study medicine produces systematic knowledge and terminology for drug and drug delivery. Pharmaceutics, medicinal/pharmaceutical drug, or shortly drug in this thesis, are some of the names to indicate drugs used to manage, treat, or prevent disease. Active pharmaceutical ingredient (API), also called active substance, is a substance or a mixture of substances intended to be used in drug manufacture, which then becomes the active ingredient that exerts pharmacological activity on the body 10,11 . The API is essential in the composition of any drug. Excipient, basically anything other than the active ingredient in the drug product, is the inactive substance that does not have biological activity and is used to facilitate the manufacturing and handling of drugs 12 . Several examples are flavoring agents that mask the unpleasant taste and odor, fillers that bulk up the drug that contains a small amount of API, and preservatives that prevent drug decomposition during storage 12 . Dosage form is the physical state of the drug comprised of a specific mixture of API(s) and excipient(s) 12,13 . It can be solid (table, capsule, powder), semisolid (cream, paste), or liquid (syrup, suspension) 12 . A given drug produced by two companies can have different dosage forms. Formulation is the process of preparing, combining, and shaping the API and excipient to produce the final drug product. Formulation also often refers to dosage form 14 . Route of administration is the path by which the drug is introduced/administered to the body. It indicates the location where the drug enters the body, such as "oral": to mouth, "intravenous" into vein, and "nasal": through nose 15 .

There are several terms that are related to a drug's therapeutic effect/pharmacological activity/biological activity. Potency is the amount of drug needed to produce an effect of a given intensity, quantified by EC50, the drug concentration required to produce 50% of its maximum effect 9 . Efficacy is the capacity of a drug to produce its maximum effect 13 . Therapeutic window (TW) is the drug concentration range within which the drug produces Chapter 1 Introduction -Literature review a wanted effect without causing toxic effects 9,16 . It lies between minimum effective concentration and maximum safe concentration. Therapeutic index (TI) is the ratio between the median lethal dose (LD50) 13 and median effective dose (ED50) 13 . The TI reflects approximately the safety of a drug 13 . A low TI indicates the close values of effective and toxic doses. A drug with higher TI is considered safer to use than one with a lower index.

As new pharmaceuticals were developed, from small molecules to macromolecules and live cells, drug delivery methods have also evolved to adapt. Three main sets of strategies are drug modification, microenvironment modification, and drug delivery systems 17 . Drug modification method alters the API's structure by adding functional groups, making analogs and derivatives, and conjugating to targeting ligands. Environment modification involves changing the biological environment to which the drug is exposed to aid its navigation toward the biological target. For example, excipients are added to the drug formulation to increase API solubility or enhance cell uptake by altering the local pH. A drug delivery system (DDS), virtually a combination of the two mentioned strategies, acts as an interface between the drug and its microenvironment. DDS can be "a formulation or a device" 16 . In the scope of this thesis, DDS is a carrier/vehicle that ideally can (Figure 1-1): 1) transport drugs to the desired site of action, 2) protect drugs against degradation during transit, 3) release drugs at the intended site at a defined rate for a defined duration, 4) be removed from the body without causing toxicity or undesired effect 5,18,19 . A DDS should be optimized for the specific drug and treatment to maximize therapeutic outcomes and ensure safety and patient compliance 20 . Several examples of DDS are illustrated in Figure 12. Early DDSs are immediate-release formulations that release the active substance immediately after administration, resulting in a fast onset of action 20 . In the 1800s, they were used for oral tablets, which release API in the stomach upon contact with water 21 . There was no control over the drug release process and its rate. Later, delayed-release formulations used enteric coatings insoluble in the acidic stomach (enteric = relating to/occurring in the stomach), which prevents drug release in the stomach. However, the coating is soluble in the slightly alkaline intestine, thus leading to drug release there 21 . The onset of action was delayed, but still, the active ingredient was released immediately after the coating removal.

Those formulations are still being used today and are widely referred to as conventional DDS.

Nevertheless, conventional DDSs are not suitable for all active substances. A widely mentioned issue is the unsustainable therapeutic effect because the drug is quickly released and eliminated from the body 9,18,21 . Such a phenomenon is illustrated in Figure 1-3. When taking a single dose, the drug's blood level may not be maintained for a long enough period to produce the desired effect. In comparison, taking multiple doses results in fluctuating drug concentration in the blood. The "valley" concentration leads to no effects, while the "peak" concentration is close to or above the safe level. In addition, treatment using multiple doses per day is inconvenient for the patient to follow. If one forgets to take medicine, the treatment can fail. Here comes the idea of controlled release DDS, designed to control the drug release rate and maintain the drug blood level over an extended time. Broadly defined, controlled DDS is a formulation or device that can provide control over the rate, time, and/or place of drug release 5 . Advantages of controlled DDS compared to conventional DDS include: i) continuous maintenance of drug blood concentration leading to sustained therapeutic effect, ii) reduction of side effects because of targeted delivery and/or sustained delivery rate, and iii) possible delivery of drugs with short half-lives and hence short duration of action. A typical case needing controlled DDS is a cancer chemotherapy drug, which is toxic to both sick and healthy cells, not only kills cancerous cells but also damages the patient's health causing serious side effects like tiredness, hair loss, and infection 23 . These undesired effects can be reduced if the drug can be delivered selectively to the tumor.

Historical perspectives on 70 years since the first commercial controlled drug delivery formulation point out three generations of controlled DDS [24][25][26] : the first generation of the "MACRO" era, the second generation of the "MICRO" era, and the third generation of the "NANO" era. The dawn of controlled DDS witnessed the development of macro-sized oral and transdermal formulations. The very first example is the Spansule® technology, introduced in 1952 by Smith, Kline, and French Laboratories allowed sustained delivery for 12 hours of a cold remed 21 . The technology was based on controlling the dissolution in the gastrointestinal fluid of API coated with natural waxes. Afterward, a key publication in 1964 by Folkman and Long demonstrated proof of concept for a constant-rate drug delivery device made of silicon tubing 27 . This work inspired the development of diffusion-controlled systems whose drug release rate can be determined by the drug diffusion process through a membrane. Examples include the Ocusert® insert that releases the drug in the eye for one week or the Transderm Scop® skin patch for relieving motion side sickness 24 . The second generation's representative DDS is a long-acting injectable formulation that delivers drugs for weeks to years, made of biodegradable polymer microparticles such as polyester, polyanhydride, and their copolymers. For example, Lupron Depot®, approved by the U.S. Food and Drug Administration (FDA) in 1989, delivers leuprolide, a protein, for up to 6 months. Injections are required every 1 to 6 months, which dramatically surpasses the past daily injection of leuprolide alone 28 . It is also one of the first commercial products of protein encapsulated in polymeric DDS, stemming from the academic research work by Langer 29 . The third generation of controlled DDS is dominantly nano-sized DDS with massive attention to cancer treatment. Several original concepts were conceived in 1960-1980s 24 , such as drug-polymer conjugate, PEGylation, i.e., conjugation of polyethylene glycol (PEG) to drugs to increase its circulation time and stability, and targeting using targeting ligands like antibodies or enhanced permeability and retention (EPR) effect, the phenomenon of nanosized objects accumulating in tumor tissue through leaky vessels. The development of diverse types of nanoparticles and nanostructures has been flourishing, from organic materials like polymer, lipid, etc. to inorganic materials like gold, silver, platinum, silica, etc. Famous DDSs of this third generation are Doxil®, a breast cancer drug encapsulated in liposomes, or Onpattro® which uses lipid nanoparticles to deliver siRNA, a gene therapy for a fatal genetic disease. Overall, drug delivery technologies throughout the years have enabled both reformulation of old drugs and use of new drugs, as well as become integrated into drug discovery process 17,30 . A must-be-mentioned example is the covid-19 vaccine Comirnaty® of BioNTech and Pfizer, a lipid nanoparticle-based formulation of mRNA 31,32 , capitalized on research about lipid nanoparticles and gene delivery in the previous decade.

Despite those successful commercial DDS, there are still unsolved challenges of drug delivery, including tumor targeting and delivery of macromolecular drugs 25 . It sounds frustrating that numerous nanoparticle formulations of targeted drug delivery that were successful in vivo have not been translated to clinics 33,34 . The EPR effect, a basic rationale used by many studies on nanoparticles, is also being questioned 35,36 . Regarding macromolecular drugs like protein, peptide, and nucleic acid, they have very short half-lives and are too big to be absorbed from the digestive tract, thus being mainly administered by injection. Current attempts to replace invasive injection involve using other routes like skin or nose and developing long-acting formulations that can deliver for weeks and months. These challenges require DDS that can overcome both the physiochemical properties of drugs and biological barriers in the human body. Moreover, there is a heated debate about the field's development progress and approaches to solving the current challenges. In science, a high number of papers and citations does not necessarily lead to the invention of new ideas 37 . As there are too many publications to follow, it is difficult to select good references and make discoveries. The same goes for drug delivery; extensive literature on DDS in the last two decades does not produce more breakthrough 38 . It is argued that research has been focused too much, blindly sometimes, on nanotechnology-based DDS 16,39,40 . The ability of targeted nanocarriers 25,41 , and generally, the extent of progress contributed by nanomedicine is debatable 42 . Both the sides of pro 43,44 and against 16,45,46 nano have made their points. Despite their different opinions, the common ground between the two sides is that more quantitative and mechanistic studies on how nanomaterials work inside the body are needed 39,41,47 . It is also necessary to diversify research directions from cancer and nanoparticles and avoid creating complex yet not practical DDSs 48 . Priorities can be put on investigating treatment for the diseases that are leading causes of death, e.g., cancer, heart diseases 49 , or have a high social impact, e.g., mental-disorder [START_REF] Cierkowska | Building pharma pipelines using a socioeconomic lens[END_REF] . Precision medicine tailored for each patient and low-cost, affordable healthcare for everyone is also likely to bring great clinical, commercial, and society impact 49 .

To summarize, the field of drug delivery and controlled release has gone a long way, from being conceived as an interesting but practically impossible concept 51,52 to becoming a critical part of the pharmaceutical industry, bringing both economic and social values. Solutions to tackle the next challenges for drug delivery technology need both applied work on specific diseases and fundamental research on mechanisms. The next section will describe the general processes of how drug delivery works, particularly the physiochemical processes behind drug delivery and drug release.

Physical chemistry of drug delivery

Understanding how drug delivery works and the relationship between DDS's structure, properties, and delivery performance is essential to create an optimal DDS. The mechanisms of actions of a DDS are related to two types of behaviors: i) those of the drug in its DDS and ii) those of the drug, carrier, and DDS as a whole (drug + carrier) in the body. The first type, the interactions between the drug and its carriers, along with their individual properties, strongly affect DDS's formulation (material preparation, drug loading) and the drug release profile. The second type, the processes experienced in the body by the drug, the carrier, and the DDS, dictate their individual and collective in vivo fates. They were categorized into absorption, distribution, metabolism, and elimination (ADME) processes 12,53 . Briefly, drug absorption refers to the way the drug/carrier moves from the administration site to the bloodstream. Oral DDS should help a poorly soluble drug permeate through the gastrointestinal tract, while intravenous administration involves no absorption step. Distribution is how the drug/carrier transports from blood to tissues and cells. It is distributed throughout the body via blood circulation and then needs to bypass the cellular membrane, which depends on tissue binding affinity 5 . Its accumulation in nontarget organs/tissues can cause material waste and potential side effects 54 . Metabolism is the transformation of a drug/carrier into an active or inactive form in the liver or other organs. The metabolism to inactive form is unfavored. Elimination is the clearance of the drug/carrier and the remaining side products from the body, mainly through the kidney and liver. Slow removal of carriers loaded with drugs and fast removal of carriers alone are needed.

In the scope of this literature review, we focus on how physiochemical properties of the drug and carrier influence the DDS's formulation, namely the choice of the carrier's material, drug incorporation, carrier's surface modification, and the DDS's activities, namely drug release, carrier degradation. Figure 1-4 illustrates such processes.

Choice of material

The material of a drug carrier must be biocompatible and adapted to drug properties and intended application.

Biocompatibility describes the degree of tolerance of the relation between a material and a given biological environment 55 . It should be considered a two-way relationship in which the material interacts with its surrounding environment without causing unacceptable harm to the body. Vice versa, the body and the biological response upon contact with the material should not interfere with the material's intended function 56 . The choice of a biomaterial can be intuitively based on the intrinsic toxicity of its chemical composition, which can be measured by LD50, the median lethal dose 57 . However, the choice is not that simple. Biocompatibility can also depend on the degradation products of the drug carrier formed by its reaction in the body 55 . It is also affected by the carrier's physiochemical characteristics, especially for nanocarriers 56 .

DDSs of various sizes can be used for different drug delivery applications 58 . Macro-size systems, bigger than 1 mm and visible under our eyes, are devices made for long-term delivery methods like implants or drug-eluting stents. Nano-size carrier, smaller than 1 μm, is suitable for intravenous formulations as they require objects whose dimension is smaller than the blood vessel's diameter 59 . The size can also affect the biodistribution of drug carriers; for instance, particles less than 5 nm are more rapidly cleared from the circulation than larger particles 60 . For nano-size and micro-size colloidal suspension, the uniform size distribution of particles is required.

No matter the size, the surface is where the DDS is in contact and interacts with the surrounding physiological environment 61 . Therefore, the surface properties significantly contribute to the physical and chemical stability of the DDS. Surface hydrophobicity and charge can impact the stability in water, especially for micro-and nano-sized DDS. Their colloidal stability can be predicted and adjusted through surface charge, according to the DLVO theory 62 . Moreover, exposure of the carrier to protein and cells, leading to their adsorption to the carrier's surface, can impede or change its designed functions, render the DDS inactive, or lead to fast clearance 60,61 . Consequently, surface modification of DDS is almost always essential to control its stability in physiological media (see 1.1.2.3 for common strategies).

Stability in water and physiological media is imperative to ensure the proper functioning of DDS and to avoid the formation of unwanted harmful degradation products. Sources of destabilization can come from temperature, pH, water, protein adsorption and binding, and enzyme metabolism 63,64 A vast library of chemical composition has accumulated over the years, from natural to synthetic, from organic to inorganic, and then hybrid and composite materials. No matter the selected material, the carrier's synthesis and preparation should use non-toxic reacting agents for biomedical applications. The carrier's material composition impacts its toxicity and stability and decides the available chemistry for subsequent steps of DDS formulation and function, including drug loading, surface modification, and drug release.

Drug loading

The drug loading capability of a DDS is evaluated by drug payload and encapsulation efficiency. The drug payload, or drug loading, is defined as the weight proportion (wt%) of drug per carrier. The higher the drug payload, the better, as a smaller amount of carrier needs to be administered, consequently reducing potential adverse effects. Additionally, encapsulation efficiency (EE%) is the ratio between the amount of loaded drug over the added in the formulation process. Higher EE% reduces material loss, which is important for the pharmaceutical industry as pharmaceuticals are expensive. The calculations of drug payload and EE% are as follows:

𝐷𝑟𝑢𝑔 𝑝𝑎𝑦𝑙𝑜𝑎𝑑 (𝑤𝑡%) = 𝑚𝑎𝑠𝑠 𝑜𝑓 𝑙𝑜𝑎𝑑𝑒𝑑 𝑑𝑟𝑢𝑔 𝑚𝑎𝑠𝑠 𝑜𝑓 𝑐𝑎𝑟𝑟𝑖𝑒𝑟 × 100 𝐸𝑛𝑐𝑎𝑝𝑠𝑢𝑙𝑎𝑡𝑖𝑜𝑛 𝑒𝑓𝑓𝑖𝑐𝑖𝑒𝑛𝑐𝑦(%) = 𝑚𝑎𝑠𝑠 𝑜𝑓 𝑙𝑜𝑎𝑑𝑒𝑑 𝑑𝑟𝑢𝑔 𝑚𝑎𝑠𝑠 𝑜𝑓 𝑖𝑛𝑝𝑢𝑡 𝑑𝑟𝑢𝑔 × 100
Several things must be considered when designing a drug loading procedure: one-step or multi-step, the location of the loaded drug, and the type(s) of drug-carrier bindings. Firstly, drug loading can be carried out in a post-synthetic or one-pot fashion, meaning after or during the process of carrier preparation, respectively 65 . The one-pot procedure is more convenient than the multi-step one and is more likely to produce a homogenous distribution of the drug in the DDS 65 . However, it requires more process optimization to combine chemical synthesis and drug incorporation 66 . Examples of how these two types of processes have been applied to different types of carrier materials are summarized by Liu et al. 67 . Secondly, the drug can be loaded in the space inside the carrier or bound to the carrier surface. Encapsulation maximizes the protection of active substances as exposure to the outer environment is limited, whereas the drug attached to the DDS surface might experience premature release. More importantly, the drug can be connected to the carrier through different bonding, covalent or non-covalent type 68 . Note that coordination bond, or coordinative covalent bond 69 , is also considered as covalent type. Non-covalent interactions include van der Waals dispersion force, hydrogen bonding, electrostatic attraction, hydrophobic interaction, and π-π stacking 70 . On the one hand, non-covalent drug loading is generally a facile process, as no formation of chemical bonds is needed. There is also no chemical modification of the drug; thus, its efficacy is kept. On the other hand, covalent bonding results in firmer, hence more difficult to be broken drug-carrier interaction, thus reducing non-specific drug release 68 .

In a typical drug loading process, the API and the carrier are mixed in a solution. There are several parameters to be optimized. The solvent used and the drug concentration are primarily determined by the drug's solubility. The stability of the drug and the carrier in the chosen solvent affects the mixing time and temperature.

The principal driving force of whether and how much the drug is loaded is the drug-carrier interaction 70 , be it non-covalent or covalent. The affinity strength between the drug and the carrier for a given interaction contributes to the drug loading capacity. The strength of drug-carrier interaction must be balanced between drug protection, transportation, and drug release. If the interaction is too weak, the carrier cannot hold the drug to stay connected long enough, leading to premature drug release or drug degradation. If it is too strong, it cannot be broken to enable drug release. Moreover, the carrier integrity should be preserved after the drug incorporation step, which is particularly important for DDS comprised of coordination bonds, as the drug can replace its constituting ligands.

The choice of bonding depends on specific API and carrier material. Understanding the origin of these interactions helps to improve the drug loading and efficiency or to apply one method to other drugs and carriers with similar properties. As an illustration, hydrophobic interaction is beneficial for the encapsulation of hydrophobic compounds. Coordination bond is only possible for certain materials with metallic components. π-π stacking exists only in molecules whose chemical structures consist of conjugated systems.

Surface modification

The motivation of surface modification is to stabilize the carrier and govern its interaction with the biological environment 61 . One can alter the carrier's surface roughness and morphology, change the carrier surface chemistry by oxidation/reduction, adding functional groups, or coat other substances on the carrier's surface 71 . The carrier's surface can be coated with substances that improve physical stability by increasing surface hydrophilicity, introducing steric hindrance, or favoring electrostatic repulsion 63 . That coating can also slow down or prevent degradation and erosion by shielding the carrier from the outer environment.

A source of DDS destabilization in the blood is protein adsorption, potentially causing recognition by the human immune system and subsequent elimination of drug carriers. A popular method for slowing down this process is surface grafting of hydrophilic and flexible polymer chains such as polyethylene glycol (PEG) 60,63 . These polymer chains can form a "cloud of brushes" surrounding the DDS surface. Provided a certain polymer chain length and surface coverage density, the flexibility of PEG chains allows various chain conformations, thus maximizing steric hindrance to protein adsorption 72 . As a result, the PEG-coated carrier has a longer blood circulation time than the non-coated one 73 . Another strategy to escape being marked "foreigner" and captured by the immune system is to camouflage the carrier surface with natural substances such as carbohydrate, lipid, protein, or cell membrane 74,75 .

Surface coating or functionalization can be done using non-covalent and covalent interactions, the same approach used to incorporate drugs. Surface modification is usually done after drug loading steps to avoid interfering with drug loading efficiency. A surface coating should offer the right degree of protection. It should stabilize the carrier long enough but should not prevent drug release and subsequent carrier elimination.

Drug release

This section focuses on mechanisms to control drug release kinetics, which means how fast the drug is released from the carrier. Spatial control of drug release is reviewed elsewhere 76 .

In most cases, an ideal drug release kinetics maintains a constant drug concentration with time, called zero-order kinetics, when the release rate is independent of the drug concentration 12 . The "burst effect" is usually avoided; it is when an initial large portion of the drug is released immediately after the carrier comes in contact with the release medium (Figure 1-5) 77 .

Figure 1-5. Schematic drug release kinetics profile of burst release versus zero-order release 77

Multiple mechanisms of controlled drug release have been reported, described as either the process controlling the release rate 78,79 or the way active substances are transported and released 80 . Fredenberg et al. defined the latter as the true release mechanism and the other as the rate-controlling release mechanism; those two are not necessarily the same processes 80 . Some true release mechanisms are diffusion through water-filled pores, osmotic pumping, and erosion (Figure 123456). In the first type, the drug release rate is controlled by the random movement of drug molecules driven by a concentration gradient between the carrier's inner and outer space 81 . In osmotic pumping devices that confine API inside a semipermeable membrane 78 , the rate-controlling step is the water diffusion from outside through a semipermeable membrane to the core. Osmosis leads to pressure build-up in the core, pushing the drug solution through a fine orifice on the membrane 78 . Thus, the final drug transportation is driven by convection 80 . Regarding erosion mechanism, material dissolution or degradation can deplete and wear out the carrier, leaving the drug exposed to water [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . Notice that most of those theories are based on studies of polymeric DDSs, as they were dominant in the early days of drug delivery, but they can still be used to some extent for studying other types of DDSs. Apart from those classical mechanisms, external stimuli such as light, temperature, magnetic field, ultrasound, etc., can be used to precisely trigger the starting time and location of drug release 83,84 . Understanding the release rate-controlling process is vital for finding ways to optimize the drug release kinetics. Diffusion through water-filled pores mechanism is strongly affected by the carrier porosity. The more porous the material, the easier the diffusion, hence the drug release process. For diffusion from an intrinsic porous carrier where water absorption is fast, the rate-controlling step might be the drug dissolution if its rate is slower than the drug diffusion rate 80 . In the case of a non-porous carrier, there is no gate for the drug to leave the carrier. The drug release rate is dependent on processes that promote pore formation, like swelling and degradation 80 . For instance, a human growth hormone encapsulated in porous, biodegradable polymer microspheres was 100% released before erosion occurred. Meanwhile, the hormone loaded in non-porous ones was not 85 . Another affecting factor is the drug's diffusion coefficient, which is inversely proportional to its size and molecular weight 79 .

Degradation is another popular way to control drug release. For releasing the API chemically connected to the carrier, the bond between the drug and carrier needs breaking. Various degradation reactions have been reported, such as hydrolysis of ester, amide, hydrazone, etc. 86 . Meanwhile, drug-loaded by non-covalent bonding can be liberated from a nondegradable carrier through diffusion or dissolution-controlled erosion mechanism. The carrier degradation mechanism will be discussed in the next part.

Carrier degradation

Degradation, in the context of drug delivery, refers to chemical reactions that decompose the DDS or damage its chemical structure. Note that erosion, also commonly seen in the literature, is the material loss that can result from both physical and chemical processes 87 .

The degradability of a DDS into natural, endogenous, non-toxic products is a highly appreciated feature. For an implant, it means that no removal surgery will be needed. For micro-and nano-sized DDS, biodegradation will relieve the concern of carrier clearance and toxicity. Degradation is also a means to control drug release, as mentioned above.

A compromise between degradation and stability is needed when designing a biodegradable DDS. The DDS should be stable when transporting the drug to the intended site but degradable during or after drug release. One can slow down or prevent the degradation process by surface coating or increase degradability by adding functional groups susceptible to hydrolysis. It can be more complicated to control the stability of biodegradable DDS than non-degradable DDS, as there are more variables in the equation of carrier stability. It is also equally important to check if degradation products are non-toxic and can be removed safely from the body.

The known sources of degradation in the living body are water, enzymes, reactive oxygen species, and salts 88 . Water is present in blood, tissues, and organs everywhere in the human body. Consequently, hydrolysis is expected to be the leading cause of degradation. In addition to the chemical nature, water diffusion and sorption parameters also vary between materials, influencing the rate of the hydrolysis reaction, hence the degradation time 88 . Enzymes can catalyze hydrolytic and oxidative degradation 88,89 . Hydrolases, a family of enzymes that include proteases, esterases, glycosidases, and phosphatases, are known to accelerate the hydrolysis of certain degradable polymer 89 . In addition, highly reactive oxygen species, including hydroxyl radical (OH), superoxide (O 2-), and hydrogen peroxide (H2O2), can be produced during the body's inflammatory response to foreign materials 89 .

Possible oxidation of biomaterials in the body caused by free radicals and the involvement of oxidative enzymes was reviewed by Williams and Zhong [START_REF] Williams | Biodeterioration/biodegradation of polymeric medical devices in situ[END_REF] . Salts can catalyze the hydrolysis by changing the environment pH 88 , or some anions present in the blood, such as bicarbonate and phosphate [START_REF] Conley | Encyclopedia Britannica[END_REF] , can act as Lewis bases or nucleophiles.

How the carrier degradation reactions can happen depends highly on the type of chemical bonds present in the carrier 87 . For example, the principal degradation mechanism of organic materials like polymers is hydrolysis, forming monomers and oligomers, provided that they consist of water-reactive groups like anhydride, ester, and amide. Inorganic materials like metal nanoparticles are susceptible to acid etching which could produce free metal ions [START_REF] Mohammapdour | Mechanisms of immune response to inorganic nanoparticles and their degradation products[END_REF] . Hybrid materials comprised of coordination bonds can undergo both hydrolysis and ligand replacement [START_REF] Canivet | Water adsorption in MOFs: fundamentals and applications[END_REF] .

The carrier degradation reactions can occur in a homogenous or heterogenous fashion.

Homogenous degradation leads to bulk erosion, whilst the latter leads to surface erosion (Figure 1-7) 79,87 . Surface erosion (Figure 1-7A) is characteristic of the decrease of carrier size over time, as the degradation reaction only happens at its surface. This scenario could occur for carriers made of a hydrophobic material whose interior is protected from water exposure. In bulk erosion (Figure 1-7B), attacking agents penetrate the carrier's interior everywhere and initiate a degradation reaction. The carrier dimension remains relatively unchanged until a fast disintegration occurs. Note that surface and bulk erosion are extreme cases, and a single DDS can experience both phenomena to different extents. In the case of drug release through degradation of a drug-loaded carrier, surface degradation is preferred over bulk reaction, as the release rate in the first case is more gradual and correlates with the degradation rate.

Figure 1234567. Illustration of (A) surface versus (B) bulk erosion 79,[START_REF] Siepmann | Mathematical modeling of bioerodible, polymeric drug delivery systems[END_REF] In this sub-chapter, we have learned about the motivation of drug delivery and how controlled drug delivery can help to improve existing disease treatment methods and make new ones. We have also seen some of the chemistry behind DDS formulation and performance. With this basic, next, we will review the drug carriers made of nanomaterials that have intrigued great attention and triggered contentious debate from the scientific community.

1.2 Nanocarriers for drug delivery

Generalities about nanoscale drug delivery systems

The boom of nanotechnology in the 21 st century has spread into medicine and healthcare applications, leading to the birth of nanomedicine, the application of nanotechnology in biomedical applications, including diagnostic, therapeutics, and sensors. By definition, a nano-sized object is an object that has at least one dimension below one-billionth of a meter [START_REF] Nni | Nano"? National Nanotechnology Initiative[END_REF] . In drug delivery, nanocarriers have a size between 1 and 1000 nm [START_REF] Duncan | Nanomedicine(s) under the Microscope[END_REF] .

The most attractive advantage of nanocarrier comes from the effect of size and miniaturization. The devil lies at the surface. Given the same volume, a nanoobject possesses a much higher surface area than a micro-or macro-size object. This increase in surface area to volume ratio results in the appearance of properties and reactivities that are not observed or not prominent at bulk scale [START_REF] Roduner | Size matters: why nanomaterials are different[END_REF] (Figure 1 -8). For instance, nanocrystallization of a drug is a method to increase its solubility based on increasing surface area. The nanosize also permits nanocarrier to be used for intravenous formulation, which is used for many anti-cancer drugs with too low water solubility to be orally administered. Nanocarriers also bring possibilities to create multi-functional DDS using the platform of nanoparticle (NP) surface. A particularly attractive function is targeting either cell, tissue, or tumor. Furthermore, owing to the high surface area that allows maximum interaction with molecules and substances in the body, it is believed that NPs can pass through biological carriers that other types of DDSs cannot 60,64 . For instance, it was reported that the adsorption of gold NPs through the gastrointestinal tract increases when decreasing their diameter from 58 to 4 nm. Such influence of varying extent on the biodistribution and in vivo fate of nanocarriers have also been observed by modifying other physiochemical properties such as morphology and surface structure. In addition, the quantum effect also brings some unique properties to inorganic materials [START_REF] Roduner | Size matters: why nanomaterials are different[END_REF] .

However, the main concerns for using drug nanocarriers are their safety, especially in the long term. The same high surface area that makes NPs influence drug biodistribution can make them too reactive and create possible toxic products [START_REF] Barnard | Knowledge is our first defence[END_REF] . The NP clearance is also an important issue, as sometimes the small-size carriers are eliminated too quickly 64 that they cannot perform their intended functions.

In brief, owing to the nano size, drug nanocarriers show a particular relation between their physiochemical properties and their interaction with the body, hence their drug delivery functions and performances. This feature can be both an opportunity and a challenge for using nanoparticles as DDS.

The most reported NPs in the literature are lipid, polymeric, and inorganic NPs (Figure 123456789). Lipid-based NPs are typically vesicle or spherical structures composed of at least one lipid bilayer 64 , where a lipid molecule consists of a hydrophilic head group and a hydrophobic chain [START_REF] Tenchov | Lipid Nanoparticles─From Liposomes to mRNA Vaccine Delivery, a Landscape of Research Diversity and Advancement[END_REF] . Liposome, considered the first nanostructure used as DDS, is a vesicle surrounded by a phospholipid bilayer membrane that can spontaneously form unilamellar or multilamellar structures in water [START_REF] Tenchov | Lipid Nanoparticles─From Liposomes to mRNA Vaccine Delivery, a Landscape of Research Diversity and Advancement[END_REF] . It was discovered by Bangham 100 and was also the first nanoparticle formulation approved for clinical use 101 . Another subclass of lipid-based NPs is lipid NPs, in which micellar structure is formed within the particle core 64 . Hydrophilic drugs can be loaded in the NPs' aqueous interior, and hydrophobic drugs can be trapped in the hydrocarbon chain region. They have been particularly suited for the delivery of nucleic acid for gene therapy applications. This lipid-based class of NPs has been popular because of the simplicity of their formulation processes, easy self-assembly of NPs, biocompatibility inheriting from endogenous lipids, large cargo loading capacity, and various controllable physiochemical properties. They account for most of FDA-approved nanomedicines 102 .

Polymeric NPs are a big class of NPs because of a wide variety of polymer compositions and architectures. Various methods for NPs preparation and drug loading are available. Delivery of hydrophobic and hydrophilic compounds, of small and macromolecules, have been reported. Several polymer NPs architectures are dendrimer, polymeric micelle, polymersome, nanocapsule, and nanosphere 64 . Dendrimers are hyperbranched polymers with three-dimensional architectures, mainly investigated for the delivery of nucleic acid and small molecules. Small molecules are loaded in the interior, while macromolecules can be conjugated to the exterior branches through a functional group. Polymeric micelles, made of amphiphilic block copolymers, are formed by the self-assembly of hydrophilic moieties towards the outer water and hydrophobic moieties towards the inner. They are suitable for capturing hydrophobic drugs. Nanospheres are probably the most populous in polymeric NPs as they can be prepared from most polymers. Polymeric NPs can be biodegradable, mainly by synthesizing a water-hydrolyzable polymer backbone. This class of polymer-based NPs is popular mostly because they have long-known biocompatibility and are easily modulated for different functions. However, they also have a higher risk of particle aggregation and toxicity compared to liposomes.

Inorganic NPs reported for drug delivery applications are mainly gold, iron, iron oxide, and silica 64 . They are known for the possibilities of precision engineering of NPs size, structure, and geometries. They also possess some unique properties due to the material's intrinsic properties. For instance, gold NPs have size-dependent optical properties that are beneficial for sensor applications as well as plasmonic properties useful for photothermal therapy 68 . Iron oxide NPs make up most of the FDA-approved formulations of inorganic NPs 102 . They are preferred for trusted biocompatibility and the iron's specific magnetic properties that can be used for imaging or magnetic-guided drug delivery. Mesoporous silica NPs are probably the first porous material used for DDS. Their porosity raises drug loading capacity more than the non-porous drug carrier. Inorganic NPs with their unique properties are beneficial for niche applications that cannot use organic materials. They also have good stability. The disadvantages are mostly toxicity and concerns caused by the pristine materials or their degradation products [START_REF] Mohammapdour | Mechanisms of immune response to inorganic nanoparticles and their degradation products[END_REF] . Metal-organic framework NPs recently emerged as a new candidate for DDS 66,103 . They offer much higher porosity than silica NPs. Their diverse chemistry enables tunable drug delivery functions and adaptable encapsulation of compounds with different properties. However, their clinical applications are limited due to toxicity concerns and complicated degradation mechanisms.

To summarize, the availability of plentiful classes of NPs with a wide range of chemistry nature and properties has added another benefit for nano-size DDSs. Evidently, there is no one-size-fits-all solution, and one should understand the pros and cons no matter the choice of material.

Despite having many beneficial features for controlled drug delivery, the current landscape of drug nanocarriers is debatable. Inheriting the excitement of nanotechnology, nanomedicine raises a lot of hope for creating the "magic bullet" that can offer both spatial and temporal control for drug delivery, especially for targeting tumors. Extensive research on NPs for drug delivery has generated promising results and many clinical trials. However, the development rate is disappointing compared to the effort and resources spent 45 . To date, only a few formulations have been approved by FDA 102 . One of the reasons is the significant discrepancy in physiology between small animal models and human 39 and how that difference influences the function of nanocarriers in the body 46,64 . The current dogma in tumor targeting, the EPR effect, is also questioned 104 . To go beyond the formulation-first approach 40 , scientists have advocated for rational design based on the target biological barrier 105,106 , more collaboration with clinicians 107 , and more mechanistic understandings 47,41 .

A basic understanding of their chemistry is important, as many physical chemistry processes and reactions such as electrostatic, hydrophobic, hydrogen bonding, and metal-ligand coordination are the driving forces behind the interaction between NPs and biological substances 108 . Inadequate reports of NPs' physiochemical properties can cause difficulties in solving their toxicity issues 109 .

In this thesis, we are interested in studying nanoparticles made of polymer and metalorganic framework that have been extensively studied in our research group. While the first drug delivery system, polymer NPs are popular in literature, the second DSS, metal-organic framework NPs are relatively young and lesser known in the drug delivery and controlled release community.

Polymer nanoparticles 1.2.2.1 Basics about polymer and its applications for drug delivery

Polymer is a substance made of hundreds to millions of identical units linked together by covalent bonds. Those repeating units are called monomers. A polymer can be visualized as a chain of many small beads (monomers) 110,111 . The simplest polymer is made of a single type of repeating unit, called homopolymer, which can be represented as -[A-A-A-A-A-A]-. Polymer made of more than one type of repeating unit is called a copolymer, whose representation can be -[A-B-A-B-A-B]- 111,112 . The number of monomers in the polymer chain, called the degree of polymerization (DP), is usually at least 100. When it is lower than 10, the substance is considered oligomer 113 .

The chemical structures of several polymers proposed for biomedical applications are presented in Figure 1-10. Notice that the end unit, or the terminal group, is different from the middle parts. In addition, a polymer is not necessarily composed of linear chains. Other polymer architectures are branched chains and cross-linked networks.

Polymer is characterized by its molecular weight (MW) and MW distribution. A polymer of a given mass comprises multiple polymer molecules with different values of MW 114 . Thus, the molecular mass of a polymer is usually defined as the weight-average Mw or the numberaverage Mn. The ratio Mw/Mn, called the polydispersity index, characterizes the MW distribution. A polydispersity index of 1 means all polymer molecules within a given mass have the same MW.

Another important polymer property is polymer crystallinity, arising from the regular alignment and packing of polymer chains. However, polymer rarely has a perfect crystallinity, thus being only semicrystalline or amorphous 114 . There are both crystalline and amorphous regions/domains within a polymer sample. As the crystalline polymer region is less permeable to water than the amorphous one, the degree of crystallinity affects the polymer's swelling, solubility, and hydrolytic rate 81 . In addition, crystallinity contributes to the polymer's rigidity, but the presence of an amorphous domain makes the polymer flexible and not brittle.

While natural polymers are essential compositions of the human body, e.g., DNA and proteins, synthetic polymers are present everywhere, from daily use like clothing, food packaging, water pipe, etc., to specific applications like bulletproof vests, hip joint prosthetics, and surgical tools 115 . Polymer has been popular since day one of controlled drug delivery. Macro-and microscale DDSs mentioned in section 1.1. were primarily made of polymeric material. The rise of biodegradable polymer started from the development of PGA, poly(glycolic acid) as degradable suture, followed by the development of poly(lactic-co-glycolic acid) PLGA and poly(lactic acid) PLA 24,81 . Later, polymer was also one of the first examples of nanomedicine; the idea of the drug-polymer conjugate was proposed by Ringsdorf in 1975 and developed by an international team led by Kopecek in the subsequent years 24 . At present, polymeric nano-sized DDS have been designed in many configurations, including drug-polymer conjugate, polymeric micelle, and dendrimer 116,117 . Polymeric nanocarriers have been studied intensely because of their structure versatility, long-known biocompatibility, and biodegradability81, as well as the expansive chemistry strategies available for polymer synthesis and polymer NPs functionalization 117 . Among various drug delivery applications of polymer NPs 118 , they are particularly attractive for cancer tumor targeting and diagnostics 119,120 . A few polymer NPs are approved by FDA 64 , and many are in ongoing clinical trials 102 .

Polymeric NPs can be categorized as nanospheres (solid matrix systems) or nanocapsules (cavities surrounded by polymeric shell or membrane) 64 . Nanocapsules trap drugs in their core, whilst nanospheres have both core and surface drug loading fashion 118 . The following section focuses mainly on the nanosphere-type polymer NPs.

Choice of polymer

An extensive library of polymer families used for controlled drug release has been built [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . Probably, the most common feature to be chosen is degradability. Non-degradable polymers have C-C backbones, such as poly(ethylene), poly(vinyl alcohol), and poly(acrylic acid). Degradable polymers must have a backbone containing hydrolyzable bonds, e.g., ester, amide, and anhydride. Non-biodegradable hydrophilic polymers are used to prepare swollen cross-linked polymer networks, i.e., hydrogel 121,122 . Non-biodegradable hydrophobic polymers are fabricated as slow-dissolving membranes for dissolution-controlled drug delivery devices 78 .

Hydrophilicity is also a key parameter for biodegradable polymers. Affinity towards water can promote or impede hydrolytic degradation. Thus, the extent of hydrophilicity affects the reaction rate and mechanism. Polymers' ability to absorb water depends on structural parameters such as chain flexibility, degree of crystallinity, and the presence of hydrophilic groups (amino, carboxyl, and hydroxyl). High flexibility and low degree of crystallinity promote hydrophilicity 114 . In addition, polymers that are more-water-soluble, thus fasterdegrading, generally have low MW and highly branched structure 81 . The hydrophilicity of a copolymer can be adjusted by varying the molar ratio between hydrophilic and hydrophobic monomers.

Poly(ester) is the most studied class of biodegradable polymers; some of them are presented in Figure 1-11. The most widely known system is polymers of hydroxy acids, such as poly(lactic acid) PLA and its copolymer with glycolic acid PLGA. PLA is considered biocompatible since lactic acid is present and produced by the human body. The physical properties of PLA can be tuned by controlling its stereo-configuration through the ratio of d-and l-isomer [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . The idea of PLGA copolymer is to combine fast-degrading PGA and slowdegrading PLA 81 . It was shown that by varying the lactide-to-glycolide ratio, 50:50, 75:25, and 85:15, the degradation time of PLGA in an aqueous environment increased from 1-2 to 5-6 months 81,123 .

Another common degradable aliphatic polyester is polycaprolactone PCL. Because of its extremely long degradation time of 2-3 years, PCL is more used for degradable implants than drug nanocarriers 81 . The application of PCL for polymeric DDS is their amphiphilic copolymer which is useful for making polymeric micelles. Poly(anhydride) has a highly water-sensitive functional group, yet its degradation mechanism can be controlled by varying the C-C moiety of its monomers [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . Poly(anhydride) containing hexyl group degrades 3-time slower than its shorter C-C chain methyl analogues 124 .

Poly(amide) for DDSs is mostly made from the polymerization of amino acids, natural ones like glutamic acid and lysine 81 . It is a favored system because of its established literature on synthesis techniques. Its degradation rate is dependent on the hydrophilicity of the constituting amino acids. The copolymer of amino acid with lactic acid is also attractive because the additional functional groups like carboxylic and amino allow further polymer modification [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] .

Polymer NP synthesis and preparation

Polymer NPs have been prepared by bottom-up or top-down techniques 119 . The bottom-up approach uses the polymerization of monomers to synthesize NPs, while the top-down approach involves the self-assembly of pre-synthesized and characterized polymers.

Several techniques of polymerization to produce NPs are emulsion-and interfacial polymerization 125 . Emulsion precipitation, the most used method in the industrial production of polymer NPs, is a broad category of several techniques: conventional emulsion polymerization, mini-emulsion polymerization, and microemulsion polymerization 126 . A conventional process uses a mixture of water, a monomer of low water solubility, a watersoluble initiator, and a surfactant 125 . Polymerization starts when monomer molecules collide with the initiator, then polymer chain growth is followed by phase separation and formation of solid particles. Solid particles are stabilized in water by surfactants. As this conventional scheme requires a time-consuming removal step of excess surfactants, surfactant-free emulsion polymerization has been developed. Instead of surfactants, polymer NPs are stabilized by ionizable initiators or ionic monomers. Another method is mini-emulsion, of which the first step is the formation of micelles that contain monomers in their core surrounded by surfactants. Then polymerization of monomer takes place due to the use of a high-shear device like ultrasound 127 . In another strategy, interfacial polymerization involves oil/water emulsification dissolution of two monomers in the oil and water phases, followed by polymerization at the liquid-liquid interface. The formation of nanocapsules or nanospheres can be favored using an aprotic or protic solvent. This method was used by Bouchemal et al. to prepare polyurethane nanocapsules with mean diameters between 200 and 300 nm 128 .

The top-down approach in which NPs are prepared from the pre-formed polymer is also popular, as it uses fewer reacting agents than the bottom-up approach 81 . Nanoprecipitation and emulsion are the most known methods. In nanoprecipitation, also known as solvent displacement, dropwise addition of polymer solution (organic phase) to water (aqueous phase) under stirring leads to the spontaneous formation of nanodroplets, thus leading to self-assembly of polymer into sphere 81 . The organic solvent is removed later by stirring or pressure-reduced evaporation. Authors who reported nanoprecipitation for the first time argued that rapid diffusion of the solvent from the organic phase to the aqueous phase causes a decrease in interfacial tension and migration of polymer insoluble in water towards the liquid-liquid interface 129 . The precipitation process of polymer into NPs was governed by the Marangoni effect, which describes movement at the liquid-liquid interface. Acetone is the most frequently used solvent for this method 125 . Surfactant is optional, but its use helps to increase NPs colloidal stability over time 129 . Nanoprecipitation is an easy technique applicable for both homo-and co-polymers and can be used for both NPs preparation and drug loading. For example, the diblock copolymer PLGA-PEG was coprecipitated with a hydrophobic drug 130 . The resulting NPs, with a diameter of around 150 nm, have a hydrophobic PLGA core confining the drug and a shell composed of hydrophilic PEG blocks. Also, the loading of hydrophilic drugs is possible by replacing water with methanol and ethanol 131 .

Emulsion, as the name suggests, is based on the emulsification process, mixing two immiscible liquids to disperse one liquid (disperse phase) in the other (continuous phase). Different types of emulsions can be formed, hence the different categories of emulsion methods: oil-in-water (O/W), water-in-oil (W/O), and double emulsion water-in-oil-in-water (W/O/W). After the emulsion formation step induced by vigorous mixing, polymer molecules self-assemble into spheres, the most energetically stable form. The mixture is then homogenized by sonication or sheer to produce monodisperse NPs suspension; and finally, the organic solvent is evaporated 81 . An emulsifier is compulsory to stabilize the emulsion. Several ingredient formulas for this preparation method were summarized by Rao et al. 125 . The O/W emulsification is used to load hydrophobic drugs, while the W/O/W is used for hydrophilic drug 81 . In the single emulsion O/W technique meaning the oil is dispersed in the water phase, the polymer and the drug are dissolved in a water-immiscible organic solvent such as dichloromethane or ethyl acetate. This organic phase is emulsified in an aqueous phase containing surfactant. In the double emulsion W/O/W, the drug and surfactant are dissolved in water; this aqueous phase is emulsified in the organic phase containing polymer, forming a W/O emulsion. This W/O emulsion is then emulsified in a larger volume of water with or without surfactant to yield the W/O/W emulsion 81 .

Microfluidic is a promising method to prepare NPs with narrow size distribution and better batch homogeneity than the previously mentioned methods 81 . Manipulating a minimal volume of liquid in microscale channels allows streamlining and simultaneous optimization of multiple steps, e.g., sample preparation, mixing, and separation on a single device or system 119 .

Previous studies have explored how formulation-related factors influence the resulting NPs size. Factors that have a major impact are polymer concentration, polymer MW, concentration and nature of surfactant, the choice of organic solvent, and the ratio between organic and aqueous phase [132][133][134] . For nanoprecipitation, some effects are possibly explained by the diffusion coefficient of the organic solvent in water 133 , as nanoprecipitation is driven by the fast diffusion of organic phase droplets into a sea of water. Specifically, the size and polydispersity index of PLGA NPs were observed to increase when they were prepared in acetonitrile, acetone, and tetrahydrofuran, in the order given, following the decreasing order of their diffusion coefficients: acetonitrile > acetone > tetrahydrofuran. For the emulsion technique, a sufficient surfactant concentration is required to obtain small particles with satisfactory polydispersity 135,136 . The size polydispersity seems to increase with a higher fraction of the organic phase. In addition, smaller and more monodispersed NPs were obtained with polymer with higher MW 135,137 , probably because increasing polymer hydrophobicity led to more efficient self-assembly and packing of polymer.

Drug loading

There are three main strategies for loading drugs in polymer NPs: surface adsorption, encapsulation, and covalent attachment to polymer backbone 119 . In the first strategy, the API is adsorbed on the polymer NP surface after the NP preparation 138 . The drug loading capacity is thus dependent on the polymer hydrophobicity and the NP surface area. However, the drug payload was lower than core encapsulation 139 , and weak drug-NPs interaction could lead to burst release. Therefore, surface adsorption deems a dated method for polymer NPs. The second strategy, encapsulation, is the most reported in the literature, as it is straightforward to implement 119 . The drug is incorporated inside polymer NPs during its preparation using one of the methods described above for dispersion of pre-formed polymer. Hydrophobic drugs can be dissolved in the organic phase, and hydrophilic drugs can be dissolved in the aqueous phase. An alternative method for incorporating watersoluble drugs is chemically attaching drugs to the terminal of the polymer backbone; then, the resulting drug-polymer conjugate is formulated to NPs as usual 119 . The obtained drug payload and EE% by chemical conjugation (3.5 wt% and 96.6%) were significantly improved than loading free drug in polymer NPs by emulsion (0.3 wt% and 6.7%) 140 . The drug conjugated to polymer is more hydrophobic than the non-modified one, which reduces unwanted drug diffusion into the aqueous phase during NPs preparation. Another advantage of this polymer-drug conjugation strategy is the possibility to incorporate more than one active substance 119 and extended drug release owing to stable drug-carrier bonding 140 . The covalent linkages between polymer and drug in the polymer-drug conjugates have been ester, amid, imine, acetal bonds, etc. 141

Processing parameters of drug loading methods could be optimized. Regarding the emulsion method, decreasing the volume of the organic phase for the same volume of the aqueous phase considerably raises the payload and EE% of a hydrophobic compound 142 . This observation was explained by the increasing viscosity of the organic phase, i.e., the polymer solution, preventing the drug diffusion from the organic phase to the aqueous phase. In addition, a low fraction of organic solvent leads to quick evaporation; consequently, the cargo is better captured in the polymer 134 . Likewise, to reduce the solvent evaporation time, other authors used a rotating evaporator instead of stirring in the atmosphere 135 .

Meanwhile, low payload and low EE% of hydrophilic drugs associated with their leaking from the polymer core to the external aqueous phase can be limited by changing the pH of the aqueous phase. Thus, their water solubility and hence their portioning to external water decreases. This strategy has been applied to both nanoprecipitation and double emulsion W/O/W techniques 143,144 .

Surface modification

Attempts to improve colloidal stability, increase blood circulation time of polymer NPs or add more functions like targeting were made by modifying their surface with an amphiphilic surface coating or incorporating hydrophilic moieties in the polymer chain before NP preparation 138 . In the first method, coating substances can be attached to the polymer NP surface through physical and chemical adsorption. Regarding the physical adsorption strategy, surface modification is carried out simultaneously with polymer NPs preparation. Surfactants, such as poloxamer, polysorbate, and polyethylene glycol (PEG), are mixed in the aqueous phase of the formulation. These amphiphilic surfactants adsorb on polymer NP surface by their hydrophobic part while their hydrophilic part is in contact with water, accordingly stabilizing the colloidal suspension and efficiently shielding protein adsorption and increasing blood circulation time 145 . For example, polystyrene NPs coated with poloxamer and poloxamine displayed longer circulation time and less liver accumulation in rabbits than uncoated ones 138 . In another demonstration, the PEG coating helps PLA NPs to escape from the uptake of monocytes, a type of white blood cell, which results in longer circulation time. As a result, blood concentration and tumor concentration of the drug delivered by coated NPs are higher than plain ones 146 . Also, the PEG coating can improve the chemical stability of polymer NPs. The team of Alonso observed that upon incubation in acidic simulated gastric fluid, only 3% of PEG-coated PLA NPs were degraded, while 9% of non-coated PLA NPs did 139 . Another coating molecule for polymer NPs is lipid, whose surface adsorption can be driven by electrostatic or hydrophobic interaction 147 . For instance, a bilayer of synthetic cationic lipid was formed on polystyrene NPs through interaction with the negatively charged sulfate group of polystyrene 148 . Meanwhile, taking advantage of hydrophobic interaction, a monolayer of neutral lipids can self-assemble on the NP surface.

Many examples have been demonstrated with PLGA NPs 147 .

Another method to achieve surface coating is covalent attachment after NPs preparation, frequently used for surface grafting of targeting ligands. Published studies have used a coupling agent to activate the polymer's terminal group or the targeting ligand. For example, the carboxylate group of folic acid, a tumor-targeting ligand, was transformed into the carboxylate ester with N-hydroxysuccinimide (NHS). The resulting NHS-folate ester was added to an NP suspension of copolymer between amino-PEG and polycyanoacetate 149 . The reaction between the NHS-folate and PEG amino group led to the formation of a covalent amide bond and successful surface attachment of folic acid on these polymer NPs.

In the second style of modifying surface polymer NPs, block copolymer or polymer blends with hydrophilic segments have been synthesized. NPs prepared from these polymers will spontaneously expose the hydrophilic part to water while keeping the hydrophobic chain in the core, making them stable colloids. Gref et al. introduced a one-step method to prepare the PEG-coated NPs from PEG-polyester block copolymers in which PEG is hydrophilic and polyester is lipophilic 73 . NPs made of these amphiphilic copolymers have a micelle-like structure with a hydrophobic core and hydrophilic surface. They form a stable suspension without a surfactant which is not always compatible. The authors observed a proportional correlation between polymer NPs' blood circulation time and the MW of PEG, which was explained by a probable thicker PEG layer that offers better protection against protein adsorption. Likewise, a copolymer between PLA and polyethylene oxide (PEO) was synthesized. Increasing the density of PEO on the NP surface led to more efficient repulsion to blood proteins 150 .

The long-circulating properties of PEG-coated and PEO-coated polymer NPs are attributed to the protection against protein adsorption on the NP surface through steric hindrance.

The hypothesis is that flexible and hydrophilic polymer chains form a dense "cloud" surrounding NPs 72 . This phenomenon was theoretically explained by Joen and coworkers considering van der Waals attraction and steric repulsion between protein and hydrophilic PEO layer, as well as the hydrophobic attraction between protein and polymer NPs core 151 . They found that the steric repulsion effect, determined by the corresponding free energy per unit surface area, increases with high surface density and long chain length of PEO.

Possible influences of PEG surface density and PEG conformation on the protein adsorption on NPs are illustrated in Figure 1-12. The optimum surface density of PEG on NPs was reported when the distance between PEG chains attained 5.1 nm 138 , and no further reduction of protein adsorption was found with a distance shorter than 5.1 nm. Another mechanism by which surface coating increases the NPs' blood circulation time is surface neutralization. It was reported that polymer blend NPs of PLA and PLA copolymer with monomethoxy polyoxyethylene (MPOE) had the zeta potential equal to 0 mV when the MPOE content in the blend was more than 3% 152 . Also, the zeta potential increased from -55 mV for PLA with no MPOE to 0 mV with higher MPOE content and longer PLA-MPOE chain, suggesting that MPOE chains on the surface shield the carboxylate end group of PLA, thus reducing NPs elimination. First, the coating obtained by physisorption after polymer NPs preparation was too loosely held that the amount of bound albumin was not reproducible. Second, the coating with higher surface coverage is not always better. The in situ surface coating during NPs preparation did provide a denser albumin coating, but it was partially denaturized, probably due to exposure to organic solvent and sonification during the NPs preparation. The most suitable method was adsorption after NPs preparation and NPs' polydopamine coating, which produced a reasonably distributed albumin coating with more preserved activity.

Drug release

The drug release mechanisms of polymer NPs were extensively studied in experimental and theoretical approaches 154 , inheriting the classical theories on macroscale DDS 78,79,155 .

Fredenberg summarized that the main release mechanisms are diffusion through waterfilled pores, diffusion through polymer matrix, osmotic pumping, and erosion 80 .

Diffusion through water-filled pores is applicable for swelling polymers that absorb water and increase their volume when being immersed. This water penetration and swelling can create water-filled pores 156 . Over time, more pores and larger pores are created. Then, the drugs trapped in the polymer matrix can diffuse through these pores to the outer environment. The drug release rate is dependent on the polymer's porous structure and any processes that promote pore formation or pore closure, such as swelling and erosion 80 . Kim and Park created porous PLGA microspheres using an extractable porogen and then controlled the drug release kinetics by treatment with water-miscible solvent to induce partial pore closure and thus slowing drug release 157 . Microscopy images of porous and pore-closed microspheres and their drug release kinetics are given in Figure 1-13. In this case, pore closure hinders diffusion, hence drug release (Figure 1-13C). In addition, slow erosion due to slow polymer degradation can also contribute to the pore formation process, thus increasing the diffusion rate. Diffusion of the drug through polymer matrix was found for the release of small hydrophobic drugs. For example, the diffusion coefficient of boron-dipyrromethene (BODIPY) loaded in PLGA microparticles did not change when PLGA degraded or in the presence of a second loaded cargo, but it considerably increased with rising temperature 158 . Thus, it was concluded that diffusion through polymer matrix is the primary mechanism, and the porosity is unimportant. However, the physical properties of the polymer, such as its crystallinity and MW, play a considerable role 80 . Drug diffusion through polymer with lower MW is preferred due to the more flexible polymer chains.

Osmotic pumping is the convection-driven drug transport through water-filled pores 80 . The polymer NPs should be non-swelling in this case.

Erosion as the true release mechanism is described as "drug release without drug transport".

Here, the drug does not actually move but is exposed to the outer environment due to the disappearance of the carrier caused by polymer dissolution or degradation. Two types of erosion are reported: bulk and surface erosion. Bulk erosion happens when the water penetration is faster than the polymer bond cleavage and vice versa. In surface erosion, the polymer degradation starts at the DDS's surface and gradually depletes the polymer from the surface to the core. Bulk erosion occurs for polymers with fast water hydration, i.e., water adsorption leading to homogenous water hydrolysis throughout NPs, leading to the sudden disintegration of the whole NPs, which offers less control on drug release kinetics than surface erosion 87 .

Polymer-drug conjugate NPs are a particular type of polymer-based DDSs that must be degradable and erodible to enable drug delivery. The covalent bond between the drug and the polymer backbone should be dissociated by hydrolytic degradation, and erosion is needed to let the degradation products and the drugs leave the polymer matrix. Trapping of dissolved degradation products can have both negative and positive impacts on polymer degradation, hence the drug release rate 80 . Systems with ester and amide linkages usually undergo slow hydrolysis, hence slow drug release in weeks and months. At the same time, systems with acetal, imine, and hydrazone can release the drug in the range of several hours to days 141 . The hydrolysis can be catalyzed by acid or bases (Figure 1 -14). The properties of polymers and drugs, as well as their interactions, can affect the drug release process. For example, a biodegradable polymer composed of a water-hydrolyzable but hydrophobic backbone might favor surface erosion, thus releasing the drug in a more controlled manner [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . A PLGA of low MW and low L:G ratio (lactic acid/glycolic acid) is more hydrophilic, which increases their water adsorption affinity, thus promoting hydrolysis and erosion 81 . High solubility of the active molecules facilitates drug diffusion through waterfilled pores. For diffusion through polymer matrix, a certain degree of solubility is needed to allow drug partitioning from the polymer carrier phase to the outer aqueous phase 79,80 .

Chapter 1 Introduction -Literature review Different extents of polymer-drug interaction due to the choice of drug loading method are directly related to the drug release kinetics. For example, a burst release of 70% after one hour was reported for the drugs loaded by surface adsorption 159 due to weak drug-carrier interaction. Release of doxorubicin (DOX) chemically conjugated to PLGA NPs was complete after 25 days, compared to 5 days for drug encapsulated inside NPs 140 . The DOX-PLGA conjugate NPs release drugs principally through the slow degradation of PLGA NPs, while the other system release drug through a diffusion mechanism.

Polymer NP degradation

The degradation of polymer NPs for drug delivery is described as the "chain scission process" during which polymer bond dissociates to form its oligomers and monomers 87 . All degradable polymer NPs reported in the literature are designed to undergo hydrolysis reactions, ideally forming their monomers. For example, the ester bond is hydrolyzed into carboxylic acid and alcohol, the amide is hydrolyzed into carboxylic acid and amine, and the anhydride is hydrolyzed into carboxylic acid. Hydrolysis reactions in the body can be catalyzed by enzymes such as esterase and lipases 160 .

Several factors that govern the degradation rate of polymer NPs are the type of chemical bond, pH, copolymer composition, and water uptake 87 . The chemical nature of the biodegradable polymer backbone is probably the most critical factor, precisely, the reactivity towards hydrolysis of water-sensitive functional groups. Anhydrides and ortho-esters are the most reactive; their polymers' half-lives are 0.1 and 4 h, respectively, whereas ester and amides are much less reactive; the half-lives of their corresponding polymers are estimated as 3.3 and 83000 years, respectively 87 . Amides are especially stable because of the electron delocalization between lone pairs of electrons of nitrogen with the carbonyl group 141 . Nevertheless, those half-lives can be modified with catalysis or by chemical modification of these functional groups. For instance, 1-chlorine ethyl acetate has a 100 times faster hydrolysis reaction rate in neutral media than ethyl acetate, owing to a negative induction effect 87 . Ethyl ester of long alkyl carboxylate chains is hydrolyzed significantly slower than short chains 141 . A study on polyanhdride showed that the degradation rate drastically decreases when the number of methylene groups in the polymer backbone increases, thus making the anhydride bond less reactive 124 . Usually, polymers composed of reactive functional groups tend to degrade quickly and follow surface erosion mechanisms, while those comprised of less reactive functional groups tend to be dominantly bulk eroding [START_REF] Siepmann | Mathematical modeling of bioerodible, polymeric drug delivery systems[END_REF] . It should be noted that both fast and slow degrading polymers have found their place for drug delivery through chemical conjugation 141 .

The pH plays an important role in dictating the degradation rate of polymer NPs, as water hydrolysis can be catalyzed by acid or base 141 . The degradation rate of a poly(anhydride) decreased ten times when the pH of the degradation medium went from 7.4 to 10 124 . Meanwhile, the hydrolysis of ester was accelerated by both acid and base 87,160 . The autocatalysis effect explained the faster reaction of PLA at low pH. Its degradation product, lactic acid, further catalyzed the hydrolysis by lowering the pH 161 . This autocatalytic effect highlights the importance of the internal environment pH in the polymer NPs degradation mechanism. It was discovered that the degradation inside PLA-based DDS was faster than at their surface [START_REF] Siepmann | Mathematical modeling of bioerodible, polymeric drug delivery systems[END_REF] . Meanwhile, the poly(ortho ester) degradation is only acid-catalyzed because the hydroxide ions cannot attack the carbon of the orthoester bond [START_REF] Siepmann | Mathematical modeling of bioerodible, polymeric drug delivery systems[END_REF] .

Water adsorption, i.e., hydration, is a crucial step in polymer NPs degradation. Hydrophilic polymers take up a large quantity of water, which increases their degradation rate.

Varying a copolymer composition can change the crystallinity or hydrophobicity of a copolymer, which in turn influences its degradation process 87 . An example is PLGA, whose MW decreased faster with decreasing L:G ratio. An NMR study showed that the bond cleavage was specific to glycolic linkages, and the more hydrophilic glycolic segment degraded faster than the lactic acid part 160,162 . Likewise, the degradation of a sebacic acid copolymer was slower when raising the content of aromatic monomers to 100% 87 . For a homopolymer, low-MW and amorphous polymers were more susceptible to degradation than high-MW and crystalline.

Furthermore, during their degradation reaction, all degraded polymer NPs experience erosion, a process of material depletion/loss. Polymer erosion is characteristic of decreasing MW and other changes in its surface, morphology, and pH. Surface-eroding polymer NPs decrease their size with time, while bulk-eroding ones keep their size constant for a period until total disintegration happens.

As was discussed in part 1.2.2.6 about drug release mechanism, the degradation of polymer nanocarrier can influence or act as the main driving force of drug release. For instance, 67% release of insulin over a month from poly(CPP-SA) NPs was achieved by tuning the CPP:SA ratio (CPP = 1,3-bis-(p-carboxyphenoxy)propane, SA = sebacic acid). Knowing that poly(SA) insulin NPs released 100% insulin after three days, the prolonged release of insulin was because of the slow degradation of CPP polymer 163 . Similarly, the slow degradation of PLGA NPs also extended five times the drug release duration from DOX-PLGA conjugate NPs, compared to the DOX-loaded PLGA NPs 140 . As a result, one might play with the type of chemical bond between the drug and the polymer to increase or decrease the desired drug release duration.

In summary, polymeric DDS is a key player from which most fundamental theories on drug delivery processes, including drug release and carrier degradation, have been built. Polymer NPs naturally also have drawn the same attention to exploring their potential uses. The established polymer chemistry enables one to make desired polymers and tune their structure to suit drug loading and surface modification by either physical or chemical means. Moreover, biodegradable polymers are still one of the most used biomaterials for drug delivery because of their long-trusted biocompatibility. Research on polymer NPs is still ongoing to find better synthesis and formulation methods.

The following section introduces a relatively young material candidate for drug carriers: metal-organic framework.

Metal-organic framework nanoparticles 1.2.3.1 Basics about MOF and its applications for drug delivery

Metal-organic framework, or MOF, according to IUPAC recommendation 164 , is a "coordination network with organic ligands that contain potential voids". MOFs are coordination compounds between metal cations and organic ligands that extend in 1, 2, or 3 dimensions through repeating coordination units. The metals are mostly but not limited to multivalent metals, from transition metals like Zn, Cu, Fe, Cr, Ni, Co, and Ti to non-transition ones like Ca, Mg, and Al. The organic ligands must be chelating ligands that can coordinate with more than one metal site, such as polycarboxylate, polyphosphonate, pyridine, and imidazole 165 . A way to describe MOF structure is the association and arrangement of secondary building units (SBUs), a concept proposed by Férery 165 . One can find the systematic classification of MOFs based on their SBUs 166 . Nevertheless, there is no systematic nomenclature. MOFs are often named after the university that reports the structure and a number chosen by the inventor, for example, HKUST-1, where HKUST stands for Hong Kong University of Science and Technology.

The most crucial feature of MOF is its high porosity. MOFs are the most porous materials. Typically, the internal surface area of 1 gram MOF is 1500-5000 m 2 , for some of them up to 7000 m 2 , equal to the area of a football field 167,168 , whereas that of porous zeolites and silica 169 is only 100-1000 m 2 . MOFs can be microporous (pore size below 2 nm) or/and mesoporous (pore size between 2-50 nm). Depending on their structure and chemical composition, they can have varying thermal and aqueous stability. A MOF was reported to withstand up to 610 o C 170 .

The second attractive property of MOFs is the seemingly limitless choice and combination of metal ions and organic ligands. Their chemistry inherits from coordination, inorganic and organic chemistry, making it possible to modify MOF by linker exchange/substitution, functionalize and create new organic ligands. As a result, there are rich possibilities to tune MOF's chemical and physical properties as well as rationally incorporate a variety of new structural, magnetic, electrical, optical, and catalytic properties 171 .

MOF materials find their use in numerous applications related to sorption and separation of gas and liquid owing to their high permanent porosity 172 . MOFs can also contribute to tackling several social issues of environment and energy 173 , like serving as adsorbents for the removal of toxic chemicals 174,175 or as catalysts for energy conversion 176 .

The group of Férey pioneered the drug delivery applications of MOF. In 2006, they reported using MIL-100(Cr) crystals to encapsulate and release ibuprofen, an anti-inflammatory drug to treat pain and fever 177 . The following year witnessed the first publication of biocompatible nanosized MOFs and their ability to load several anti-cancer drugs and imaging 178 , which ignited the research direction of MOF nanoparticles (nanoMOF) for drug delivery. MOF has several advantages compared to other existing DDSs. First, its unique high porosity allows MOF to absorb a high amount of drug per carrier, leading to high drug payload and loading efficiency 103,177 . In addition, MOF having both hydrophilic metal clusters and hydrophobic organic ligands can interact efficiently, thus well incorporating drug molecules with different physiochemical properties, from hydrophilic to hydrophobic APIs 178 , from small 179 to macromolecules 180 . The possibility of using both organic chemistry and coordination chemistry offers versatile methods for drug loading, modification, and functionalization. Moreover, MOF porosity enables separate compartments in the core and at the surface. As the drug is protected in the MOF pore, surface decoration with targeting and stabilizing agents does not interfere with drug release and encapsulation, resulting in more effective protection of the active substance and easier control of surface modification and drug loading 181 . Finally, the ability to add other properties like optical or radiation enhancers makes it possible to fabricate MOF for combined therapy 182,183 .

MOFs have been researched for the treatment of various diseases, including cancer, infection, etc. 66,184 . One clinical trial is known from the Lin's team using MOF for cancer radiotherapy 185 .

Choice of MOF

Biocompatibility and stability in water are the two principal criteria for selecting MOF for drug carriers. According to the discussion in part 1.1.2., MOF should be made of non-toxic metal and organic ligands, with caution on the median lethal dose and the intended administration dose 103,186 . Endogenous metals and organic ligands are preferred. Apparently, the toxicity of a MOF does not depend only on the individual toxicity of its components. Biocompatibility studies in vitro and in vivo of several MOFs can be consulted in the reviews of Mendes and Ettlinger 184,186 .

Various MOFs display different degrees of stability in water vapor and water liquid [START_REF] Canivet | Water adsorption in MOFs: fundamentals and applications[END_REF]187,188 . It took a few years for researchers to find the right affecting factors. Some of them are related to MOF structure 189 ; for instance, a shorter and more rigid linker makes a MOF more stable. MOF with large cages and pores is more prone to degradation issues than one with small cages. Also, the form of inorganic SBUs and their arrangement have an effect; MOF made of isolated metal clusters like UiO-66(Zr) and MIL-88(Fe) are more reactive than those built from infinite metal chains like MIL-140(Zr) and MIL-53(Fe). In addition, the nature of metal cations can impact MOF stability in several ways. MOF chemical reactivity can be estimated from the metal cations' reactivity towards metal-ligand exchange 189 . For instance, given the same organic ligand, Cr(III) MOF is more inert than Fe(III) due to a slower metal-ligand exchange constant from M 3+ to M(H2O)6 3+ (M = metal). For a given charge, a smaller metal cation is less reactive. The high stability of MOF is also usually attributed to the nuclearity of its inorganic building unit 186 , illustrated by UiO-66 190 , NU-1000 191 , and PCN-222 192 , composed of the multi-nuclear Zr6 metal cluster. An even more substantial basis for choosing MOF stable in water is the metal-ligand coordination bond strength 186 , which can be rationalized by the principle of hard and soft acids and bases 193 . It was shown that MOF-5 and MOF-177 decompose quickly in water due to their weak metal-oxygen coordination bond 194,195 . Also, ligands with higher pKa values than the water pH increase MOF aqueous stability 196 .

The most studied MOFs for drug delivery are namely MIL-100(Fe), ZIF-8(Zn), and UiO-66(Zr). Their ligands, SBUs, and cages are demonstrated in Figure 1-15. MIL-100(Fe) (MIL = Matériaux de l'Institut Lavoisier de Versailles) consists of Fe(III) and benzene-1,3,5tricarboxylate, commonly called trimesate 197 . The SBU is oxo-center-trimers (Fe3O) of iron 6coordinated with the trimesate carboxylate's oxygen. There are two kinds of pores, small mesoporous cages (25 Å), accessible exclusively through pentagonal opening windows (4.8-5.5 Å), and large mesoporous cages (29 Å) having both hexagonal (8.6 Å) and pentagonal windows. MIL-100(Fe) is famous because it is the first MOF to have pores and pore windows large enough to encapsulate efficiently drug molecules 177,178 . Its stability in water up to 50 o C was reported 198 . ZIF-8(Zn) (ZIF = Zeolitic imidazolate frameworks) is made of Zn(II) and imidazolate 199 . The SBU is a zinc ion tetrahedrally linked to four 2-methyl imidazole groups. ZIF-8(Zn) with (sod) topology has pores of 11.6 Å diameter with hexagonal pore windows of 3.4 Å. ZIF-8(Zn) has been studied as a DDS for its pH-sensitive reactivity 200 and its friendly synthesis that allows in situ drug loading 201 . These features compensate for the small pore window that could prevent drug molecules from entering ZIF-8(Zn)'s cage. ZIF-8 is also inert to water 198 . UiO-66(Zr) (UiO = Universitetet i Oslo) is made of Zr(IV) and 1,4-benzene dicarboxylate, commonly called terephthalate linker 190 . The SBU comprises a Zr6O4(OH)4 cluster in which each zirconium is 8-coordinated with terephthalate linkers. There are two cages: a tetrahedron cage of 7.5 Å and an octahedron cage of 12 Å; both have a triangular pore window of 6 Å 190,202 . UiO-66(Zr) is known for its exceedingly high 12 coordination number and high stability in water 203 , which combines with high porosity and sufficiently large pore window to make UiO-66 a good candidate for nanoMOF drug carrier 204,205 . Its ability to withstand defects creates additional possibilities for drug loading and functionalization 206 . A worth-mentioning MOF family for drug carriers is bioMOF, whose organic ligands are biomolecules or the intended drug, alleviating the toxicity concern of synthetic ligands. The first published examples are bio-MOF-1 and Bio-MIL-1. Bio-MOF-1 208 is composed of Zn(II) and adenine, one of the four nucleobases in DNA. BioMIL-1 209 is made of mixed Fe(II)/Fe(III) and nicotinic acid, a compound used to treat vitamin B3 deficiency 210 . The group of Rosseinsky synthesized bioMOFs using highly symmetric biomolecules or a symmetric coligand. Finally, several honorable mentions of MOF-based DDSs are the flexible MIL-53(Fe) 211,212 and the natural-origin CD-MOF 213,214 .

NanoMOF synthesis

In contrast to polymer NPs preparation, MOF NPs must be synthesized from precursors. Different sources can supply the energy needed to drive MOF nucleation and crystal growth, hence different techniques for MOF synthesis, including solvothermal, electrochemical, microwave, and mechanochemistry. They all have the potential for large-scale production of MOF 215 . Solvothermal, electrochemical, and microwave synthesis can be scaled up using flow chemistry configuration.

In controlled drug delivery, a key criterion for optimizing MOF synthesis is precise control of nanoMOF's size and size distribution. Moreover, toxic solvents and reacting agents should be avoided, like dimethylformamide (DMF), a solvent frequently used in MOF synthesis 103 . Alternative solvents and schemes for the green synthesis of MOF are available 216,217 . No matter the synthesis method, activation is always crucial to remove trapped solvent and excess linkers molecules in the framework to make the MOF's permanent porosity accessible. MOF's high porosity is directly related to its ability to absorb drugs in the subsequent process. Commonly used activation methods are vacuum drying, solvent exchange, and supercritical drying 218 .

In the solvothermal/hydrothermal synthesis technique, reactions occur in closed vessels under autogenous pressure and at a temperature above the solvent's boiling point. It is called hydrothermal when the solvent is water. One can play with reaction temperature, time, solvent, nature of precursor, reagent concentration, and pH 218 . Synthesis of nano-sized MOF has been realized by reducing the reaction time, temperature, or pressure to limit the crystal growth 103 . However, heterogeneous heating supplied by convection by the oven causes low synthesis yield and high polydispersity of nanoMOF size. Thus, growth inhibitors such as acetic acid, pyridine, and polyvinylpyrrolidone have been used. NPs smaller than 100 nm from several MOFs have been obtained 103 ; however, long reaction time is a limitation.

Microwave-assisted synthesis is an alternative to the solvothermal technique, as microwave irradiation produces more homogenous heating, resulting in homogenous nucleation and growth of NPs. The technique relies on the interaction between electromagnetic energy of microwave frequency and polar solvent molecules/ions in solution/ions in a solid 219 . Direct and homogenous irradiation of the reaction mixture results in efficient and rapid heating, promoting localized spots for MOF nucleation and crystal growth. Monodispersed nanoMOF can be obtained in under an hour. Several nanoMOFs potentially being used for drug delivery were obtained by this method 103,215 ; the obtained NP dimension is up to 200 nm. Primary reaction conditions are temperature, power level, heating time and rate, and reactant concentration 220 .

Electrochemical synthesis uses redox potential to cause MOF formation. The metal ions formed from anodic dissolution are continuously introduced to a mixture with organic linkers and electrolytes in an electrochemical cell 215 . No metal salts are needed, and no residual anions are produced 215 . One can optimize reaction time, organic linker concentration, and temperature 215 . Published electrochemical syntheses that may be relevant to drug delivery applications can be found for ZIF In mechanochemical synthesis, mechanic breakage of reactants' intramolecular bond initiates chemical transformation 219 . It has received attention for MOF synthesis for limited or no use of solvent, room temperature condition, and short reaction time. The metal source can be pure metal, oxide, hydroxide, or salts 221 . Typical methods are ball milling, grinding, extrusion, and heating. Note that the product purification may still require a solvent 215 . Ball milling and manual grinding have been used to prepare drug-loaded MIL-100(Fe) 222 and surface-coated ZIF-8(Zn) NPs 223 .

Drug loading

An outstanding feature of nanoMOF-based DDSs is that they can absorb both liquid and gas. Methods for loading and delivery of therapeutic gases can be found elsewhere 103,224 .

The loading methods present in the following are for active substances soluble in liquid solvents.

Drug loading can be done after or during the nanoMOF synthesis process, i.e., postsynthetic and one-pot methods 225,226 . In the one-pot style, nanoMOF synthesis and drug incorporation are carried out in situ. The API is mixed with the synthesis precursors and becomes integrated within the resulting nanoMOF 66 . In the post-synthetic fashion, drug loading is done after nanoMOF synthesis and activation steps. Usually, nanoMOF is soaked in a drug solution, and drug molecules are absorbed and trapped in its pores. The solvent is mostly water to be safe for drug delivery applications.

From the operation point of view, one-pot drug loading is easier to be done on a large scale than a multi-step process. It is a simple procedure and requires a short manipulation time. One-step encapsulation of caffeine in nanoZIF-8(Zn) was complete after 2 h, while it took three days for the post-synthetic method to achieve the same caffeine payload 227 . Moreover, this in situ drug loading method offers a uniform distribution of APIs throughout NPs 228 . It is particularly suited for nanoMOF with small pore windows; the drug can still be encapsulated in MOF's cages without entering its pore opening. A typical example is ZIF-8(Zn), which has a relatively small window of 3.4 Å diameter. Successful loading of anti-cancer drugs such as doxorubicin (DOX) 228 , camptothecin 229 , and 3-methyladenine 230 was reported. Simultaneous loading of two drugs was also possible and obtained around 40 wt% total payload 231 . Nevertheless, it is not straightforward to generalize this method to other MOFs and drugs. Firstly, since the drug must not be altered or degraded, mild MOF synthesis conditions, including water solvent, room temperature, and atmospheric pressure, are required, which imposes constraints on the synthesis procedure. Secondly, the success of drug integration within MOF depends on the drug's properties and interaction with the MOF's ligand. Specifically, Paseta et al. explained that the active substance should have compatible solubility with MOF ligand, hence strong affinity 232 . It should exhibit better interaction with the ligand than the solvent, thus being able to compete with the solvent adsorption and prevent solvent trapping inside MOF.

Another attractive in situ pore encapsulation technique is using mechanochemical synthesis, which eliminates the use of solvent 66 . Apart from being convenient and promising for industrial applications 233,234 , the solvent-free drug loading method avoids the competition between drug and solvent guest molecules for the MOF's metallic binding sites, thus increasing drug payload. Studies using ball milling or grinding of a mixture of metal salts, organic ligands, and active substances have attained an acceptable payload of 15-20 wt% 235,236 .

Biomimetic mineralization, a relatively new method, is helpful for incorporating therapeutic biomacromolecules like protein and nucleic acids since they are too big for most of the existing MOFs. The difference from a typical one-pot drug loading is that the biomolecules act as the nucleation site for nanoMOF crystallization 180 . The first proof-of-concept successfully applied this method for a series of biomolecules (protein, enzyme, DNA) and nanoMOFs 237 . It claimed that enzymes encapsulated in MOF maintained their activity after heating at 80 o C and boiled in DMF at 153 o C. It was proposed that specific interaction between biomolecule and MOF facilitates MOF formation and growth around the biomolecule. Furthermore, the Gassensmith team managed to encapsulate a whole virus inside ZIF-8(Zn) 238 and studied the growth mechanism of ZIF-8(Zn) on virus 239 , showing different configurations of ZIF-8 depending on the interaction between the Zn 2+ precursors and the virus 239 . A significant consideration of biomimetic mineralization is that the guest molecules are too large compared to the MOF host to allow diffusion out. Therefore, drug release, in this case, needs to rely on the MOF carrier degradation 240,241 .

The synthesis of bioMOF, whose ligands are active molecules, while not being a poreencapsulation method, still fits in the category of one-step drug loading. There is no compulsory need for high porosity, large cage, or large pore window size, and the toxicity concern is also eased. In this case, MOF degradation is needed to allow the release of ligands, i.e., the drugs. Long and extremely long drug release duration due to the slow degradation of several bioMOFs were reported: 70 days for bioMIL-5 made of Zn(II) and azelaic acid in MHCA (Mueller-Hinton cation adjusted) broth 242 , and 29 days for MOF made of Ca(II)/Mg(II) and bisphosphonate drugs in acidic aqueous medium 243 The post-synthetic pore-encapsulation method is more prevalent in the literature, owing to its facile applicability for a variety of drugs and nanoMOF carriers 66 . However, the loaded drugs are not necessarily distributed uniformly throughout nanoMOF as the pores closer to the NP surface are more kinetically accessible than those in the bulk region. Pore encapsulation of drugs in the pre-formed nanoMOF is essentially the adsorption of drugs, which is highly dependent on the host-guest interaction between the nanoMOF carrier and the drug. As done with polymer NPs, the drug-carrier interactions in nanoMOF can be classified into non-covalent and covalent types. Non-covalent interactions allow the physical adsorption of drugs, including van der Waals, hydrogen bonding, and electrostatic interaction. The van der Waals is a universal interaction that can be used for any drugs and MOFs. Peng et al. tested the hypothesis of five possible interactions and bonds, including coordination, electrostatic forces, hydrogen bonds, π-π interactions, and van der Waals interaction. They proved that it was the van der Waals that enabled the successful loading of single-stranded DNA in IRMOF-74-I(Ni) 244 . Hydrogen bonding is also useful for MOF-based DDSs because many of them consist of hydroxyl groups, such as the interaction between ibuprofen's carboxylic and MIL-53(Fe)'s hydroxyl 212 or that between caffein's carbonyl and UiO-66's hydroxyl 245 . The same consideration can be applied to π-π interaction, popular MOF nanocarriers such as MIL-100(Fe) and UiO-66(Zr) have ligands containing aromatic rings. The existence of π-π interaction was demonstrated by theoretical calculation and modeling in the cases of 5-fluorouracil (5-FU) and UiO-66(Zr) and of DOX and nanoMIL-100(Fe) 246,247 .

Knowing the available interactions between the drugs and MOF carriers helps to design better MOFs for drug loading. An illustration is the research of the Sun group on the loading of 5-FU in a chiral MOF 248 . Their first attempt reported burst 5-FU release due to their MOF degradation. They tried to strengthen drug-MOF interaction by synthesizing another MOF with many sites for hydrogen bonding. As a result, they obtained a higher drug payload and a more progressive drug release 249 . In another study, after synthesizing MOF-867, the authors modified the organic ligand's bipyridyl group to N + -CH3 to introduce positively charged sites in the MOF structure 250 . Efficient electrostatic interaction between this cationic MOF-867 and anionic-nature diclofenac sodium leads to a high payload of 54.6 wt%. The corresponding neutral MOF did not absorb at all, even after seven days of being soaked in the drug solution. Nevertheless, non-covalent bonding is considered weak, which raises the concern that subsequent processing of drug-non-covalently-bonded-MOF, such as surface modification, can cause premature drug release 226 .

An alternative method is using more robust covalent bonds 226 . A prodrug or MOF-drug conjugate (if one takes the terms from polymer-based DDS) is created, where the drug is covalently attached to MOF through its functional group or SBU's metal 251 . This covalent drug-MOF bond should be stable, but the drug must be cleavable under specific conditions. A limitation is the increasing possibility of non-uniform drug distribution within MOF. As covalent bonds are stronger than non-covalent ones, a molecule already covalently bonded to the pore near or at the nanoMOF surface will not likely diffuse deep in the NP's bulk. To conjugate the drug to the organic linker, one usually needs to perform post-synthetic modification of the organic linker. This idea was realized for nanoMIL-101(Fe) 252 . A prodrug of anti-cancer cisplatin, ESCP (ethoxysuccinato-cisplatin), reacted with the -NH2 functional group of the MOF's terephthalate ligand (BDC). However, the obtained drug loading efficiency was low; roughly 40% of NH2-BDC was converted to ESCP-NH-BDC. Coordination between drugs and the MOF's metal cations is more common in MOF research, such as between UiO-66(Zr) and the drug's phosphate group 205 or between Mg2(olsalazine) and olsalazine's amine group 253 . The active substance can be modified to take advantage of the metal-drug coordination, like in the study loading carboxylate-functionalized BODIPY in nanoUiO-66(Zr) 206 . The functionalized BODIPY even replaced the constituting terephthalate ligand of UiO-66 without collapsing the framework, using a solvent-assisted ligand exchange method. A similar strategy was employed for nanoMIL-100(Fe). When functionalizing API with phosphate groups, high drug payload and high efficiency were attained 254,255 , which is impossible with non-modified API. Co-loading of two phosphate drugs was successful 256 .

Post-synthetic surface attachment of drug is quite rare in nanoMOF studies. It offers another possibility for drug loading of MOF with a small pore window but provides weak protection for the loaded cargo exposed to the outer environment 180 . It is generally expected that attaching the drug to the NP surface is likely to cause burst drug release, but some researchers showed that DOX strongly complexed to ZIF-8(Zn) surface did not suffer from rapid release 201 .

All the above-mentioned methods demonstrate the usefulness of MOF-based DDSs as there are various methods to incorporate and protect an API, owing to the diverse chemistry of MOF.

In principle, drug loading capacity in nanoMOF depends on the API's physiochemical properties, namely the size, shape, and hydrophilicity. These properties guide one to select the most suitable MOF for a given API; this was done to screen a series of MOFs for salicylic acid encapsulation 257 .

Several MOF's structural parameters important for MOF's drug loading capacity are surface area, pore size, and pore opening size 225 . The more valuable parameters the MOF possesses, the higher possibility of achieving good drug payload and encapsulation efficiency. Generally, MOFs possessing high internal surface area are more likely to have high drug payload 177,258 . A similar argument applies to pore size. The larger the pore diameter, the larger the empty void volume, and the higher the amount of drug nanoMOF can accommodate. MIL-100(Cr) with smaller cages adsorbs less ibuprofen than MIL-101(Cr) 177 .

The pore size difference can be used to load multiple active substances. Loading of the large enzyme in the large cage was followed by loading of smaller enzyme in the smaller cages, achieving an experimental payload matching the theoretical maximum loading capacity 259 (Figure 1-16A). Likewise, the diameter of the pore opening window decides the drug's accessibility to the MOF carrier. If the window is too small, the drug cannot enter. If the pore is too big, the drug can enter but easily get out, which might lead to uncontrolled release. The small size of the opening windows explained the lower payload of several compounds in MIL-53 and MIL-88A than in MIL-101-NH2 and MIL-100 178 . Another work designed a MOF to have a large enough pore window through which single-stranded DNA can pass 244 . The team of Gref demonstrated the compartmentalized encapsulation of two antibiotics in small cages and large cages of nanoMIL-100(Fe) 260 (Figure 1-16B). Soaking MOF in amoxicillin which can pass only the hexagonal window 9 Å, was done first, followed by soaking in potassium clavulanate which can pass both the hexagonal and pentagonal windows 9 Å and 5 Å, respectively. Moreover, tuning how drugs and nanoMOF carriers interact with each other is another way to obtain efficient drug loading. A study by Gaudin et al. determined a quantitative relationship between the MOF structure and hydrophobic caffein uptake 261 . They observed that it increased when MIL-88B(Fe) has polar and hydrogen-bond donating functionalized groups like -NH2 and -OH, which was explained by the increase of hydrogen bonding between caffein and functionalized MIL-88B(Fe). Notably, an opposite trend was observed when loading the same compound in UiO-66(Zr) 204 because different solvents were used, and these two studied MOFs have different hydrophobicity. The authors' follow-up article showed that caffein does not interact with UiO-66(Zr) through the functional groups -NH2 and -OH 245 , thus explaining the poor caffein uptake observed previously 204 . This study highlighted the importance of determining the anchoring sites for drugs inside MOF to optimize the proper parameters of drug loading formulation and MOF carrier design.

Regarding the drug loading method based on coordination linkage, the complexing affinity of the guest molecules towards the MOF's metal cations, in comparison to that of MOF's organic ligands, are important factors. It was demonstrated after the loading of azidothymidine-triphosphate (AZT-TP) in nanoMIL-100(Fe) up to 9% of this MOF's constituting ligands, trimesate were leached, which can be rationalized by a stronger complexing strength of triphosphate towards Fe(III) than carboxylate 262 . The azidothymidine-monophosphate (AZT-MP), however, caused only 2% of trimesate leaching, even though the payload of AZT-MP (36 wt%) was higher than AZT-TP (24 wt%), suggesting that the triphosphate derivative exhibits stronger interaction with nanoMIL-100(Fe) than the monophosphate one. Souza et al. presented that one-pot mechanical grinding is preferred to the post-synthetic soaking for loading 5-FU in HKUST-1 because the former yields selectively pore encapsulation while the latter could lead to surface adsorption, which associated with potential burst drug release 222

The location of the loaded drug also receives attention, for it can affect the drug release process. The drug located near the nanoMOF surface can diffuse or be cleaved from the framework faster than the drug loaded deep in the MOF particles' bulk. This behavior is especially critical for post-synthetic drug loading, in which the drug might favorably locate in the MOF's pores close to the surface. A work on RNA delivery proposed that siRNA was loaded inside the MOF's pores because the RNA-loaded-MOF has a lower internal surface area than the empty one 263 . Porcino et al. used solid-state NMR spectroscopy to prove that adenosine triphosphate was exclusively coordinated to Al(III) in the bulk nanoMIL-100(Al) 264 . More advanced techniques like scanning near-field optical microscopy and fluorescent imaging were also used to verify the guest's location inside the pore of ZIF-8(Zn) 265 .

Surface modification

Modifying nanoMOF's outer surface can improve its colloidal and chemical stability, prolong drug release kinetics, prevent burst effect, increase blood circulation time, and add targeting function 266 . Apart from the criteria for a DDS's surface coating, that of a porous material like MOF should not intrude into it or block its pores, preventing drug release. To minimize this interference, surface modification is performed after the drug loading step in a typical workflow of nanoMOF formulation for drug delivery.

Three main surface modification methods have been used for nanoMOF drug carriers: noncovalent, covalent, and core-shell methods 266 . Compared to the first two methods, the coreshell method creates a dense continuous coating surrounding the MOF core. The other two methods usually involve the attachment of individual molecules to nanoMOF surface with varying coverage density by utilizing non-covalent or covalent interaction between nanoMOF and the coating. Most examples of the non-covalent route are coating natural substances or biomolecules, probably owing to the mild synthesis condition. Namely, electrostatic interaction was used to attach the lipid bilayer to nanoMIL-88A(Fe) to create an exosome vesicle surrounding MOF 267 . Exosomes are endogenous vesicles with lipid bilayer membrane. This method circumvents the challenge of loading drugs in exosomes. No drug leakage up to 10 h and no drug release in the artificial lysosomal fluid was observed. The hypothesis is that MOF decomposed under acidic pH, leading to an increase in osmotic pressure inside the exosome and, in turn, the destruction of its bilayer 267 . In addition, π-π interaction is commonly used because of the usual presence of aromatic rings in MOF's ligands. For example, oligonucleotides were adsorbed on nanoMIL-101(Fe) through its π-π interaction with the terephthalic acid organic ligand 268 . A functionalized polymer pyrene-PGA-PEG was grafted on UiO-66(Zr) surface, taking advantage of the π-π interaction between the polymer coating's pyrene and the MOF's terephthalate and benzoate 269 .

The covalent surface modification method is more reported, probably because it offers more strongly bounded coatings for a DDS. Chemical reactions used are condensation, click chemistry, and coordination 266 . For instance, condensation was behind the coating NH2-PEG of nanoMIL100(Fe) 270 , where the polymer's terminal amino group reacts with the MOF's surface carboxylate sites to form an amide bond. Both colloidal stability and cellular uptake were ameliorated. The same amide coupling was used to create the poly(N-isopropyl acrylamide) (PNIPAM) coating for UiO-66-NH2 271 . In another work, DNA was grafted on UiO-66(Zr) surface using the click catalyst-free strained-promoted alkyne addition reaction 272 between the DNA's dibenzylcyclooctane and UiO-66's azide group, both of whom were rationally designed for the pristine DNA and nanoMOF 273 . Surface modification through coordination is deemed the most generalizable method and has been demonstrated for many types of grafted molecules. Folic acid, a small molecule, was attached to two zirconium MOF NPs through coordination between the folic acid's carboxylate and the MOF's Zr surface sites 274,275 . The two studies both obtained reducing NPs aggregation in water but observed mixed results on stability in phosphate buffer. Coordination bonding is also possible for coating natural and synthetic polymers. Heparin, a biopolymer with anticoagulant properties, was coated on nanoMIL-100(Fe) through the coordination between heparin's sulfate group and Fe(III) sites 276 . The authors improved the MOF's colloidal stability in water and phosphate buffer spiked with albumin (a blood protein) and prolonged drug release kinetics while maintaining the nanoMOF's drug loading capacity 276 . Coordination bonding was also used for different derivatives of cyclodextrin (CD), including native form 277 , CD with different lengths 278 , and phosphate-functionalized 279 , relying on coordination between metal cations and hydroxyl or phosphate groups. The phosphatemetal coordination has been proved efficient for other families of MOF based on zirconium; it was used to make phosphate-terminated-DNA and siRNA coating 280,281 .

A particular effort was given to PEG coating, as it is considered one of the most reliable strategies to prepare long-circulating NPs, i.e., "stealth" NPs 73 . PEG can protect NPs against protein adsorption by electrostatic repulsion and steric hindrance 72 . Surface grafting of PEG chains has been implemented for several MOF families, including UiO-66 282 , MIL-88A 178 , MIL-100 178 , and ZIF-8 231 . PEG was functionalized with amino, carboxylate, or alkyne groups to promote coordination or click chemistry with nanoMOF. In the case of UiO-66(Zr), the PEG coating significantly prolonged the drug release at pH 7.4 compared to the uncoated MOF 282 . In contrast, the PEG-coated nanoMIL-100(Fe) experienced unwanted drug release and had lower drug loading capacity than the uncoated one 178 , hypothetically because some of the PEG's flexible polymer chains can penetrate MOF pores and trigger drug leakage. The stealth effect of PEG was still effective; the PEGylated MIL-100(Fe) accumulated less in the liver and spleen than in the bare-surface one. To avoid pore inclusion, other authors connected PEG to nanoMIL-100(Fe) through bulky cyclodextrin (CD) molecules, which cannot pass the MOF pore window 279 . They proposed different designs for linkages between PEG and CD 279,283,284 . Another strategy is to use branched polymers that connect to the nanoMOF surface and PEG through their branches, serving as an interface between them. As an example, Cutrone et al. functionalized dextran with alendronate and PEG, where alendronate serve as anchoring sites for MIL-100(Fe) 285 .

Overall, all reported surface modification methods have been demonstrated to improve, to different extents, the colloidal stability of nanoMOF in water and physiological media, presumably by increasing hydrophilicity and enhancing the steric/electrostatic repulsion between NPs. Changes in colloidal stability of the coated nanoMOFs are mainly monitored by measuring their size and zeta potential (ZP), which is an indication of surface charge 62 .

For instance, coating cancer cell membranes on a Zr MOF leads to a slight increase in particle size and a significant change of ZP from positive to negative values +24.5 mV to -28.5 mV 286 . It is known that the ZP of nanoMOF changes from positive to negative when increasing the pH to basic, and the free terminal carboxylate group is protonated. However, a study reported that the lipid-coated nanoMIL-100(Fe) has the same ZP no matter the pH, indicating that the lipid layer shields the nanoMOF's charged surface. The formulation of surface coating also plays a role. Cutrone et al. observed that lowering the mass ratio of dextran coating over nanoMOF increases the level of nanoMOF aggregation in water 285 . The group of Fairen-Jimenez recently disclosed the surface modification method with methoxyl-PEG-PO3 using metal-phosphate coordination that allows one to redisperse freeze-dried nanoMOF in water, which has never been done before 287 (Figure 1 -17). The zeta potential of PCN-222(Zr) reversed the sign after the surface modification. The authors also found evidence for the partial inclusion of PEG in the MOF pore, which reduced 1-5 wt% of drug payload, but they argued that the dispersibility of lyophilized NPs is worth this inconvenience.

Furthermore, surface coating is a tool to control the chemical stability of nanoMOF, consequently influencing the drug release kinetics. The coating acts as a shield to protect from the outer environment; therefore, the stability of the coating is vital to ensure the level of nanoMOF protection. For instance, unwanted detachment of heparin over time results in the same UiO-66(Zr) degradation rate in phosphate compared to the bare-surface one 274 . In most cases, the authors attempt to make the cases for their surface modification methods by showing that the surface coating slowed down the nanoMOF degradation and drug release compared to the non-coated one 269 .

LázaroI and coworkers compared three different surface modification protocols (coordination modulation, post-synthetic ligand exchange, click modulation) and six coating molecules (folic acid, biotin, heparin, poly-l-Lactide, PEG, PNIPAM). They found some relation between the choices of coating method and the stability, therapeutic activity, and toxicity of UiO-66(Zr) 274 . Firstly, all surface functionalized-nanoMOFs have better dispersion in phosphate buffer than the un-coated ones, except for one biotin-modified due to low surface coverage. For the same grafting molecule, the coordination modulation (CM) folatecoated nanoMOF kills more HeLa cancer cells than the post-synthetic (PS) folate-coated one, despite the first having lower drug content than the latter. The CM-coated has folate both at NPs' surface and in bulk (folic acid is a small molecule that can penetrate UiO-66 pore), whereas PS-coated has mainly at the surface. Those two coated MOFs have similar degradation rates; thus, folic acid content is removed more slowly in the CM-coated. Also, it was found that the presence of folic acid promotes MOF uptake by HeLa cells by endocytosis. Altogether, the CM-coated nanoMOF displays a greater therapeutic effect on cancer cells, as it retains higher content of folic acid over time. As seen in the discussion of drug loading, the coating stability also depends on the bonding strength between the coating substances and nanoMOF. In research on coordination methods for surface grafting of oligonucleotides, a linear correlation was found between the obtained surface coverage of phosphate-functionalized DNA and the surface density of metal nodes 288 . It also discovered that the higher the MOF metal cation's coordination number, the stronger the metal-phosphate coordination bond and the higher the DNA surface coverage. Likewise, another group that studied phosphate-CD coating observed that the coating stability is directly related to the density of the phosphate group on CD 289 . XPS and solid NMR spectroscopy provided proofs for the M-O-P metal-phosphate coordination bond 264 .

Drug release

The driving forces of the drug release process from MOF are mainly diffusion and MOF degradation 290 . To control the drug release process in MOF, one can play with: i) the drug-MOF interaction through MOF functionalization, ii) drug diffusion from MOF pores through alternating pore size, MOF morphology, iii) kinetics of MOF degradation through surface modification 103 .

Many studies have pointed out the relationship between drug release kinetics and the hostguest interaction between MOF and loaded drugs. A mathematical model to predict guest release based on host-guest and guest-guest interaction was established 291 . The three-stage release of ibuprofen from MIL-100 and MIL-101(Cr) was explained according to different absorption sites of ibuprofen within the pores, close to the pore wall or mobile inside pore 177 . Likewise, the 2-maxima release profile from MIL-100(Fe) film was attributed to molecules physisorbed and chemisorbed 292 . Moreover, the type of nanoMOF carrier-drug bonding dictates the drug release mechanism. Drug-loaded MOFs using electrostatic interaction undergo drug release in phosphate buffer saline (PBS) due to ion-exchange between the anionic drug and the phosphate anions 208,293 . Drug-loaded MOFs using coordination bond experience MOF degradation in the same PBS medium. The stronger the coordination bond, the slower the drug release. It was discovered that monophosphate drug was released faster from MOF than triphosphate drug because the former has weaker interaction with nanoMOF 262 .

When diffusion through the pore is the release rate-controlling step, the role of MOF structure becomes increasingly important. For example, drugs residing at nanoMOF surface, thus having less protection from MOF, will likely experience burst drug release 178,248 . Attempting to extend drug release time, Fairen-Jimenez and coworkers have proposed a strategy to hinder drug diffusion from MOF pores by mechanical and thermal treatment causing partial pore collapse 294,295 . The release of the drug calcein from UiO-66 was prolonged from 2 to more than 30 days 294 . Likewise, an exceptionally long release of ibuprofen from the flexible MIL-53(Fe) and (Cr) was because of its being strongly confined in MIL-53's small pore 212 . A comparative study was carried out on release in water at 37 o C of a hydrophilic and a hydrophobic active substance from three different MOFs, namely MIL-100(Fe), MIL-127(Fe), and UiO-66(Zr) 296 . All those three MOFs are stable under this drug release condition; less than 2% of ligand release was detected. Therefore, the main driving force of drug release should be diffusion. It was shown that large cavities and pore windows of MIL-100(Fe) facilitate drug diffusion from its pores, no matter the hydrophobicity nature of the drug. Conversely, the nature of the drug strongly influences drug release in the case of UiO-66(Zr), from which hydrophilic aspirin is released faster than hydrophobic ibuprofen. Probably because UiO-66(Zr) is more hydrophobic than the other two MOFs, its drug release is strongly linked to the loaded drug's hydrophobicity. Other authors demonstrated that prednisolone which has very low water solubility, could not be fully released (100%) in aqueous media. Adding ethanol to the release medium helped raise the 1-day release from 60% to 75% 297 .

For biodegradable MOF, drug release is generally driven by nanoMOF carrier degradation.

As discussed in the next section, MOF degradation reaction varies in different media, and so does the drug release kinetics. For instance, the study of AZT-TP-loaded nanoMIL-100(Fe) detected a significant portion of drug release and ligand release in PBS but found no drug release in Tris buffer (Tris(hydroxymethyl)aminomethane) or water 255 . Another work on alendronate-loaded UiO-66(Zr) observed slower drug release in PBS at pH 5.5 than at pH 7.4, hypothetically because of slower degradation of this MOF at acidic pH, but no proof on ligand release was given 205 . Meanwhile, ZIF-8 degradation at acidic pH makes it attractive for targeting cancer cells where extracellular microenvironments are more acidic (pH 5.7-7.8) than healthy tissues 229 .

To investigate the release-rate controlling process, one should always check the release of MOF's constituting ligands during drug release. The drug release processes from four different MOFs were studied 298 . The researchers discovered that in water where there is no MOF degradation, drug release is governed by drug pore diffusion and drug-MOF interaction. Whereas in PBS, the drug release kinetics overlap with the linker release, indicating that drug release is governed by MOF degradation. Nevertheless, it is not always straightforward to rationalize the drug release mechanism. An investigation on a series of bMOFs synthesized from zinc-adeninate SBUs and dicarboxylic ligands and then loaded with cationic drugs by electrostatic integration showed an unusually prolonged drug release duration (49-80 days) 299 . Initial burst release during the first 4 hours was found for two MOFs, bMOF-100 and bMOF-102, due to their immediate MOF degradation after exposure to simulated body fluid (SBF). Then, the drug release became very slow, probably because the drug was trapped in the collapsed structure (one can relate to the amorphization treatment mentioned above 295 ). Drug release from bMOF-1, another MOF in the series that is stable in the drug release medium, was found to depend on the cation exchange process. Despite those observations, the main reason behind this prolonged release phenomenon has not been explicitly explained yet.

In addition, some specific drug-drug interactions can impact the drug release mechanism.

The two studies on topotecan and doxorubicin (DOX) encapsulation in nanoMIL-100(Fe) found an inverse relationship between drug payload and drug release %, which is somewhat counterintuitive 300,301 . Low drug release % was detected at high drug loading wt% and vice versa. Both these two active molecules have intermolecular π-π stacking among themselves, resulting in their aggregation at high concentrations. At high drug payload, DOX molecules aggregate inside nanoMOF's pores, consequently stabilizing the MOF matrix against degradation, possibly by blocking reacting agents from penetrating the pores. Therefore, the rate of MOF degradation, hence the rate of drug release, decreased.

NanoMOF degradation

In general, the coordinative nature of the metal-ligand bond makes nanoMOF prone to (bio)degradation 103,186 . Its stability and destruction principally involve the competition at the metal center between the constituting ligands and the external environment's coordinating species 189 . In water, two main MOF degradation mechanisms are hydrolysis and ligand replacement [START_REF] Canivet | Water adsorption in MOFs: fundamentals and applications[END_REF] .

The ligand replacement involves the attack of water, resulting in metal coordination with water molecules and the formation of a free ligand [START_REF] Canivet | Water adsorption in MOFs: fundamentals and applications[END_REF] :

M n+ -L n-+ H2O M n+ -(OH2) + L n-
Water hydrolysis involves the breaking of metal-ligand bonds and dissociation of water, producing hydroxylated cations and free protonated ligands:

M n+ -L n-+ H2O M n+ -(OH) -+ HL (n-1)-
For MOFs composed of carboxylate ligands, it is generally agreed that their degradation reactions in basic conditions undergo the ligand replacement mechanism 189 . In strongly acidic conditions with pH lower than pKa of ligand, protonation of ligand happens, causing dissociation of coordination bonds and MOF degradation. For example, incubation of a zinctartrate MOF in SBF leads to the release of Zn 2+ ions 302 , and hydrolysis of MOF-1201(Ca) produces Ca 2+ ions, lactate, and acetate 303 .

Nevertheless, MOF's stability in water does not guarantee its stability in physiological conditions. Several anions present in the blood, such as phosphate, sulfate, and bicarbonate [START_REF] Conley | Encyclopedia Britannica[END_REF] , can act as Lewis bases or nucleophiles and owe some affinity to replacing the constituting organic ligands of MOF 189 . As a result, free organic ligands are released, and metal-containing species are formed. The nature of the metal-containing degradation products highly depends on replacing anions. The extent of degradation is directly related to the relative strength of metal-ligand bonding versus metal-replacing anion bonding.

While nanoMOF degradation can ease carrier elimination from the body, investigating these degradation products is crucial to verify MOF's toxicity. Several authors argue that a fast MOF degradation can lead to a spike in the local concentration of metals or organic linkers, which might cause toxicity 186 . A study found that ZIF-8(Zn) exhibits cytotoxicity towards both primary and cancerous cells due to the leaching of Zn 2+ ions produced by MOF degradation in cell culture media 304 . For biodegradable nanoMOF-based DDSs, MOF degradation affects their drug release kinetics, which can both benefit and backfire on their drug delivery performance. A progressive MOF degradation leads to controlled drug release, whereas fast degradation causes burst drug release. Note that sometimes, rapid degradation can be useful. For example, when MOF is coated on titanium prosthetics, degradation of MOF-74(Mg) coating induces an acidic environment and kills bacteria, which helps to protect the prosthetic from bacteria and increase the age of the prosthetics 305 .

Numerous MOFs are degradable in phosphate-containing media due to the high affinity of phosphate towards high valence metal cations. Phosphate buffer saline (PBS) is a typical simulated physiological medium to cause degradation of ZIF-8(Zn) 306 , MIL-100(Fe) 307,308 , MIL-88(Fe) 309 , and UiO-66(Zr) 298,310 . The group of Falcaro extensively studied the drug delivery application of ZIF-8(Zn) and its degradation mechanism 306 . As expected, nano-sized ZIF-8 degrade faster than micro-sized ones since the NPs with the larger surface area are more reactive. They both lost their crystallinity and became amorphous after the reaction with phosphate. The NPs morphology also changes from a rhombic dodecahedron to a spherical-like shape. The metal-containing degradation products are zinc-phosphate NPs, evidenced by the EDS map showing the appearance of P and disappearance of N, and the IR spectrum showing the new vibration modes of PO4 3-in the degraded micro-ZIF-8(Zn) (Figure 1 -18). release of trimesate ligand increases with the PBS concentration. The amorphous shell consists of iron phosphate, as shown by EDS and IR spectroscopy, and probably some iron oxide, as suggested by Mossbauer spectroscopy 308 . The core-shell structure of the degraded MIL-100(Fe) was an interesting finding because it does not entirely fit the categories of the bulk and surface erosion mechanisms as demonstrated for polymer NPs. The global dimension of the microparticle did not change, a feature of bulk erosion. However, the amorphization did not happen homogenously but concentrated at the particle surface, a feature of surface erosion. The maintenance of crystallinity and shape of MIL-100(Fe) was also observed in another study. Phosphoric acid was used to selectively etch the MIL-100(Fe)'s large cages and create hierarchical porosity. A spongy structure was obtained (Figure 1-19E, F), and releases of trimesate and Fe 3+ ions were detected 311 . Degradation mechanisms of nanoparticle-based drug delivery systems: the case of polymer conjugate and metal-organic framework

Another factor that further complicates the nanoMOF degradation mechanism's investigation is the diversity of MOF composition. In consequence, individual studies of each MOF are needed, but they should use systematic reaction conditions to allow comparison of different MOFs. Horcajada and coworkers found the degradation times of some MOFs in simulated body fluid (SBF) follow the order: MIL-100(Cr) < MIL-101(Cr) < MIL-53(Cr) (3, 7, and 21 days, respectively) when soaking ibuprofen-loaded MOFs compacted disk in SBF with the same mL SBF per mg drug 177,212 . Between MIL-100 and MIL-101 with the same constituents but different structures, MIL-100 degrades faster because of larger pores and pore windows, leading to better pore accessibility for phosphate ions HPO4 2-in SBF. MIL-53's long chain structure and small pores make it less reactive than the other two made of isolated metal clusters 189 . Likewise, the phosphate/metal ratio is a crucial parameter. The higher the ratio, the quicker the degradation process should be. Thus, one should be cautious of this condition when comparing the degradation reactions of different MOF families.

The medium pH also plays a significant role in dictating the MOF degradation mechanism.

In an aqueous medium with no coordinating species other than hydroxide, degradation of a carboxylate MOF is possible at both acidic and basic pH. Under basic pH, the carboxylate ligand might be replaced by hydroxide ions. Under acidic pH, it might be protonated, providing that the pH is lower than the pKa of the organic ligand. At acidic pH, the ZIF family is particularly unstable 189 . For instance, ZIF-8(Zn) degrades in acetate buffer pH 5 after only a few minutes, compared to 7 days in phosphate buffer (PB) pH 7.4 200 . Protonation of imidazolate linker at pH 5 (pKa of imidazole = 14.05) reduces zinc-imidazole bond strength, thus inducing the ZIF-8(Zn) decomposition. Contrastingly, ZIF-8 was found stable in the basic medium of boiling sodium hydroxide for one day 199 . This stability was explained by its hydrophobic nature, preventing the hydroxide from approaching the Zn(II) centers.

Degradation at basic pH follows another mechanism involving a competition between organic linkers and hydroxide to coordinate with metal cations. For example, UiO-66(Zr) is more stable in acidic pH than in basic pH; complete release of its terephthalate linker was detected at a pH above 9. Another study determined the trimesate ligand release from nanoMIL-100(Fe) loaded with antibiotics, which was affected by both the medium and drug loading 314 . The gentamicin-loaded nanoMIL-100(Fe) degrades faster than the empty one in PBS and slightly slower under acidic conditions. In the presence of pancreatin, the gentamicin-loaded and the empty nanoMIL-100(Fe) have similar degradation rates. Furthermore, contradictory results were reported when adding protein to the medium. Generally, protein corona absorbed on nanoMOF surface is expected to protect it from degradation. Indeed, several authors found that nanoMIL-100(Fe)'s degradation by phosphate ions diminishes when adding bovine serum albumin to phosphate medium 297,307,314 . However, the presence of albumin in the medium elevates the degradation of ZIF-8(Zn) NPs 316 .

Degradation of drug-loaded MOF is extensively used to achieve controlled drug release from nanoMOF carriers. A typical example is the pH-responsive DDS based on selective degradability at the acidic pH of the ZIF family. A study reported minor DOX release at a pH higher than 6.5, opposite to 93% release in 7-9 days at pH 5-6 228 . The same trend was witnessed for ZIF-90 317 . Likewise, Christodoulou and coworkers observed the controlled release of prednisolone phosphate (PP) associated with progressive degradation of drugloaded nanoMIL-100(Fe) in PBS, contrasting to the burst release of prednisolone and prednisolone sulfate (PS) associated with the fast release of trimesate ligand (Figure 1-20A, B) 297 . The authors also observed that PS was released in a sulfate-containing medium while PP was not (Figure 1-20C, D) because sulfate could not displace the phosphate drug from MIL-100(Fe).

Drug loading can also influence the nanoMOF degradation mechanism. The presence of the drug helps the nanoMOF carrier to resist degradation at pH 7.4, as seen in the case of π-π stacked topotecan and DOX 300,301 (explained in part 1.2.3.6). However, in the same study of DOX, its drug release from nanoMIL-100(Fe) significantly increased upon incubation of loaded MOF in artificial lysosomal fluid at pH 4.5, probably because of more intense MOF degradation 301 . This behavior opposes the previous finding that this MOF is more stable in PBS at pH 5.3 297 . Nevertheless, the degradation of drug-loaded MOF can be faster than that of the empty one, as observed for gentamicin-loaded nanoMIL-100(Fe) 314 . Likewise, an in vivo study found that busulfan-loaded nanoMIL-100(Fe) is less stable and degrades faster than the empty ones, thus being eliminated faster 318 . When the drug is held within MOF by coordination bonds, the extent of drug release and MOF degradation is affected by the competitive substitution between MOF's constituting ligand and MOF's drug ligand. The present writer demonstrated in chapter 3 that adenosine triphosphate is more efficiently and firmly coordinated with nanoMIL-100(Al) than adenosine monophosphate. Therefore, the degradation mechanisms of MOF loaded with these two drugs are also different. To summarize, nanoMOF is a promising porous drug carrier that offers high payloads for hydrophilic and hydrophobic drugs and various methods to control drug delivery functions. MOF can also be biodegradable. Nevertheless, like other DDSs containing inorganic components, toxicity, biocompatibility, and body clearance remain the main concerns for using MOFs for patients. There is a lack of systematic in vitro and in vivo studies compared to an abundant number of articles on proof-of-concept applying MOF for different drug delivery applications. Other problems with clinical translation are smaller size with homogenous distribution, scale-up synthesis, and stability/storage of dispersible powder. In particular, MOF degradation processes should be studied case by case to evaluate if MOF degradation products cause toxicity or accumulate in the body without being removed. In the very first step, the nature of MOF degradation products must be deciphered. Therefore, one of the works to be done is to understand the molecular mechanism of nanoMOF degradation in biological media and the nature of their degradation products. More importantly, leading researchers have argued that each MOF has its own degradation products, and consequently, each MOF must be investigated separately 186 . Thus, suitable characterization techniques and methodologies are much needed to analyze the degradation reactions of different MOF families.

Characterization techniques for drug delivery systems

One cannot stress enough the essential role of material characterization for any application.

For drug delivery, routine characterization of a DDS is needed to validate the design and formulation as a fundamental step to ensure its proper functions in vivo 319 . It is said that a DDS has two identities 186 : synthetic and biological identity, i.e., the properties as-synthesized and the behaviors in the living body. This section focuses on characterization techniques for the synthetic identity, specifically, the physiochemical properties of a DDS.

Routine characterization techniques for nanocarriers

Basic physiochemical characterizations of a nano-sized DDS are size and size distribution, morphology, surface properties, composition, physical stability, chemical stability, batch-tobatch consistency, and starting material characterization 319 . It is impossible to cover the techniques for all types of NPs. The techniques below are the most general and widely used for NP-based DDSs.

The size and size distribution of a colloidal suspension are typically the first things to be measured. Dynamic light scattering (DLS), or photon correlation spectroscopy, relates the size of sub-micron particles with their Brownian motion as larger particles move more slowly 320 . The result of DLS provides the average hydrodynamic diameters and size polydispersity of NPs in the suspension phase. The hydrodynamic diameter is the diameter of a sphere with the same translational diffusion coefficient as the particle. A polydispersity index equal to or below 0.2 is acceptable to be called a monodispersed sample. However, DLS has its limit for samples that are polydisperse or contain aggregation/agglomerate 321,322 . Electron microscopy is a method of choice when one wants to study individual NPs.

Transmission electron spectroscopy (TEM) and scanning electron microscopy (SEM) are the most used techniques. They can inspect not only the size but also the morphology of individual particles. However, it might be time-consuming if one wants to get a statistical value. It should be noted that the size observed under TEM/SEM corresponds to solid particles without considering the solvation shell. DLS and TEM/SEM are often used to ensure global and local views of DDS's size distribution.

The surface properties of NPs can be characterized by measuring the zeta potential (ZP), an indication for but not equal to the surface charge 321 . ZP is the surface potential at the slipping plane, the diffusion region where ions are bound loosely to a colloidal particle 62 . Typically, ZP measurement is based on electrophoretic light scattering that can be integrated with the DLS instrument. The magnitude of ZP could be related to the colloidal stability of a suspension with the general guideline that a colloidal suspension with ZP in the range of 0 ± 10 mV might be highly unstable. However, one should be cautious when interpreting ZP as some sterically stable NP suspensions can have low ZP values 321 .

The composition analysis of a DDS involves the determination of all components' concentration in the formulation, including the carrier, drug, surface coating, and exicipients 319 . Typical techniques for confirming the presence and the connection of these components with the carrier are X-ray diffraction (XRD) for crystalline materials and spectroscopy techniques such as infrared (IR), Raman, and nuclear magnetic resonance (NMR) spectroscopy. XRD is based on the diffraction of X-ray when encountering a regular and repeating arrangement of atoms. IR spectroscopy is based on the characteristic vibration of functional groups, giving signature bands whose energy lies in the IR region. NMR spectroscopy is based on the resonance of each nuclear when activated by radio frequency, emitting characteristics energy of atoms with specific chemical bonds and environments. Quantitative analysis of drug loading and surface coating should cover both the bound and free components in the formulation 319 .

Measuring drug loading usually deals with a small amount of substance, thus requiring highly sensitive analytical instruments like high and ultra-high performance liquid chromatography (HPLC and UHPLC, respectively) and liquid chromatography/mass spectrometry (LC/MS). HPLC relies on the separation of the molecules of interest by a column filled with adsorbents, followed by the quantification by a UV-vis detector. The biggest strengths of HPLC and UHPLC are high accuracy, precision, and especially low limit of detection. However, these techniques require an elaborate procedure with many instrumental and methodology parameters. Also, fluorescent and UV-visible spectroscopies are useful for drugs that have these specific properties. The problematic step is to separate the free and the loaded drug properly.

The physical stability of DDS in water and biological media is conveniently assessed by following its size and size distribution with time to check possible aggregation. Likewise, chemical stability, including degradation, drug leakage, drug release, should be monitored over time using similar analytical techniques for drug loading measurement. Typically, drug-loaded or unloaded nanocarrier samples are incubated with a chosen medium at 37 o C under continuous mixing induced by rotator, magnetic stirring, etc. After a given time interval, the suspension is separated by centrifuging. Then, the solid precipitate and the liquid supernatant are analyzed separately to find the amount of drug release and ligand release or analyze the structure of the degraded nanocarrier. The drug integrity after being released is also important to check. Note that drug release kinetics should be represented as the time evolution of % of its initial payload. Stability under the storage conditions (temperature, pH, lyophilization/freeze-drying) should also be verified.

Batch-to-batch consistency is a key criterion to ensure reproducibility, which is important for clinical translation required by FDA. Any of the properties above can be used to demonstrate reproducibility.

Starting materials characterization, even when they are commercial products, is highly recommended, at least for the drug and the carrier's material 319 . It is even more critical when synthesizing one's own carrier material (polymer and MOF NPs in this thesis). The characterization techniques of these two materials will be described later in the next section.

When analyzing a DDS, it is advantageous to have techniques that can offer both qualitative and quantitative information at once, simplifying and easing the research and development work. NMR spectroscopy is one of a few techniques capable of doing such things. It is beneficial not only for material's routine quality control but also for in-depth studies about reaction mechanisms. Taking advantage of the expertise in the host laboratory, the present PhD thesis is aimed to explore NMR spectroscopy to characterize the drug delivery processes of nanoparticles, eventually expanding the toolbox to monitor the chemical evolution of DDSs. The following subsection will be dedicated to summarizing the techniques used for starting material characterization, then introducing the scientific background of NMR spectroscopy and its application for drug delivery. Finally, in situ NMR spectroscopy will be covered.

Starting material characterization techniques

Some essential characterization techniques for polymer and metal-organic framework materials are presented.

For polymer, IR and liquid-state NMR (lqNMR) spectroscopy are routinely used to verify the chemical structure of the as-synthesized polymer, especially for the polymer's terminal group. 1 H and 13 C NMR measurements provide high-resolution structure determination and quantify its composition. The presence of functional or terminal groups can also be checked by IR spectroscopy. Additionally, mass spectrometry (MS) measures the mass-to-charge ratio. Polymer chemists use it extensively in matrix-assisted laser desorption/ionization (MALDI) mode to detect macromolecules with big molecular weights and long chains. The electrospray ionization (ESI) MS technique is more apt to detect small molecules. Regarding the degradation studies, liquid NMR and IR spectroscopy can analyze the chemical nature and chemical structure evolution of a degraded drug carrier. For instance, several studies on the degradation of PLA utilized lqNMR to detect and quantify the lactic acid and lactide produced 323,324 . MS can help to distinguish the monomers or oligomers in trace amounts produced from degradation or drug release processes 325 .

For the routine characterization of MOF, the most used techniques include XRD, IR spectroscopy, and N2 gas adsorption. XRD measurement of single or microcrystal is the technique of choice to determine the crystal structure of any potentially new MOF. For a MOF with a known structure, powder XRD serves to verify the crystallinity and the phase purity of the synthesis. The MOF porosity is usually evaluated by measuring its N2 gas adsorption/desorption isotherm, from which its specific surface area can be calculated using the Brunauer-Emmett-Teller (BET) model. IR spectroscopy is used to check the presence of the organic ligands and if they are coordinated to or trapped in the framework. Electron microscopy, such as SEM and TEM, is used to observe the morphology and size of MOF. SEM can be coupled with energy dispersive X-ray spectroscopy (EDS) to analyze the elemental composition and distribution. Also, thermogravimetric analysis (TGA) can provide information about MOF composition and its thermal stability. Inductively coupled plasma atomic emission spectroscopy (ICP-AES), shortly ICP, is used to determine its metal content.

A final note will be given to the use of NMR spectroscopy for MOF materials. Liquid NMR spectroscopy is usually used to characterize the organic ligand when the ligands are modified prior to MOF synthesis. It can also be used to analyze the digested MOF solution to get information about the linker ratio, functionalized ligand, leftover substances, and residual solvent. Solid state NMR (ssNMR) spectroscopy is especially helpful in probing the local environment. Deviations from periodic structures in MOFs can happen after chemical modification or functionalization, often not detected by XRD and IR spectroscopy. In this case, ssNMR spectroscopy can probe local atoms, their neighboring environments, and chemical bonds while keeping the MOF intact, in contrast to liquid NMR, which requires solubilization or digestion of MOF samples.

Regarding the degradation study, XRD and BET measurement are the first tools to check the integrity of MOF, as loss of crystallinity and loss of porosity are reasonable indications for MOF structure collapse. Note that MOF partially degraded can still retain its crystallinity 188 . The evolution of chemical structure upon degradation can be evaluated by IR and NMR spectroscopy. Solid-state NMR spectroscopy can detect changes in the ligand's molecular state or the metal cations' coordination sphere.

In this thesis, to investigate the drug delivery mechanisms of nanoMOF, we also use several techniques that are not necessarily in the MOF characterization routine, namely XPS, XANES, and SAXS. X-ray absorption near-edge structure (XANES) spectroscopy, a category of X-ray adsorption spectroscopy, uses X-ray irradiation to excite core electrons from the bulk of the material. Edge energy is the energy required to excite those core electrons in orbitals, e.g., 1s, 2s, 2p, 3s, corresponding to K1, L1, L2, and M1 edge energy, respectively. The edge energy for each element is unique. The shape and feature of this edge and near-edge structure contain information about the oxidation state, local atomic environment, and coordination structure (Figure 1 -22). In general, XANES is usually employed to study the coordination structure of transition metals which are very popular in catalysis and electrochemistry. It has also been used in drug delivery for verifying the integrity of the drug loaded in the carriers 327 . , where 2θ is the scattering angle, and λ is the wavelength of the incident X-ray. Figure 1-23 gives a simplified scheme. SAXS can give information about the NPs' size, size distribution, shape, and surface roughness, as well as their inter-particle interaction. The wave vector can be understood as the inverse of the actual distance in the sample. The small 𝑞 measured in SAXS reflects the interaction at the nanoscale of the sample, such as the interaction among particles/grains or between particles and media. With large 𝑞 measured in wide-angle X-ray scattering (WAXS), interaction at the atomic scale of the sample, i.e., the crystal distance, was probed. WAXS is essentially XRD for soft matter. SAXS profile can be divided into several regimes (Figure 1 -24). Three regimes of interest are Guinier, Power law, and Porod regime. The Guinier regime is where the scattering intensity can be related to an object's size as follows: 𝐼(𝑞) ≈ 𝐼(0)𝑒 -𝑞 2 𝑅 𝑔 2 /3 , where Rg is the radius of gyration, and 𝐼(0) is the intensity at zero scattering angle (𝑞=0). The Power law regime is the middle zone of a SAXS profile that follows a power law dependence 𝐼(𝑞) ∝ 𝑞 -𝛼 , which gives information about the particle shape and geometry. The Porod regime is the near-end zone of high q in a SAXS profile, where the scattered intensity follows 𝐼(𝑞) ∝ 𝑞 -4 . This Porod regime gives information about the degree of surface roughness relative to the particle size. Several studies on polymer and MOF-based DDSs have used SAXS to follow MOF formation, drug encapsulation, and drug release kinetics from polymeric micelles 330,331 . WAXS was used to follow the diffraction peak of ZIF-8(Zn) NPs during their degradation 306 .

NMR spectroscopy

Nuclear magnetic resonance (NMR) spectroscopy is based on the unique magnetic resonance signature of the nuclei of a given isotope to give information about a substance's atomic structure. NMR is the energy emitted by an atom, under a static magnetic field, after being excited with radio frequency.

An atom can be perceived as a magnet because of its intrinsic nuclear spin. An NMR-active atom must have a non-zero spin. Under the Earth's magnetic field, the magnetic moments of atoms orient randomly. The magnetic fields of individual atoms cancel each other, and in total, no magnetic field is felt. However, in the presence of an external magnetic field, these atoms, these individual magnets, are aligned following the main axis of the external magnetic field. This phenomenon is schematically represented in Figure 1-25. If we consider these nuclear spins have their own energy levels, their initial states are degenerate. When aligned, these degenerate states split into lower and higher energy levels, representing states aligning with and against the external magnetic field. Now, if those atoms are given a radio frequency (RF) radiation matching the energy gap between these ground and excited states, the nuclear spins are excited and transit to the excited state, i.e., spin aligned against the external magnetic field. This spin flip is called the Zeeman transition (Figure 1 -25). After turning off this RF radiation, the excited spins transit back to their ground state and emit the resonance frequency called the Larmor frequency. The nature of the Larmor frequency is the precession frequency of the nuclei around the static magnetic field. The Lamor frequency is dependent on the external magnetic field B0 following an equation: 𝜈 0 = 𝛾𝐵 0 , where 𝛾 is the gyromagnetic ratio. The gyromagnetic ratio is unique to each nucleus with a unique nuclear number and mass; therefore, the first feature of NMR spectroscopy is element-and atomicsensitive. The gyromagnetic ratio indicates how sensitive a nuclear spin is to excitation by an external magnetic field. Nuclei with a high gyromagnetic ratio and sufficient natural abundance produce more sensitive NMR measurement, i.e., more intense signal for the same sample quantity and measurement time. Several popular NMR nuclei include 1 H, 13 C, 31 P, 19 F, 27 Al, 7 Li. However, the measured NMR frequency of an atomic nucleus in a molecule is not exactly equal to its Lamor frequency. Electrons orbiting around an atom also create a local magnetic field which slightly alters the magnetic field felt by the nucleus. In other words, the nucleus is shielded by electrons. How effective this shielding depends on this nucleus's neighboring atoms. A neighboring electronegative atom pulls the electron density, thus exposing the nucleus to a stronger magnetic field. This nucleus shielding allows the distinction of chemically inequivalent environments, for instance, 1 H in the methyl group -CH3 and the methylene group -CH2.

This shift in NMR frequency caused by nucleus shielding is used against a reference compound's resonance to define a characteristic value in the NMR spectral scale. To eliminate the dependence of resonant frequency on magnetic field (𝜈 0 = 𝛾𝐵 0 ), division by the Larmor frequency at the instrument's magnetic field is needed. Overall, the chemical shift is defined as

𝛿(𝑝𝑝𝑚) = 𝜈 𝑠𝑎𝑚𝑝𝑙𝑒 -𝜈 𝑟𝑒𝑓𝑒𝑟𝑒𝑛𝑐𝑒 𝜈 𝑠𝑝𝑒𝑐𝑡𝑟𝑜𝑚𝑒𝑡𝑒𝑟
. For a given nucleus, e.g., 1 H, 13 C, or 31 P, a reference is chosen for the position 0 ppm, where there is almost no shielding from the surrounding atom. Typical ranges of 1 H NMR chemical shifts are presented in Figure 1-26.

The main interactions that affect the NMR frequency are schematically represented in Figure 1-27. The Zeeman interaction on the order of a hundred MHz has the most pronounced effect on the NMR frequency, followed by the quadrupolar interactions (x10 6 Hz). The chemical shift and the dipolar coupling are on the order of kHz, while the scalar coupling is within a few tens of Hz 334,335 . All these interactions are anisotropic, meaning they are dependent on the sample orientation. They affect the shape and the width of an NMR peak, thus giving information about the sample's structure. The scalar coupling, or the J-coupling (Figure 1-27B), is a through-bond interaction between spins. It depends on the covalent bond connectivity, character (sing or double bond), distance (one-, two-or three-bond), and angle 337 . The effect on NMR spectra is peak slitting, whose distance is related to the J-coupling constant. Different patterns of multiplet can then arise. The dipolar coupling (Figure 1-27C) is through-space interaction between spins. It is independent of the electronic structure and depends on the internuclear distances 337 . The quadrupolar coupling (Figure 1-27D) only exists for nuclei with spins higher than ½ that have a non-spherical quadrupolar charge distribution. These asymmetrically distributed charges interact with the electric field gradient generated by surrounding electrons.

Figure 1-27. NMR interactions, including (A) chemical shielding σ affecting the effective magnetic field felt by the nucleus, (B) scalar coupling J depending on through-bond contact between two nuclear spins, (C) dipolar coupling D depending on through-space distance between two nuclear spins, and (C) quadrupolar interaction between their nuclear charge distribution and local electrical field

gradient 337,338 Degradation mechanisms of nanoparticle-based drug delivery systems: the case of polymer conjugate and metal-organic framework Furthermore, NMR spectroscopy has different styles of measurement to get information about interatomic interaction. In a simple NMR measurement, irradiation with single or multiple RF pulses is followed by the acquisition of the resulting signal that decays to equilibrium in microseconds or seconds 334 . The time evolution of this free induction decay (FID) signal is Fourier-transformed to the signal intensity versus frequency NMR spectrum (Figure 1-28), which is named a one-dimensional (1D) NMR spectrum. In a two-dimensional (2D) NMR measurement, another time scale is added by varying the evolution time between pulsed irradiation and data acquisition. The collected data set is subjected to separate Fourier transformation of evolution time and detection time domain, producing a 2D NMR spectrum of signal intensity versus two frequencies, F1 and F2, which is actually a threedimensional dataset considering the intensity data. The 2D NMR spectrum is usually presented by a contour plot, the horizontal cross-section of the 3D plot (Figure 1-29) 339 . Such correlation spectra between identical nuclei, e.g., 1 H-1 H, or between different nuclei, e.g., 13 C-1 H, 31 P-1 H-, etc., support information for determining molecular structure. Another equally significant feature of NMR spectroscopy is its quantitative nature. Given several appropriate conditions, the area of an NMR peak is proportional to the number of nuclei contributing to that signal 340 . Considering an NMR signal of intensity 𝐼 corresponding to the number 𝑁 of observed nuclei, then: 𝐼 = 𝛼𝑁, where 𝛼 is a constant that remain the same for all resonances in one NMR spectrum. Feasible applications are structure identification, quantification of a substance or a mixture of substances, and purity determination. Proton ( 1 H) NMR spectroscopy is the most used owing to its availability in organic compounds and high NMR sensitivity. Many other experimental parameters contribute to the accuracy and precision of quantitative NMR measurement 341 . Relaxation delay should be at least five times the relaxation time T1 to ensure the magnetization reaches equilibrium so that one records the NMR signal at its highest possible intensity. Acquisition time should be long enough to record the complete FID signal. The number of scans should be compromised between a sufficient signal-to-noise ratio and manageable manipulation time. In the case of liquid samples, a solvent suppression pulse sequence can be used to maximize the NMR intensity of the sample. Quantitative NMR measurement can be carried out in relative or in absolute manners 341 . Relative quantification is the easiest method, where the molar ratio between two compounds A and B can be calculated by their relative NMR integration using the equation:

𝐼 𝐴 𝐼 𝐵 = 𝑛 𝐴 𝑛 𝐵 × 𝑁 𝐴 𝑁 𝐵
, where n is the molar and N is the number of nuclei. Usually, a reference compound with a known concentration is added to the sample, and the sample composition can be determined. This internal reference can be mixed homogenously with the sample or placed in a coaxial insert inside the liquid NMR tube. The latter avoids the interference of reference compound with NMR signals of analyte and potential peak overlapping 342 . In an absolute quantification method, a calibration should be done beforehand to establish the linear relationship between an analyte's NMR peak integration and its molar. The same set of experimental parameters should be used for calibration and sample measurements to ensure the calibration's validity. When it is impossible to use a reference compound, the ERETIC method (Electronic reference to access in vivo concentration) was introduced, using an electronic reference signal to create an artificial reference NMR peak 343 .

To sum up, NMR spectroscopy is a useful tool to decipher molecular structure because it is 1) atomic-selective, 2) sensitive to the surrounding environment of atoms and bonds, 3) non-destructive, 4) quantitative, and 5) flexible to measure samples at different states like solid, liquid, and gases. A wide range of materials, from synthetic inorganic and organic materials 344 to macromolecules like proteins 345,346 , can be analyzed by NMR spectroscopy. It is worth mentioning that the non-invasive and non-destructive nature enables NMR spectroscopy to be used for studies of biological samples or clinical studies of drugs inside those samples 347 . However, the main limitation is low sensitivity 334 . A few tens to hundreds of milligrams of material are usually required, while only a few milligrams are needed for other techniques like IR spectroscopy. Also, studying compounds containing paramagnetic nuclei can be highly challenging. In everyday use, NMR is the principle of magnetic resonance imaging (MRI), a staple of medical diagnostic technique 348 .

A distinctive feature of NMR spectroscopy is that the technique varies for sample phases and intermolecular interaction. In the following sections, we will cover liquid-state, solidstate, and high-resolution magic-angle spinning NMR spectroscopy, as well as NMR relaxation and relaxometry.

Liquid-state NMR spectroscopy

The most used NMR spectroscopy technique is liquid(-state) NMR spectroscopy (lqNMR), sometimes called solution NMR spectroscopy. It is dedicated to liquid-phase samples such as solution, pure liquid, and liquid mixture. In a solution, all solutes exhibit the Brownian motions, tumble and rotate quickly; eventually, they distribute homogenously in the solvent, and there is no preferred orientation 349 . Therefore, the anisotropic NMR interactions are canceled, and only the isotropic values of NMR interaction are observed 337 , resulting in narrow NMR peaks (width a few Hz) in a lqNMR spectrum. The spectral resolution, the ratio between the NMR spectrum's total width and the NMR peak's linewidth in the liquid phase, is much higher than in the solid phase (the reason will be discussed later). Thus, lqNMR used to be called high-resolution NMR spectroscopy. The sample should be homogenous, and any precipitation should be filtered. For 1 H NMR, solvents with no proton are preferred to avoid signal saturation and maximize the sample signal. Deuterated solvents like D2O, CDCl3, d6-DMSO, etc., are common. A liquid sample of 500-600 μL is typically placed inside a cylindrical glass tube of 5 mm outer diameter, called an NMR tube 350 . LqNMR is the staple tool of organic chemists to determine the molecular structure of as-synthesized compounds.

Although there is no general framework on how lqNMR can be used in drug delivery applications, one may find it helpful in characterizing a pharmaceutical formulation and a DDS. The group of Alonso has used proton NMR spectroscopy to characterize the composition and structure organization of lipid NPs. NP suspensions of tripalmitin, lecithin, and PEG-stearate were prepared (all tripalmitin, lecithin, and stearic acid are lipids) using deuterated water 351 . Quantitative analysis by 1 H NMR peak integration of NPs suspension, 1

H NMR peak integration of the starting ingredients' physical mixtures, and 2D HMQC peak volumes showed that tripalmitin has the highest weight proportion of the NPs, while lecithin and PEG-stearate together account for only 20%. Then, 1 H NMR spectra were recorded at different temperatures (Figure 1 -30). When the temperature was raised to 55 o C, the signal intensity at 1.2 ppm increased, assigned to the melting of stearate linked to PEG. At 66 o C, the melting of tripalmitin and lecithin resulted in their signal appearing in the NMR spectrum. As the solid species cannot be detected by liquid-state NMR measurement, it was concluded that the lipid NPs composed of a tripalmitin solid core surrounded by PEG coating, which was consistent with the high wt% of tripalmitin. This structure also agreed with the difference between the NMR spectra of NPs suspended in deuterated water (D2O) and isolated NPs dissolved in deuterated chloroform (CDCl3). The first showed only one broad and small peak at 1.2 ppm, while the latter contained signals of most of the functional groups in the aliphatic region below 2.8 ppm. Only the PEG coating, still in contact with liquid water and mobile to some extent, was visible in the lqNMR of NP suspension. In a subsequent study of these lipid NPs, the same authors combined previously used 1 H NMR measurements with diffusion coefficient and relaxation measurement to prove that a liquid lipid could be incorporated inside the solid tripalmitin core and exist as a mobile liquid phase there 352 . Finally, the release process of encapsulated phosphate porogen in PLGA microspheres incubated in fetal bovine serum for 22 days was studied by 31 P NMR 355 . The obtained series of NMR spectra displayed an upfield increase of chemical shift and a significant decrease of linewidth, which indicates the conversion of diphosphate to monophosphate as PLGA degraded to glycolic and lactic acid and the internal environment pH dropped.

Monitoring self-diffusion coefficients of drugs have found many applications for characterizations of DDS, e.g., identification of species in the formulation, probing the drugcarrier interaction, monitoring drug release process, and even studying the in vivo fate of drug carrier 356 . This has been done by diffusion-ordered NMR spectroscopy (DOSY), which yields a two-dimensional spectrum with NMR chemical shift in the F2 dimension and the diffusion coefficients in the F1 dimension 357 . DOSY is a useful technique for distinguishing components in a complex mixture and by their diffusion coefficients. Liu et al. used DOSY to determine the structure type of microemulsion 353 . In microemulsions made of less than 20 wt% aqueous phase, the diffusion coefficient of water is 10-time lower than that of isopropyl myristate (oil phase), suggesting an O/W structure. Another study found a relation between the diffusion coefficients of the drug encapsulated in microemulsion and its transdermal permeation rates across human skin 358 . The author suggested using this technique to optimize the formulation composition without having to conduct in vitro study in advance.

Additionally, quantitative NMR might be useful for studying the kinetics of drug release, degradation, or any reaction happening during the drug delivery processes. A possible benefit of using quantitative NMR rather than the intricate HPLC technique is that it is easier to set up quantitative NMR protocol to distinguish mixtures of released drug and carrier degradation products. Liu and coworkers used quantitative 1 H NMR to follow the degradation process of polymersome based on PLA-PMPC, copolymer between poly(lactic acid) and poly(2-methacryloyloxyethyl phosphorylcholine) 359 . Figure 1-32 displays the proportion of remaining lactate units over time. A pH-dependent degradation of PLA was likely to cause the polymersome-to-micelle transition, which could be used for pH-trigger drug release. Another study examined the role of the PLA's stereo configuration on the degradation process of PLA nanofiber in air, water, and physiological serum 323 . The time evolution of PLA MW, an indication of degradation, was determined by 1 H NMR integration of the tertiary proton group in PLA. In another application, liquid NMR spectroscopy was used to determine the content of the surface functionalization groups after hydrolysis of silicon and silica NPs, taking advantage of distinctive 1 H NMR peaks of the NPs and the coated molecules 360,361 . The authors optimized the digestion protocol and the NMR acquisition parameters. The quantitative method was applied to study the effect of PEG input on the PEGylation of silicon NPs 360 . Although lqNMR is easy to use and produces high-resolution routine characterization results, not all substances we use are liquids. The pharmaceutical industry also produces a lot of solid dosage forms, which are easier to be stored than liquid and suspension dosage forms. That is where solid(-state) NMR spectroscopy (ssNMR) comes in. As the name says, it is dedicated to solid samples of various forms, be it powder, crystal, amorphous, soft matter, or any solid piece of material 334 .

The most distinguished difference between ssNMR and lqNMR is that Brownian motion does not exist in the solid state. All the anisotropic NMR interactions not detectable by lqNMR are now visible in ssNMR spectroscopy. Figure 1-33 illustrates the influence of anisotropy on ssNMR spectra. Each nuclear spin with a slightly different orientation has a discrete NMR frequency. As there are many orientations, there are a huge number of resulting NMR peaks. Therefore, a broad featureless signal is eventually observed in solidstate NMR spectra. The isotropic averaging by Brownian motion in liquid can be mimicked by mechanical spinning around an axis at high speed; thus, all nuclei spin are rotated at the same frequency respective to their neighbors 362,363 . The spinning frequency should be of kHz order for sufficient averaging. In addition, the anisotropic part of chemical shielding, dipolar interaction, and first-order quadrupolar interaction is proportional to the term (3cos 2 θ-1), which is equal to zero at θ = cos -1 (√1/3) ≈ 54.74 o , the so-called magic angle 363 . The concept that combines these two effects is named magic-angle spinning (MAS) 364 . By spinning the sample around an axis tilted by 54.7 o from the static magnetic field axis (Figure 1-34A) under a jet of compressed air, a narrow liquidlike NMR spectrum can be obtained for solid samples (Figure 1-34B). The specimen holder is a cylindrical zirconia rotor with an outer diameter of typically 4 mm (Figure 1-34C). Faster MAS rates lead to better peak resolution but require a smaller rotor. Small rotors can be problematic because of challenging rotor fabrication and limited sample quantities, hence low NMR sensitivity. Currently, the MAS rate ranges from 5-100 kHz for cylindrical rotors of diameters from 7 to 0.7 mm (Figure 1-34C), which can correspondingly contain between 500 and 1 mg of sample 334 .

However, the anisotropic interactions are both a curse and a blessing. The line broadening caused by these interactions significantly reduces the spectral resolution but can provide crucial information about the 3D structure and local geometry in solid materials in their native forms, such as interatomic space distance and molecular motion 334 . Several NMR experiments are designed to reintroduce some of these desired interactions without affecting the spectral resolution.

There are other methods to increase the quality and expand the capabilities of ssNMR. Cross-polarization coupled with magic-angle spinning (CPMAS) is one of the most common experiments in ssNMR because it enhances the signal sensitivity for nuclei with low natural abundance and low gyromagnetic ratios such as 13 C, 15 N, and 31 P by transferring magnetization from naturally abundant spins which are most often 1 H 365,334 .

For quadrupolar nuclei, only the first-order component of quadrupolar interaction is removed by MAS. Significant broadening is still problematic for spectral resolution, especially when the remaining second-order quadrupolar interaction is intense and the chemical shifts of several resonances are close. Multiple-quantum magic-angle spinning (MQMAS) is the main experiment to resolve this quadrupolar broadening by correlating individual line components along the isotropic dimension in a 2D spectrum 334,366 . SsNMR spectroscopy has various applications for pharmaceuticals and drug delivery. It has been used to distinguish solid polymorphs of APIs, characterize the APIs within a pharmaceutical formulation and identify their purity 368 . Polymorph identification is particularly important because those different states can give rise to different properties and biological activity of the drug. Also, NMR crystallography, a combination of ssNMR and simulation to determine the structure of amorphous drugs, has gained attraction recently, pioneered by the team of Emsley 369 . For the DDS, ssNMR has been used to investigate the drug-carrier interaction, the state, and the location of the loaded drug.

Atsutoshi Ito et al. investigated the recrystallization during storage of troglitazone loaded in polyvinylpyrrolidonesolid dispersions prepared by solvent and co-grinding methods 370 . The two formulations, despite having similar XRD patterns, exhibit different rates of troglitazone recrystallization during the next seven days. However, their initial 13 C CPMAS spectra showed a slight difference in peak intensity and line width of troglitazone, which was associated with different drug conformation within the solid dispersion. This example demonstrates the ability of ssNMR spectroscopy to investigate local short-range structural properties compared to long-range sensitive XRD. Other authors encapsulated phosphatecontaining porogen in PLGA microspheres to study their hydration mechanism 355 . Decoupled 31 P MAS NMR spectra displayed a shift of the 31 P resonances of the dibasic and monobasic sodium phosphate porogen after the encapsulation, thus evidencing their successful encapsulation.

Mesoporous silica and MOF-based DDSs benefit the most from ssNMR techniques; there are more studies about them in the liteturature [371][372][373] . One of the first ssNMR reports on mesoporous silica DDSs investigated the molecular state of ibuprofen encapsulated in two MCM-41 materials of 35 and 116 Å pores (Ibu35 and Ibu116, respectively) 374 . The 1 H MAS NMR of free crystalline ibuprofen showed a broad featureless line, whereas that of ibuprofen loaded in mesoporous silica has sharp lines, indicating that ibuprofen is mobile in the MCM-41 matrix (Figure 1-35A). Correlation MAS NMR measurements of 13 C-1 H and 29 Si-1 H showed weak interaction between the drug and the silica matrix, partly contributing to the high drug release rate. This high mobility of drug molecules in mesoporous silica was also observed in the system of phenylphosphonic acid and SBA-15 375 . An opposite scenario occurred to the aminopropyl-SBA. The drug was rigid and interacted strongly with the carrier through the aminopropyl group 375 , proved by similar broad 1 H MAS NMR spectra of free and loaded phenylphosphonic acid (Figure 1-35B). In the first report of nanoMOF drug carrier, results from ssNMR supported the encapsulation of ibuprofen (IBU) inside the pore of MIL-100(Cr) and MIL-101(Cr) 177 . 13 C MAS NMR showed a significant shift of the α-CH3 peak of the ibuprofen's carboxylic acid after drug loading and the absence of the carboxylic peak due to ibuprofen's interaction with

Chapter 1 Introduction -Literature review paramagnetic chromium cations. In the subsequent studies that attempted to detect the interaction of nanoMIL-100(Fe) with its surface coating, the paramagnetism of MIL-100(Fe) was problematic. Therefore, a strategy was proposed to use a diamagnetic structure analog, MIL-100(Al) 264 . The mechanism behind the successful loading of phosphate drug in MIL-100(Al) and the selective location of loaded drug adenosine triphosphate (ATP) and cyclodextrin-phosphate (CD-P) surface coating was demonstrated by MQMAS and 31 P- 27 Al correlation NMR (Figure 1 -36). The MQMAS series of the pristine, ATP-loaded, CDP-coated, and loaded-coated nanoMIL-100(Al) reveal distinctive signals of Al species in the core interacting with the ATP and those at the surface interacting with CD-P. This selective interaction was further justified by 31 P- 27 Al MAS NMR, which showed effective interaction within 5Å between MIL-100(Al) and ATP/CD-P.

Likewise, ssNMR helped to understand the interaction between indomethacin and the MIL-101 framework. The 13 CP MAS spectra of the loaded MOF confirmed the presence of indomethacin and the solvent tetrahydrofuran, which was in contrast with the loaded mesoporous silica SBA-15. SBA-15 has a long cylindrical channel with a 9 nm diameter, while MIL-101(Al)-NH2 has spherical cavities with small windows of 1.2 nm diameter; as a result, the solvent can migrate from SBA-15 more easily. More importantly, the 1 H-1 H and 13 C-1 H MAS NMR spectra show very weak interaction between indomethacin and the framework but intense interaction with tetrahydrofuran 376 . High-resolution magic-angle spinning (HR-MAS) NMR spectroscopy has been developed to measure heterogeneous samples containing solid and liquid phases, such as suspension, gel, tissues, cells, etc. Those samples cannot be measured by lqNMR because it will miss the solid part and produce very broad peaks for the liquid part. The source of line broadening in the NMR spectra of these samples comes from the residual dipolar coupling and the magnetic susceptibility gradient in heterogeneous samples 377 . The dipolar coupling is explained in the previous section. The magnetic susceptibility factor describes the sample's permeability to the magnetic field. The higher the magnetic susceptibility, the higher the magnetic field created in the sample induced by the external magnetic field 377 . Samples placed in different homogenous liquid media have different discrete chemical shifts. In the case of a heterogeneous medium, different interfaces between the sample and the medium cause varied magnetic susceptibility. Also, a gradient of distance-dependent dipolar coupling between the solvent's and sample's nuclear spins results in a gradient of chemical shift. These two effects combine into broad signals in the resulting NMR spectra 377 . MAS can also be used to reduce this line broadening effect, as the magnitude of the solvent-induced magnetic field also contains the term (3cos 2 θ-1). In some cases, a spinning rate of only a few kHz is sufficient to average this magnetic susceptibility 378 . In a simplified way, HR-MAS is the marriage of lqNMR and ssNMR, using the MAS concept to obtain high-resolution lqNMRlike spectra. HR-MAS NMR found many applications for biological samples like blood or tissues 379 or foodstuffs in the food industry 380 . In material science, HR-MAS techniques have appeared in studies of polymer gel, polymer membrane, surface-modified NPs, etc. 378 In many cases, the samples for HR-MAS are liquid-like or fluid. Those samples are packed in tightly sealed inserts which are then placed in the NMR rotor to prevent sample dehydration or liquid leakage into the spectrometer (Figure 1 -37). Regarding applications in drug delivery, HR-MAS NMR has been used to characterize active compounds in pharmaceutical formulations and study the drug-DDS interaction. For instance, lidocaine and prilocaine, two local anesthetics in commercial products, a waterbased gel (Kamistad Gel) and an O/W emulsion-based cream (Emlacream) could be detected by 1 H HR-MAS NMR. Furthermore, the concentration of lidocaine in the gel could be determined from the calibration of different gel samples 381 . In another study, 2D 13 C-1 H measurement in a HR-MAS NMR probe confirmed the presence of galactose and lactose on the surface of mesoporous silica NPs. It only took 2 h compared to at least overnight for an equivalent ssNMR experiment 382 .

A study of polymeric hydrogel 383,384 built a model for drug transport accounting for both drug-drug and drug-polymer interactions. The diffusion of the drug in hydrogel showed a different behavior from that of the free drug in D2O. At higher drug concentration, the diffusion coefficient of the drug in D2O solution decrease as a result of increasing viscosity. Meanwhile, the diffusion coefficients of drugs loaded in hydrogel rose, explained by the saturation of adsorption sites for drugs in the hydrogel. The interaction between hydrogel and drugs helped to prevent the drug-drug interaction that leads to drug aggregation.

A study combined the use of ssNMR, lqNMR, and HRMAS NMR to work out the structural organization of nanocrystal suspension of Carbamazepine (CBZ) and CBZ-saccharin (SAC) with hydroxypropyl methylcellulose (HPMC) as the solvent and sodium dodecyl sulfate (SDS) as the stabilizing agent 385 . Solid-state 13 C NMR of the suspension confirmed the presence of the CBZ dihydrate and SAC and revealed the presence of solid-phase HPMC, suggesting the solidification of HPMC during the suspension preparation process. The solidified HPMC phase on the API nanocrystal surface helps to stabilize their suspension. In addition, 1 H lqNMR spectra showed the presence of APIs and SDS stabilizer only in the supernatant centrifuged from the CBC and CBZ-SAC nanocrystal suspension. The absence of APIs and stabilizer's signal in the lqNMR spectra of the original suspension was due to the fast exchange between the bound and free states in nanocrystals. Finally, 1 H HR-MAS NMR was used to detect and quantify the semi-solid phase in these suspensions.

NMR relaxation and relaxometry

NMR relaxation describes the response of nuclear spins toward equilibrium when they are subjected to an external magnetic field and excited RF waves. Like the chemical shift, relaxation time is also an intrinsic property of a nuclear spin that depends on its surrounding environment. There are two main types of relaxation time in NMR spectroscopy: longitudinal relaxation time T1 and transverse relaxation time T2. The corresponding relaxation rate is the inverse of the relaxation time, R1 and R2, equal to T1 -1 and T2 -1 , respectively. The T1 relaxation is the process when the net magnetization grows along the z-axis from zero to its possible maximum when a magnetic field is turned on, i.e., from when nuclear spins are aligned randomly to when they are all aligned following the magnetic field 386,387 (Figure 1 -38). How fast this process is can be described by the expression: 𝑀 = 𝑀 0 (1 -𝑒 -𝑡/𝑇1 ), where T1 is the longitudinal relaxation time. The T2 relaxation process describes the relaxation of the transversal component of magnetization oriented in the xy plane perpendicular to the applied magnetic field: 𝑀 𝑥𝑦 = 𝑀 0 𝑒 -𝑡/𝑇2 , where T2 is called the transverse relaxation time 388 . When a 90 o -RF pulse is applied, the net magnetization M is rotated into the xy plane, but not all nuclear spin components are necessarily aligned toward the M direction (Figure 1 -39). After the RF pulse is turned off, those individual nuclear spins dephase randomly, and Mxy decreases. It reaches 37% (= 1/𝑒) of its initial values after a time of T2. The fluctuation of Mxy relates to the line width of the NMR signal. When T2 is long, Mxy decays slowly, the line width is narrow, and vice versa. T1 and T2 relaxation times depend largely on the size and the motion of the nucleus. Solid molecules that move slowly transfer their energy slowly, too; thus, it takes more time for the net magnetization to grow, hence long T1. Meanwhile, T2 becomes shorter, causing large linewidth in ssNMR spectra. For instance, in the previously mentioned study of ibuprofen in mesoporous silica 374 , the highly mobile ibuprofen had a T2 8-time longer than the free crystalline ibuprofen and a short T1. An opposite scenario was observed for daunorubicin (Dau) loaded in poly(butylcyanoacrylate) (PBCN) NPs. Its T1 relaxation time in the free drug solution or the physical mixture with PBCN was shorter than in the Dau-loaded-PBCN NPs.

The limited mobility of Dau could be due to its hydrogen bonding with the polymer chains 389 . In addition, T1 could give information about the location of a drug in microemulsion. For example, the T1 of lidocaine in a microemulsion composed of water, isostearyl isostearate (oil phase), and surfactants are close to that in the neat isostearylic isostearate and are shorted than in the neat water, suggesting that lidocaine was primarily located in the oil phase 390 . This observation agrees with the higher solubility of the drug in isostearyl isostearate than in water. It is worth mentioning that the T2 relaxation could be useful for NMR imaging. This technique was used to monitor quantitatively the drug release from swelling polymer tablets 391,392 .

NMR relaxation time varies with different magnetic field strengths. Relaxometry refers to the study of variation of relaxation time as a function of magnetic field strength. However, a standard NMR spectrometer only has a single magnetic field strength. A relaxometer can measure the changes in relaxation time T1 as a function of the magnetic field with only one instrument. Fast-field cycling (FFC) NMR relaxometry probes the evolution of longitudinal relaxation rate T1 -1 versus magnetic field from a few kHz to around 100 MHz 393 . The obtained dependence is called a nuclear magnetic resonance dispersion (NMRD) profile. As the T1 is related to the molecular motions, i.e., rotation and translation, changes of T1 at a low magnetic field correspond to the process that lasts between nano to milliseconds, the range of chemical exchange process 394 .

A study combined relaxation measurements at a high field (100 MHz) and FFC relaxometry to study the relationship between the structure of a hydrogel made of xanthane, a natural hydrophilic polymer, and the pH-dependent drug release 395 . The hypothesis is that changing the xanthane fraction in the hydrogel could affect the polymer chain dynamics, thus influencing the hydrogel's swelling and drug release rate.

In situ NMR spectroscopy

In situ means the same location, and in situ in chemistry typically means in the reaction mixture. In situ characterization techniques in chemistry are highly appreciated for providing information about reaction mechanisms 396,397 . They can trace the process dynamics and the evolution of reactants and intermediates without disturbing the system's status, which, in most cases, cannot be done in ex situ studies. In the context of drug delivery, in situ techniques can be used to study the mechanisms of drug release and carrier degradation, as well as the interaction of carriers with substances in biological media. It is especially beneficial for microparticle-and nanoparticle-based DDSs whose ex situ studies usually need a separation step by centrifugation or filtration. In addition, the labor and time spent on kinetics studies are reduced by avoiding sampling after specific time intervals.

In the literature, many in situ studies employ fluorescence-related techniques to monitor NPs in real time. A study in 2011 followed the degradation of a PEG-dextran hydrogel, showing a good correlation between in vitro and in vivo measurements 398 . PEG was conjugated with a fluorescent dye, and its stability over time was checked. The release of degradation products was tracked through the fluorescent intensity (Figure 1 -40). Another study used the phenomenon of FRET (Förster/Fluorescence resonance energy transfer) to examine in situ the physiochemical properties and the drug release ability of PEG-PCL (polyethylene glycol-polycaprolactone) NPs 399 . Before NP formulation, BODIPY and Nile red were attached to the end group of PEG-PCL through a click chemistry reaction. The release of rhodamine was quantified in situ by the calibration of the FRET acceptor/donor ratio versus the quantity of encapsulated rhodamine. The limitation of this approach is that one needs to incorporate a fluorescent tag in the sample and ensure they stay attached over time. Based on these examples above using Raman and IR spectroscopy, it is reasonable to expect that NMR spectroscopy is applicable for characterizing in situ DDSs, as it can provide both qualitative and quantitative information while being a non-destructive technique. For NPs, a prominent class of DDs, NMR spectroscopy might be able to distinguish their bulk and surface structure 403 . LqNMR spectroscopy can measure diluted NPs dispersion, and suspensions with high viscosity or high concentration are suitable for ssNMR and HR-MAS NMR 403 . However, it should be noted that in most cases, solid-state species cannot be detected in a lqNMR spectrometer. Liquid species can be recorded in a ssNMR instrument but with compromised spectra quality compared to a dedicated lqNMR spectrometer 404 .

Different applications of in situ NMR spectroscopy in drug delivery have been reported. For example, Westesen and colleagues used 1 H lqNMR to characterize the state of several poorly water-soluble drugs incorporated in lipid NPs 405 . The loaded drug changed from being mobile (narrow line width) in lipid emulsion to being confined (broad line width and signal intensity decrease) in lipid suspension. It was concluded that lipid suspension is more suitable for controlled drug release, as highly mobile drugs might experience burst release.

A similar principle was used to assign the 13 C CPMAS NMR spectra of poly-nbutylcryanoacrylate nanocapsule suspension in water 406,407 . The CP measurement recorded both narrow and broad signals overlapping with each other. Comparison with the reference spectra of the solid polymer strengthened the assignment of broad signals to the polymer at the solid phase. The narrow signals were assigned to the surfactant dissolved in the aqueous phase and the triglyceride core of the nanocapsules. Also, the transformation of linewidth could be used to indicate the degradation mechanism of particle suspension 406 . If the solid particles decompose into small fragments, the mobility of these polymer fragments increases, accordingly decreasing the linewidth of their 13 C CPMAS signals. If the degraded solid particles are solubilized in the liquid phase, well-resolved signals can be observed in direct excitation mode.

Furthermore, different NMR methods have been developed to monitor several dynamic processes in situ. The Kenneth's group is one of the few active players in this research area. Firstly, they succeeded in using solid-state 13 C CPMAS NMR spectroscopy to monitor the crystallization from solution 408 . For the proof of concept, glycine containing two polymorphs was used. The NMR technique in this paper only detected the evolution of species at the solid phase. Later, a "Classic NMR" experiment was published, which allows the detection of both solid and liquid phase 409 . Using a home-made NMR pulse program that alternatively operates two different pulse sequences for solid-and liquid-NMR, the corresponding spectra are alternatively recorded, giving information about the time evolution of solid and liquid species together (Figure 1 -42). These in situ NMR methods have been applied for studying crystallization processes of several polymorphs 404 . In 2021, the same group published their in situ NMR investigation on the formation mechanism of a nickel MOF 410 , using a ssNMR spectrometer in an attempt to monitor species evolution at both solid and liquid phases. 1 H MAS NMR spectra provided quantitative information about reaction kinetics in the liquid phase. Nonetheless, the authors did not manage to obtain information about the solid phase. Degradation and drug release processes of the polymer conjugate NPs were studied qualitatively and quantitatively using mass spectrometry and liquid NMR spectroscopy to characterize the molecular structure of released species. We discovered a complex NP hydrolysis process forming diverse oligomers and monomers of lactic acid (LAC) and the drug ITA. The slow degradation process led to a low release of free drugs, although raising the pH from 5.3 to 7.4 induced a slight increase in the amounts of released products. TEM images showed that bulk erosion is likely to play the main role in the degradation of PLA-ITA NPs. The overall results and methodology can be of interest for designing polymer-drug conjugate NPs and understanding their degradation and drug release.

Introduction

Tuberculosis (TB) is the 13 th leading cause of death worldwide and the second deadliest infectious disease, only after COVID-19 (above HIV/AIDS) 1 . According to the World Health Organization's global report on TB in 2021 2 , nearly 1.5 million people die from TB yearly. It is caused by a bacterium named Mycobacterium tuberculosis (Mtb), infecting about onequarter of the world's population, of whom 5-15% will fall ill 1 . Mtb typically affects the lungs (pulmonary TB) and spreads through air droplets when a sick person coughs or sneezes. Sometimes, Mtb can remain dormant in the body for an extended period and reactivate even years after infection 3 . The current treatment method for this dangerous disease includes a standard 6-month use of four first-line antibiotics: isoniazid, rifampicin, ethambutol, and pyrazinamide 2,[4][5][6][7] . This drug regimen can successfully treat approximately 85% of TB-infected patients 8 , but it is complicated, time-consuming, and associated with numerous side effects, frequently leading to poor patient compliance or mismanagement from health workers. Consequently, drug-resistant Mtb bacteria strains have emerged 9,10 .

There is a clear need for more efficient TB treatments to overcome drug resistance and shorten the drug course. In addition to the development of novel antibiotics 11 , increasing focus has been put on using nanocarriers to deliver novel and existing anti-TB molecules. In comparison with the administration of free drugs, nanoparticulate drug delivery systems (DDSs) can offer drug co-encapsulation, protection from biological barriers, and prolonged drug release, thus reducing dose frequency. Also, nanoparticles can be designed to target tissues and cells infected by Mtb, such as alveolar macrophages, thereby simultaneously improving the bioavailability of the active substances and decreasing their adverse effect, overall increasing the therapeutic efficiency.

In this context, the original active ingredient in this study is itaconic acid (ITA), a human metabolite playing a role in the immune response 12 and having potential antimicrobial and anti-inflammatory activity 13 . It has never been used before for TB treatment, but Michelucci et al. showed that ITA might inhibit the growth of Mtb 14 . However, ITA as a free drug cannot penetrate inside the infected cells. That is why it is necessary to design a suitable carrier. A possibility to achieve targeted intracellular delivery of ITA is taking advantage of the more acidic environment inside the TB-infected cells or granuloma (small area of inflammation 15 ) than inside normal cells 16,17 . A suitable DDS that releases ITA upon a pH gradient becomes necessary in this case.

Among the types of pH-responsive drug nanocarriers, polymer nanoparticles (NPs) are the most popular because of their abundant methods to control the incorporation and release of both hydrophilic and hydrophobic drugs, as well as long-known biocompatibility and biodegradability 18,19 (reviewed in chapter 1.2.2). Two primary strategies to make pHresponsive polymer NPs are charge shifting and acid-sensitive linkage 20,21 . In the first strategy, polymers containing ionizable groups undergo swelling or dissociation upon pH variation. The other strategy designs polymers composed of pH-sensitive functional groups in their backbone or pH-sensitive covalent linkage with drugs. The latter is called polymerdrug conjugate.

Nevertheless, incorporating ITA into polymers is not straightforward due to its high solubility in water. Also, to control the drug release from polymer-drug conjugate, one should understand the degradation mechanism of this specific system.

In this chapter, we demonstrate a new approach to build a bioconjugate for itaconic acid based on polymers of lactic acid and investigate the degradation and drug release mechanism of the resulting poly(lactic acid)-itaconate (PLA-ITA) NPs.

Poly(lactic acid), or polylactide (PLA), whose monomer is lactic acid, a human metabolite, has long been used for various implants, micro-and nano-carriers because of its trusted safety for human use 22,23 . The physical properties and the degradability of PLA could be adjusted by varying the proportion of d-and l-lactic acid stereoisomers 24 .

The drug ITA will be conjugated to PLA to avoid the issues associated with the physical encapsulation of hydrophilic molecules in hydrophobic polymer NPs 25 . For synthesizing a polymer-drug conjugate, essentially a polymer prodrug, the "drug-initiated" method is a trendy approach in which a drug is used to initiate the polymerization of monomers to form short polymer chains. Compared to the conjugation method with functionalized polymer, the "drug-initiated" method is simpler, consists of fewer manipulation steps, and is more likely to produce homogenous polymer chains with the same molar mass and terminated with a drug 25 . Pioneering studies in the "drug-initiated" synthesis of polymers were carried out by Tong and Cheng 26 . They employed native functional groups of the drug for ringopening polymerization or pre-functionalized drugs for reversible deactivation radical polymerization techniques [26][27][28][29] . In a recent study in our group, pipemidic acid, an antibacterial drug, was successfully conjugated to poly(ε-caprolactone), which was then formulated into NPs under a nanoemulsion procedure 30 .

For studying the degradation process of the PLA-ITA polymer-drug conjugate, liquid NMR spectroscopy is a promising tool that provides structural characterization at the molecular level (chapter 1.3.3.1). It has been used to investigate the core-corona structure of lipid and polymeric nanoparticles [31][32][33] . For instance, a 1 H solution NMR study showed that NP structure assembled from amphiphilic poly(lactic acid)-poly(ethylene glycol) PEG-PLA depends on the solvent hydrophilicity and the copolymer's molecular weight 33 . Further applications of liquid NMR spectroscopy to track the location in the carrier, the release kinetics, and the diffusion coefficients of the active substances have been reported and reviewed [34][35][36] . Herein, liquid 1 H and 13 C NMR spectroscopy were mainly used to determine the nature and the amount of the released products.

The present study consists of three parts:

1) Elaboration of the PLA-ITA polymer-drug conjugate NPs, including a polymer synthesis followed by a formulation optimization step

2) Identification and quantification of possible degradation products of PLA-ITA NPs, by using mainly liquid NMR spectroscopy with the support of mass spectrometry

3) Monitoring the stability of original and degraded PLA-ITA NPs

Synthesis of PLA-ITA polymer-drug conjugate

In the first step, the synthesis of PLA-ITA polymer-drug conjugate was inspired by the report of itaconate-terminated poly(l-lactic acid) oligomers of Okuda and coworkers in 2012 37 . They showed that it was possible to ring open itaconic anhydride by the OH-terminal group of a PLA chain using tin octanoate, paving the way to a one-pot process using the standard lactide polymerization catalyst. Reasonable chemoselectivity (87:13) for the more reactive C-5 carbonyl was observed providing β-mono ester of itaconic acid. To maximize the amount of itaconic acid while maintaining the self-assembly capacity of the oligomer, we planned to implement a two-directional growth of the polymeric chain, thus allowing installing two terminal hydroxyl groups for reaction with itaconic anhydride (Figure 2-1). Thus, conventional tin-catalyzed PLA synthesis using d,l-lactide ring-opening polymerization was undertaken starting from 1,3-propanediol as initiator. After completion, the mixture was cooled, and an excess of itaconic anhydride was added. After 3 additional hours of heating at 100°C, the crude material was washed by acid and purified by precipitation, affording the desired material as a white solid. The synthesized material was characterized by 1 H and 13 C NMR, IR spectroscopy, and mass spectrometry to determine its structure. The 1 H NMR spectrum of the obtained product (Figure 2-2) shows the characteristic peaks of the repeating unit of lactic acid at 5.17 and 1.57 ppm and the signals of the propane diol at 4.20 and 1.99 ppm. The chemical shift of lactate in PLA-ITA is similar to that in PLA alone and agrees with the literature 38 . The signature peaks of the two terminal itaconate moieties at 5.88 and 6.45 ppm correspond to the protons of the methylidene groups =CH2. The smaller singlet at 5.83 can be tentatively assigned to the minor α-mono ester of itaconic acid as observed previously 37 . Peak integration indicated a number-average molecular weight Mn of 3850-4050 g.mol -1 , which roughly corresponds to 25 lactic units for both hydroxyl groups of the propanediol. Table 2-

Optimization of polymer NP formulation

Once the polymer PLA-ITA was produced, we optimized the formulation of PLA-ITA NPs, based on the NP size and polydispersity. The data of mean hydrodynamic diameter and the polydispersity index (PdI) were given in appendix A.2. Among the tested formulations, the single emulsion formulation of 15 mg.mL -1 NPs, using dichloromethane (DCM) as organic solvent and polyvinyl alcohol (PVA) as emulsifier gave the most concentrated suspension with the lowest particle size polydispersity. The obtained PLA-ITA NPs had an average diameter of 331 ± 12 nm and a PdI of 0.09 ± 0.05.

Nevertheless, the single emulsion formulation was not suited for the degradation study. In a first attempt, PLA-ITA NPs prepared by single emulsion were incubated at pH 7.4 and 5.3 for over one month. The obtained supernatant was measured by liquid NMR spectroscopy. The 1 H NMR spectra (Figure 2-3) show difficulties in accurate analysis. Firstly, the peaks of interest of ITA in the region 5.5-6.5 ppm have very low intensity, hindering both qualitative and quantitative characterization. This low intensity is due to low ITA loading in 7.6 wt% of PLA-ITA, which corresponds to maximum ITA release below 3 mmol.L -1 . However, it is not easy to increase the wt% of ITA because it will increase the solubility of the bioconjugates in aqueous media, risking its ability to form NPs in water. The second issue seen from To avoid these issues, PLA-ITA NPs were prepared by nanoprecipitation without using PVA emulsifier. In the absence of emulsifier, formulation with small volume and low NP concentration was used to prevent potential aggregation. The attained mean hydrodynamic diameter was 343 ± 36 nm, and the PdI was 0.49 ± 0.07. In addition, to improve the NMR sensitivity, the concentration of the analyzed sample was increased to the maximum. Specifically, samples of degradation products were freeze-dried, and the obtained precipitates were redissolved again in a minimal amount of D2O. The 1 H NMR spectra of freeze-dried degradation supernatant of PLA-ITA nanoprecipitation are displayed in Figure 234. They present several peaks assigned to methylidene protons CH2= of ITA in the region 5.7-6.6 ppm (Figure 2-4A), which contrasts with two singlets in the same region of free itaconic acid (see Figure 2-2 in section 2.2). There are also more signals assigned to methylene -CH2-protons in the region of 3.25-3.70 ppm (Figure 2-4B) compared to a single peak in the spectrum of itaconic acid solution. Such an increase in NMR signature was not observed when incubating ITA at similar degradation conditions of PLA-ITA NPs (Figure 2-5A), indicating that this phenomenon is due to the degradation reaction of PLA-ITA and not to the interaction between itaconic acid and phosphate salts.

Likewise, there are several quadruplets and doublets corresponding to the methanetriyl -CH protons and methyl -CH3 protons of LAC in the region 4.0-5.5 and 1.2-1.8 ppm, respectively (Figure 2-4C, D). Such multiplication of LAC's NMR signatures was detected in lactic acid aqueous solution and when incubating lactic acid at similar conditions of PLA-ITA degradation (Figure 2-5B). Espartero et al. evidenced that lactic acid in solution and lowmolecular-weight PLA can form different oligomers from monomer up to octamers, leading to multiple characteristic peaks for free lactic acid, hydroxyl and carboxylic end groups, and main chain units in 1 H NMR 40 . Moreover, the NMR spectrum becomes more complicated for a mixture of d,l-isomers. Likewise, the existence of oligomers made of itaconate and lactate was also demonstrated by mass spectrometry (MS). The mass spectra of degradation products in PBS at pH 7.4 for two months (see appendix A.3.) showed peaks of negatively charged (z = -1) fragments of lactic acid (m/z 89.3) and itaconic acid (m/z 129.1), as well as peaks of mixed esters between ITA and LAC monomers or propane-1,3-diol (Table 23).

ITA-propane-2- diol C8H12O5 [M-H] - 187.1 LAC-propane-2- diol C6H12O4 [M-H] - 147.1
The remaining parts of the PLA-ITA NPs were analyzed by the MALDI MS. Some of the main peaks from the positive mode MALDI spectra of original and degraded PLA-ITA were listed in Table 2-4. In the spectrum of degraded PLA-ITA (Figure 23456789) recorded without added cations, the peaks for [M+Na] + and [M+K] + were present because the instrument's glass chamber is made of Na and K. When Na + was added, the peaks for [M+K] + decreased their intensity, thus supporting the peak assignment. Moreover, the MALDI mass spectrum of residual PLA-ITA NPs after degradation exhibits signals of PLA-ITA with two-time shorter chain lengths than the original one (Table 234), indicating i) the PLA-ITA NPs were degraded into smaller fragments, consistent with the NMR and ESI-MS data, but ii) the degradation is still incomplete even after 1-2 months. From the experimental evidence so far, it appears that degradation of PLA-ITA NPs not only produced itaconic acid and lactic acid but also created oligomers of ITA-LAC and ITApropane-1,3-diol. It was reported by Kulkarni and co-workers that PLA incubated in a basic medium undergoes random scission of polymer backbone 41 . The current study shows consistent results. Random hydrolysis of ester bonds between ITA and the PLA chain and among lactate units leads to the formation of a mixture of monomers and oligomers.

Quantification of drug release and polymer degradation

Furthermore, quantitative 1 H NMR allowed us to determine the % release of ITA drug and LAC monomer, and the effect of pH on drug release and NPs degradation. Table 2-5 shows that ITA-containing and LAC-containing species formed faster at neutral pH 7.4 than at acidic pH 5.3. It took 2 months for the degradation at pH 5.3 to produce a similar level of % ITA release to that of degradation at pH 7.4 after 1 month.

Ester hydrolysis can be catalyzed by both acid and base. Here, knowing that the starting pH of original PLA-ITA NPs is slightly acidic (pH = 5-6), incubation of PLA-ITA at pH 7.4 causes a bigger pH increase than at pH 5.3. Therefore, hydrolysis of ester bonds between LAC and ITA in PLA-ITA conjugate NPs possibly follows a BAC2 type mechanism catalyzed by OH - anion 42 . A similar random ester bond cleavage was previously reported for the decomposition of low-molecular-weight PLA in basic media 43 , in which the rate of lactic acid formation differs from the rate of ester bond loss. In addition, the slow degradation of PLA-ITA NPs, hence the slow release of ITA, could be related to the lack of a hydroxyl terminal group. de Jong et al. demonstrated that the alcohol terminus of PLA participated in its hydrolysis process both in acidic and basic environment 44 .

Nevertheless, the slow release of ITA conjugated with PLA NPs driven by pH change is more advantageous than the burst release of unconjugated drug 45 . To get a global view of how the degradation process happens, the size, morphology, and stability of original and degraded PLA-ITA NPs were monitored for 2 months. On the one hand, the size distribution and zeta-potential of the original NPs stored at 4 o C for 2 months had less than a 10% deviation (data not shown). On the other hand, the degraded PLA-ITA NPs suspension aggregated, whose size became too big to be measured by DLS. Thus, we used TEM to observe the evolution of NP size and morphology. Figure 2-10 shows that as time progressed, while the original NPs kept their morphology, the PLA-ITA NPs degraded at pH 7.4 became transparent. This evolution could be a sign of a bulk erosion mechanism, in which water diffusion into the NPs is faster than the water hydrolysis reaction 46 . Water molecules penetrate, and the chain scission is initiated throughout the bulk of NPs without NP disintegration. Before a certain degree of hydrolysis, the NP size is roughly unchanged 46 .

Noteworthily, degradation at pH 5.3 after 1 month seemed to produce swollen NPs (Figure 2-10C). This swelling phenomenon was not observed in degraded pure PLA NPs without ITA at the two studied pH (Figure 2 -11), suggesting that the swelling was related to the presence of hydrophilic ITA within the conjugate NPs. Plausibly, at this pH 5.3, PLA-ITA NPs which is more hydrophilic than pure PLA NPs has a higher tendency to swell.

Conclusions and perspectives

We presented a method to synthesize a polymer-drug conjugate based on lactic acid and itaconic acid and formulate it as NPs. Also, we set up a method to study drug release and degradation qualitatively and quantitatively using liquid NMR spectroscopy, which enables the analysis of the nature and the amount of released species during NPs degradation and drug release (recap scheme below).

Using this method with the support of mass spectrometry, we found that the degradation of PLA-ITA polymeric NPs in different pH environments did not proceed by simple hydrolysis of the itaconic terminated PLA chain but rather by a complex route. A variety of oligomers and monomers made of LAC and ITA formed. Ester hydrolysis appeared to be slightly favored at neutral pH than at acidic pH. The degradation process, hence drug release, remained very slow. Combining information from electronic microscopy, we propose that bulk erosion was likely to happen. Altogether, this work highlights the importance of investigating the degradation and drug release products early in nanoparticle formulation development.

Looking forward, the knowledge obtained in this mechanistic study may bring some suggestion to the design of polymer-drug conjugate NPs. The synthesis and characterization methodology presented here could be adapted for formulating other polymers and drugs and studying the degradation of NPs made of polymer-drug conjugates or drug-loaded polymer NPs.

Experimental methods

Chemicals

All reagents and solvents from commercial suppliers were used without further purification.

For polymer synthesis: Itaconic anhydride, tin(II) 2-ethylhexanoate and d,l-lactide were purchased from Sigma-Aldrich Chemical Co. Diethyl ether was distilled from sodium/benzophenone ketyl. Toluene and dichloromethane were distilled from calcium hydride under a nitrogen atmosphere. All reactions involving air-or water-sensitive compounds were routinely conducted in glassware flame-dried under a positive pressure of nitrogen or argon.

For polymer nanoparticle preparation and degradation: Salts, including KH2PO4, Na2HPO4, KCl, and NaCl, were purchased from SigmaAldrich. Organic solvents, including dichloromethane, acetone, and ethyl acetate, were commercial analytical-grade products. Deionized (DI) water was taken from the PURELAB Ultra ELGA purification system. Poly(lactic acid) used to prepare PLA NPs was purchased from Sigma Aldrich Resomer® R 202 S, Poly(d,l-lactide), MW = 10000-18000 g.mol -1 .

For the NMR experiments, the deuterated solvents were deuterium oxide (D2O) and deuterated trichloromethane (CDCl3) from Innova-Chem, +99.8% atom D.

Synthesis of polymer-drug conjugate PLA-ITA

A solution of anhydrous 1,3-propanediol (freshly distilled over CaH2, 285 mg, 3.75 mmol) and d,l-lactide (10.0 g, 69.4 mmol) in anhydrous toluene (5 mL) was degassed using three freeze-thaw cycles. Tin(II) 2-ethylhexanoate (130 mg, 0.3 mmol) in toluene (0.2 mL) was added through a syringe, and the resulting mixture was placed in an oil bath at 140°C. After 8 h, the obtained gel was cooled to room temperature, and itaconic anhydride (2.0 g, 17.8 mmol) was added. The reaction mixture was heated at 100°C for 3 h. After cooling, the mixture was concentrated under reduced pressure. The residue was taken up into CH2Cl2 (150 mL), and washed with 0.1 N HCl (10 mL) and brine (10 mL). The organic phase was dried over MgSO4 and concentrated under reduced pressure. The residue was taken into CH2Cl2 (5 mL) and added dropwise into Et2O (100 mL). The solid was filtered and dried under vacuum. The process was repeated to precipitate an acetone solution of the polymer into water. After drying under vacuum over P2O5, the title compound was obtained as a white solid (6.2 g).

Preparation of PLA-ITA nanoparticles

Single emulsion

The organic phase was prepared by dissolving 60 mg of PLA-ITA conjugate polymer in 1.5 mL dichloromethane (DCM). The aqueous phase was polyvinyl alcohol (PVA) solution of 0.5 wt%. 4 mL of the aqueous phase was added to the organic phase, and immediately afterward, the resulting mixture was vortexed for 20 seconds. Subsequently, it was sonicated at 20% power for 1.5 minutes and then at 10% power for 30 additional seconds (Bandelin tip sonicator), using a sonication tip dip inside the sample vial. The sample vial was placed in an ice bath to avoid heating the suspension due to sonication. The sample was stirred by magnetic stirring overnight at room temperature to evaporate the residual organic solvent. A sheet of laboratory paper was used to cover the vial from dust.

Nanoprecipitation in small volume

The organic phase was prepared by dissolving 10 mg PLA-ITA polymer in 0.75 mL acetone.

The aqueous phase was deionized (DI) water. The organic phase, i.e., the polymer solution, was added dropwise (1 drop of roughly 10 L per second) into 2 mL water under continuous gentle magnetic stirring. The sample was stirred overnight at room temperature to evaporate the residual organic solvent. A sheet of laboratory paper was used to cover the vial from dust. After solvent evaporation, each batch was sonicated for about 10-15 s, before all batches were combined into a single sample in a Falcon tube.

The NP suspension was stored in the fridge at 4 o C.

The concentration of the formulated PLA-ITA NPs was estimated from the weight of the polymer divided by the volume of DI water remaining after evaporation of organic solvent.

To determine the volume of solvent inside the sample vial, the empty vial containing the magnetic stirrer was weighed beforehand, and the sample vial was weighed after overnight evaporation.

Degradation of PLA-ITA nanoparticles

Briefly, the degradation of PLA-ITA NPs was tested in phosphate buffer saline (PBS) at pH 7.4 and 5.3 in DI water. After a given period, the suspension in PBS was separated into supernatant and pellet, named degraded supernatant and degraded pellet. The obtained degraded supernatant was freeze-dried. Afterward, the freeze-dried precipitates were dissolved in deuterated water D2O to be measured by liquid NMR spectroscopy.

Incubation of 4.5 mg.mL -1 PLA-ITA NPs in PBS 10 mmol.L -1 was carried out as follows. From the 5 mg.mL -1 stock NPs aqueous suspension, 12.6 mL was pipetted into a 15 mL Falcon tube. Then, 1.4 mL of PBS 100 mmol.L -1 at pH 7.4 or 5.3 was pipetted into this tube. The sample tube was closed, mixed, and placed in a rotator inside an oven set at 37 o C. After 1 and 2 months, roughly half of the incubated sample tube was transferred to another 15 mL Falcon tube to avoid centrifuging a full tube. These two tubes were centrifuged at 10000 RCF (relative centrifugal force) for 15 minutes. The supernatant from two tubes was put in another tube. The pellet was dried at room temperature. The degraded supernatant was freeze-dried overnight using a Christ Alpha 2-4 LSCbasic lyophilizer. Maximum amount of the freeze-dried precipitate was dissolved in 600 µL D2O; the solution was then transferred to an NMR tube.

PBS 100 mmol.L -1 in H2O was prepared as follows. Stock A was NaH2PO4 solution of 0.1 mol.L -1 , and stock B was KH2PO4 0.1 mol.L -1 . Both these salts were quickly dissolved in DI water. To 40 mL of stock A and stock B, 3.2 g of NaCl (1.37 mol.L -1 ) and 0.08 g of KCl (27 mmol.L -1 ) were added and dissolved. Different proportions of stock A and stock B were mixed to obtain different pH. 20 mL PBS 100 mM with pH 7.4 and 5.3 was comprised of 4 mL stock A and 16 mL stock B, and 19.5 mL stock A and 0.5 mL stock B, respectively. The pH of PBS 100 and 10 mmol.L -1 was checked by a digital pH meter.

2.6.5 Characterization of PLA-ITA polymer and PLA-ITA nanoparticles

For characterization of the synthesized PLA-ITA, IR spectra were obtained as a solid or neat liquid on a Fourier Transform Shimadzu IRAffinity-1 spectrometer. Only significant absorptions are listed. The 1 H and 13 C NMR spectra were acquired on Bruker Avance 300 (300 MHz and 75 MHz, for 1 H and 13 C, respectively) spectrometers. Recognition of methyl, methylene, methine, and quaternary carbon nuclei in 13 C NMR spectra was based on the Jmodulated spin-echo sequence.

The size and colloidal stability of NP suspension were evaluated by dynamic light scattering (DLS) and zeta-potential (ZP) measurements using a Zetasizer Nano ZS90 apparatus from Malvern Panalytical. Samples of 0.15 mg.mL -1 of NPs were prepared in DI water for DLS or 1 mmol.L -1 KCl for ZP measurements. pH measurements were conducted on a Hanna Instruments edge® pH digital meter.

Morphology and size of NPs were observed by TEM under a 120 kV and 80 kV JEOL JEM-1400 microscope. NP suspensions were deposited on copper grids and observed with staining. For the original polymer NPs, negative staining was done with phosphotungstic acid. For the polymer NPs degraded at pH 7.4, following the negative staining with phosphotungstic acid, the TEM grid was washed twice with water to remove the residual salts from the degradation medium. For the polymer NPs degraded at pH 5.3, the TEM grid was washed twice with water prior to negative staining with uranyl acetate to prevent precipitating uranyl acetate in phosphate. Note that negative staining of these aciddegraded samples with phosphotungstic acid (pH 7.4) did not work, probably due to incompatible pH. Then the TEM grid was washed twice with water to remove the residual salts.

Solution NMR spectra for the degradation study were recorded at room temperature. ESI (electrospray ionization) mass spectra were recorded on an LTQ-Velos Pro Thermofisher Scientific spectrometer. MALDI-TOF (Matrix-assisted laser desorption-ionization time of flight) mass spectra were recorded in an UltrafleXtreme apparatus (Bruker Daltonics) using the facility from Small Molecule Mass Spectrometry platform of ICSN (Centre de Recherche de Gif -http://icsn.cnrs.fr). The ESI MS and MALDI MS measurements were performed in negative and positive ion modes, respectively.

For ESI MS measurement of the water-soluble degradation products, the residual phosphate salts in the pellet had to be removed. Specifically, after freeze-drying the degraded supernatant, the acquired powder was sonicated in 2 mL methanol for 20 minutes. The sample is left still for 1 hour before centrifuging 10000 RCF for 5 minutes and removing the supernatant. This liquid was evaporated under a fume hood for 2 nights, leaving the solid, which was then suitable for ESI MS measurement.

Studying degradation mechanism of metal-organic framework nanocarriers

Abstract

To make a drug work better, the active substance can be incorporated into a vehicle for optimal protection and control of drug delivery time and space. For making the drug carrier, the porous metal-organic framework (MOF) is one of the latest candidate materials as it can offer high drug loading capacity, various possibilities to design an effective drug delivery system, biocompatibility, and biodegradability. Nevertheless, the MOF degradation process is complex and not easily predictable, and the toxicity concern related to the degradation products has troubled the clinical translation of MOF. In the current chapter, we present an in-depth degradation mechanism of the aluminum-and iron-based nanoMIL-100 materials exposed to phosphate buffer saline (PBS). Using a combination of analytical tools, we propose a stepwise reaction mechanism between nanoMIL-100 and phosphate, which starts with the substitution of labile water ligand and continues with the substitution of trimesate ligand. As a result, aluminum phosphate/iron phosphate and free trimesate were formed.

Coordination between the metal(III) and phosphate, trimesate release, and the correlation between these two processes were proved. Furthermore, the extent of material erosion, i.e., bulk or surface erosion, was assessed by studying the transformation of nanoparticles' surface, morphology, and interaction with water. The impact of drug loading and surface coating on nanoMIL-100 degradation and the affecting factors were also determined. This study provides solid evidence about the possible degradation products of nanoMIL-100 in a simulated physiological condition. The findings of this study could be the groundwork for investigating the degradation of nanoMIL-100 carriers in a more complex scenario.

Introduction

The ongoing quest to search for the right drug delivery system (DDS) to treat leading death causes and the blooming of nanotechnology in the 21 st century has resulted in an everexpanding library of drug carriers. One of the most recent materials for DDS is metalorganic framework (MOF) (reviewed in chapter 1.2.3), a coordination network between metal and large-size organic ligands in a fashion that creates large voids in MOF structure. This exceptional porosity gives MOF the ability to absorb a large quantity of drugs in a small amount of carrier. Therefore, the amount of foreign materials that need introducing to the patient body decreases, potentially reducing the treatment's unwanted effects. Additionally, MOF's diverse chemistry arising from its organic and inorganic origins brings various methods to twist and tune its functions for drug delivery. As a result of these two mentioned advantages, MOF reported in the literature has a high payload for a wide variety of APIs 1,2 and has been used for the treatment of various diseases 3 . Multiple drug loading and surface modification strategies have been developed (see section 1.2.3). Nevertheless, due to the coordinative nature of its chemical bonds, the clinical applications of MOF have faced the challenge of unpredictable and non-generalizable degradation behaviors and associated toxicity. Therefore, each MOF needs individual analysis of the degradation mechanism and nature of degradation products.

In chapter 3, we want to study the degradation mechanisms of drug carriers made of MOF nanoparticles (nanoMOFs), currently being invested for biomedical applications slightly more than microscale MOFs (microMOFs).

The MOFs of our interest are nanoMIL-100(Fe) and nanoMIL-100(Al). The MIL-100 (MIL = Matériaux de l'Institut Lavoisier de Versailles) is the first reported example of nanoMOF used for drug delivery 4 . The structure of MIL-100 is displayed in Figure 3-1. MIL-100 consists of trivalent metal cations such as Al(III), Fe(III), and Cr(III), and tridentate ligands benzene-1,3,5tricarboxylate, also called trimesate or BTC [5][6][7] . Its secondary building unit (SBU), the M(III) trimer (M = metal), consists of three metal cations (III) coordinated with one central bridging oxygen and six carboxylate groups from several trimesate ligands. The 6-fold coordination sphere is completed with hydroxyl and water ligands. Two metal cations coordinate with one water molecule per cation, and one cation coordinates with one hydroxyl group. Every four SBUs are interconnected through trimesate linkers, creating tetrahedra. These tetrahedra are the nodes to build the MTN type MIL-100 framework, which possesses two kinds of pores: i) large mesoporous cages of 29 Å diameter, accessible through hexagonal (8.6 Å) and pentagonal windows (4.8-5.5 Å), and ii) small mesoporous cages of 25 Å diameter, accessible exclusively through pentagonal opening windows. The generally accepted formula is M3O(H2O)2OH(C6H3(CO2)3)2 (for M = Fe, Al) 6,8 . Several advantages of MIL-100 as a drug vehicle are the following. First, the big cages and large pore windows of MIL-100 are beneficial for capturing active molecules, leading to satisfactory drug payload and encapsulation efficiency. Second, it can create several types of bonding and interaction with drug molecules, resulting in controlled drug release. In the MIL-100 family, nanoMIL-100(Fe) is one of the most studied MOF-based DDSs. Its high drug loading capacity and controlled drug release have been demonstrated for numerous examples of active compounds, including anticancer [10][11][12] , antiretroviral 13 , antibiotic 14,15 , and anti-inflammatory 16 . For example, the anticancer drug busulfan was loaded in nanoMIL-100(Fe) with a payload of 25 wt%, five and sixty times higher than its payload in polymer nanoparticles and liposomes, respectively 10,17,18 . In addition to the general structural advantages of MIL-100, MIL-100(Fe), in particular, is made of biocompatible iron. A pioneering in vivo study found no immune or inflammatory reactions over three months after injecting high doses (110-220 mg.kg -1 ) of nanoMIL-100(Fe) in Wistar rats 19 . It also detected trimesate and iron eliminated from rats.

Likewise, nanoMIL-100(Al), the aluminum structural analog of MIL-100(Fe), has shown some potential for drug delivery applications. Feng et al. described a MIL-100(Al) gel with a remarkable payload of 62 wt% for the anticancer drug doxorubicin (DOX) 20 . In another publication, DOX was efficiently encapsulated (74% encapsulation efficiency) in a core-shell nanoMIL-100(Al) coated with γ-cyclodextrin-citrate and was released progressively depending on the initial drug loading 21 . NanoMIL-100(Al) was also biocompatible in vitro and did not induce cell toxicity for several lung and liver cell lines 22 .

Notably, both nanoMIL-100(Al) and nanoMIL-100(Fe) have shown specific interaction with phosphate-containing molecules, predictable from the thermodynamically favored complexation between Al(III)/Fe(III) and phosphate ions 23,24 . This affinity towards phosphate has been employed to achieve high drug payloads and stable coating for nanoMIL-100. For example, gemcitabine monophosphate was loaded up to 30 wt% with more than 98% efficiency, in contrast with less than 1 wt% of non-modified gemcitabine 11 . Agostini and coworkers pointed out a similar tendency for azidothymidine (AZT). Both monophosphate and triphosphate derivatives of AZT (AZT-MP and AZT-TP, respectively) were better incorporated in nanoMIL-100(Fe) than the pristine AZT 25 . More significantly, the authors also illustrate stronger binding with Fe(III) of AZT-TP than AZT-MP. In a study of nanoMIL-100(Al) and adenosine triphosphate (ATP), the authors evidenced the existence of the coordination Al-O-P bond between ATP and MIL-100(Al) upon ATP loading 26 , which explains the successful incorporation of phosphate-functionalized drugs in nanoMIL-100 carrier. Regarding surface modification, Agostoni et al. proposed a one-step coating process with phosphate-functionalized cyclodextrin (CD-P) 27 . The presence of strong complexing phosphate groups resulted in a stable coating obtained within minutes in water. It was supposed that bulky CD-P could not pass through the pore opening of nanoMIL-100(Fe).

The resulting coating neither alters MOF surface area (necessary properties for high drug loading) nor induces MOF degradation. In the subsequent study with nanoMIL-100(Al) 5 , solid-state NMR spectroscopy evidenced the Al-O-P coordination bond between the surface Al(III) and CD-P.

However, this strong interaction with phosphate also causes degradation of nanoMIL-100 at high phosphate concentration respecting to nanoMOF concentration. As phosphate is present in human blood in non-negligible amounts, it is significant to study the reaction of nanoMIL-100 analogs in phosphate, which is the target of our study.

Despite the abundant literature on drug loading and drug release, only a few studies have addressed this critical topic of the nanoMIL-100 degradation mechanism. Several studies found that cell medium containing phosphate salts causes MIL-100(Fe) degradation, commonly believed to produce free iron(III) ions and trimesic acid 13,14,28 . The release of trimesic acid was not observed in other simulated physiological media that do not have phosphate salt ingredients. Notably, Li et al. discovered that despite witnessing more than 30% trimesate ligand loss, the MIL-100(Fe) nano-and micro-particles maintained their sizes and shapes upon exposure to phosphate buffer 29 . Based on the data of Raman, Mossbauer, and IR spectroscopy, the authors suspected the formation of an amorphous shell of iron phosphate and/or iron oxide surrounding an intact crystalline core. Christodoulou and coworkers also found that the nanoMIL-100(Fe) and microMIL-100(Fe) degraded without collapsing or changing their dimensions using in situ atomic force microscopy and highresolution transmission electron microscopy 9 . Moreover, MIL-100(Fe) degradation started at the particle surface, where the crystal planes at the surface disappeared within 2 hours of incubation in phosphate buffer saline (PBS) at pH 7.4. Nevertheless, to our knowledge, none of these mentioned studies provide detailed evidence about the atomic-scale structure of degraded MOF.

More significantly, chemical modification of nanoMOF can influence its degradation process. Simon-Yarza et al. reported that the nanoMIL-100(Fe) loaded with the anticancer busulfan was degraded and eliminated more slowly than the non-loaded ones 18 . Another example showed that the degradation of MIL-100(Fe) loaded with prednisolone phosphate led to a more progressive formation of trimesate ligand than degradation of the empty one 30 .

In this context, this project attempts to decipher the degradation mechanism of two analogs, nanoMIL-100(Al) and nanoMIL-100(Fe), both of whom are potential candidates for DDSs. Notably, the iron-trimesate MOF shows better loading capacity towards DOX than the aluminum-trimesate 21,31 . Using a drug loading formulation of 1 mg.mL -1 nanoMOF mixing with DOX solution, at the same input DOX loading, the payload, hence the loading efficiency of nanoMIL-100(Fe) is always slightly higher than that of nanoMIL-100(Al) (1.2-1.5 times).

Other authors found that nanoMIL-100(Al) degraded in modified Eagle medium (MEM), a cell culture medium, while nanoMIL-100(Fe) did not 22 . From the existing literature, the iron and aluminum-based MIL-100 have some differences in their chemistry, but it is not certain yet how that difference will play a role in their degradation and drug release mechanisms.

To learn the effect of phosphate groups on drug delivery processes like drug-carrier interaction, carrier degradation, and drug release, we chose the active compounds and the surface coating that bear phosphate moieties. Adenosine triphosphate (ATP) and adenosine monophosphate (AMP) play crucial roles in the human body's energy transformation process 32,33 , thus usually prescribed for treating weight loss or dietary imbalance related to cancer. Cyclodextrin-phosphate (CD-P) was used as a surface coating to take advantage of our in-house non-covalent surface coating method 27 .

An array of techniques is used to evaluate the changes in nanoMOF structure and composition upon degradation. It includes standard pertinent characterization techniques for MOF, such as XRD, IR spectroscopy, gas adsorption, TEM, and SEM. Also, the chemical environment at the nanoMOF surface is studied by XPS. Furthermore, the metal coordination sphere of nanoMIL-100(Al) and nanoMIL-100(Fe) are probed by ssNMR and XANES spectroscopy. The degradation reaction kinetics is determined by quantitative liquid NMR spectroscopy. Finally, SAXS and NMR relaxometry probe the behaviors and interactions of MOF nanoparticles in their suspension. Here, NMR spectroscopy techniques gave valuable information for understanding the degradation process of nanoMOF.

The present chapter has four main parts:

1) Synthesis and characterization of nanoMIL-100(Al) and nanoMIL-100(Fe)

2) Investigation of the degradation mechanism of original nanoMIL-100(Al) and nanoMIL-100(Fe) at the molecular scale by analyzing the nature of degradation products and the alteration of MOF's coordination sphere

3) Study of the influence of drug encapsulation and surface modification on the degradation of nanoMIL-100(Fe) and (Al) at the molecular scale, which includes studying the interaction between nanoMOFs and the loaded drug/surface coating before the degradation 4) Study of how the size, morphology, and interparticle interaction of nanoMIL-100 suspension evolve during the degradation to get insights about the process at the nanoscale

Synthesis and characterization of nanoMIL-100(Al) and (Fe)

In the first step of this study, we synthesized nanoMIL-100(Al) and nanoMIL-100(Fe) and then characterized their properties relevant to a drug carrier.

Size, shape, and water stability

One of the most important criteria for a colloidal drug delivery system is to have uniform size distribution, which is the aim of the synthesis of nanoMIL-100 materials. NanoMIL-100(Fe) was reproduced from a published procedure 8 and had a mean hydrodynamic diameter of 226 ± 15 nm and polydispersity index (PdI) of 0.13 ± 0.03 (n = 11) (Figure 3-2). However, nanoMIL-100(Al) synthesized following another study 34 had an inhomogeneous size distribution (Figure 3-3A). An additional short centrifugation step was needed to remove the large-size particles. As a result, we obtained nanoMIL-100(Al) with a mean hydrodynamic diameter of 147 ± 14 nm and PdI of 0.10 ± 0.02 (Figure 3-3B). TEM images in Figure 3-4 also display the two nanoMIL-100 particles with uniform shapes and sizes.

. Magnification of (A-B) images is 10k, and of (C-D) images is 30k

Despite having similar monodisperse size distribution in EtOH, original nanoMIL-100(Al) and nanoMIL-100(Fe) behaves differently in deionized (DI) water. The original nanoMIL-100(Fe) suspension in water had an average diameter of around 200-250 nm with PdI around 0.2, only when dispersing nanoMIL-100(Fe) wet pellet in water with a ratio above 1 mg.mL -1 . Below this ratio, dispersing nanoMIL-100(Fe) in water resulted in particle aggregation, whose average size measured by DLS was higher than 1000 nm. Meanwhile, the average diameter of the original nanoMIL-100(Al) suspension in water is more than 1000 nm, no matter its concentration. These behaviors were more clearly observed in TEM images (Figure 3-5A to C). In the nanoMIL-100 suspension in EtOH and nanoMIL-100(Fe) suspension in water, particles were uniformly distributed (Figure 3-4 and Figure 3-5D to F). Globally, they have a faceted shape and a dimension of around 100 nm. However, in the nanoMIL-100(Al) aqueous suspension, these particles exist as agglomerates (low magnification TEM image in

Crystallinity, porosity, and functional group

Apart from the nanoparticle size in the suspension phase, the two MIL-100 analogs possess similar features of crystallinity, porosity, and functional groups. First, the as-synthesized nanoMIL-100(Al) and nanoMIL-100(Fe) have similar experimentally measured XRD patterns (Figure 3 -6), as expected from their analog crystal structure. Also, the position and the width of the main diffraction peaks are similar to those in the calculated XRD pattern obtained from MOF crystal (Figure 3456), assuring the crystallinity of the as-synthesized MOF. B) nanoMIL-100(Fe). The theoretical patterns were calculated using Mercury software and the crystal structure .cif files from reference 5,6 Regarding porosity, the two MIL-100 MOFs (Figure 34567) display a type-II N2 adsorption isotherm. Two uptakes at P/P0 of 0.05 and 0.12 were previously explained by the presence of two types of cages with different pore sizes and pore opening windows 5 . They also have high BET surface area; the average value for nanoMIL-100(Al) is 1632 ± 325 m 2 .g -1 (5 synthesis batches) and for nanoMIL-100(Fe) is 1621 ± 327 m 2 .g -1 (4 synthesis batches). Our obtained values agree with the previous works 8,34 , indicating that the activation step in the present synthesis was adequate. In addition, the IR spectra (Figure 345678) have several fingerprints of coordinated trimesate in MIL-100 structure, in accordance with the published spectra of MIL-100(Fe), MIL-100(Al), and trimesic acid. The assigned vibration bands and peaks are summarized in Table 3-1. For instance, the band centered around 3250 cm -1 corresponds to the stretching vibration of the carboxylate group's -OH. The peaks at 1625-1630 and 1380-1400 cm -1 were attributed to the stretching of carboxylate C=O and COO, respectively. Those located around 760-770 and 700-710 cm -1 represent out-of-plane vibration of C-H in trimesate's benzene ring. The stretching band of aromatic C-C was also found around 1450-1460 cm -1 . Notice in the IR spectrum of nanoMIL-100(Fe) that there is a low intense peak at 1715 cm -1 , which can correspond to the stretching band of carboxylic C=O of trimesic acid trapped inside the MOF pores even after the washing steps. Small peaks near 1265 cm -1 , like the stretching band of carboxylic C-OH of trimesic acid, were also detected in both nanoMIL100(Al) and nanoMIL-100(Fe). These observations support the presence of residual trimesic acid trapped in the pores of our synthesized nanoMIL-100(Al) and nanoMIL-100(Fe), most likely in marginal quantity. Wave number (cm -1 )

MIL100Al

MIL100Fe trimesic acid

The standard physiochemical properties of nanoMIL-100 are summarized in Table 3-2.

Chemical composition

The chemical composition of the synthesized nanoMIL-100(Al) and (Fe) was determined by TGA, elemental analysis, and ICP.

Summary

We successfully prepared nanoMIL-100(Al) and nanoMIL-100(Fe) suspensions with almost similar features: particle size, chemical structure, crystal structure, and porosity. The exception of different colloidal stability in water of the two analogs may affect their degradation processes.

Degradation mechanisms of original nanoMIL-100(Al) and (Fe) at molecular scale

In the second step of this study, we assessed any possible modification in nanoMOF's molecular structure upon its degradation reaction. NanoMIL-100(Al) or nanoMIL-100(Fe) was incubated in phosphate buffer saline (PBS), composed of phosphate monobasic and dibasic salts, at 37ºC to mimic the physiological medium. After a given degradation period, the suspension in PBS was centrifuged to separate the liquid supernatant, named "degraded supernatant", and the solid precipitate, named "degraded pellet''. These two recovered degradation products were characterized by a variety of techniques. First, the degradation products in the liquid phase and the reaction kinetics were monitored by quantitative liquidstate NMR (lqNMR) spectroscopy. Second, the crystallinity, porosity, and functional group of degraded MIL-100 were measured by X-ray diffraction (XRD), gas adsorption, and IR spectroscopy, respectively. Then, a semi-quantitative analysis using X-ray photoelectron spectroscopy (XPS) was performed to get details on the chemical environment and atomic composition at the surface of degraded nanoMIL-100. Finally, an in-depth investigation into the atomic scale structure of the degraded nanoMIL-100(Al) and (Fe) was conducted separately by solid-state NMR (ssNMR) spectroscopy and X-ray absorption near-edge structure (XANES) spectroscopy, respectively. A comparison of the degradation mechanisms of two MIL-100 analogs is given at the end.

Reaction kinetics at liquid phase

LqNMR spectroscopy detects trimesate in the nanoMIL-100's liquid degradation products.

The typically obtained 1 H NMR spectra of the degraded supernatant consist of signature aromatic proton peaks of trimesate/trimesic acid in the region 8. 3-8.5 ppm (Figure 3-10), in agreement with the literature that nanoMOF degradation leads to formation and release of free ligand. Notably, in Figure 3-10, the 1 H NMR spectrum of nanoMIL-100(Fe)'s degraded supernatant has only one narrow peak at 8.37 ppm assigning to trimesate; trimesic acid has only one peak of aromatic protons, too 40 . However, the 1 H NMR spectrum of the aluminum MOF's degraded supernatant contains not only a similarly narrow peak at 8.44 ppm, but also several additional small peaks in the region from 9 to 8.5 ppm. The intensity of these peaks is roughly 4-5 times lower than that of the principal peak. Still, they are positioned within the region of aromatic protons and were detected in all nanoMIL-100(Al) degraded samples. Thus, they are also assigned to trimesate species in nanoMIL-100(Al)'s degraded supernatant. The origin of these additional trimesate peaks is not clear yet.

Considering the main 1 H NMR signal of trimesate in nanoMIL-100's degraded supernatant, as a function of degradation time, its chemical shift negligibly deviates by maximum 0.01 ppm, and its intensity increases. By quantitative NMR spectroscopy, the peak area of trimesate can be used to monitor the rate of trimesate formation, hence the kinetics of MOF degradation reaction (details in section 3.8.6). In the case of MIL-100(Al), the area of the additional small peaks was also considered. The kinetics of trimesate formation during nanoMIL-100 degradation, determined by quantitative NMR spectroscopy, is shown in Figure 3-11. The percentage of released trimesate was calculated as the amount of trimesate in the degraded supernatant (measured by lqNMR) divided by the initial trimesate content (determined by MOF digestion, for details, see 3.8.6). Different concentrations of nanoMOF (0.25, 0.5, and 10 mg.mL -1 ) and PBS (1 and 10 mmol.L -1 ) were used, resulting in varying phosphate-to-metal molar ratios (P/M, where M is Fe or Al) of 0.3, 0.5, 5.5, and 11. Figure 3-11 shows that globally, the profiles of trimesate release from both nanoMIL-100 consist of two stages: fast and gradual trimesate release. The first stage happened in the first three hours after incubation in PBS, when there was a rapid generation of free trimesate. Subsequently, the degradation reaction rate gradually decreased until it reached equilibrium after 24-48 hours. This plateau of trimesate release after a given incubation period in PBS was also observed in other works dealing with MIL-100(Fe) 9,13,14 . Unlike the incubation in phosphate buffer, the control incubation of nanoMIL-100 in water produced a negligible amount of trimesate, less than 2% in the case of nanoMIL-100(Al) and non-observable in nanoMIL-100(Fe), once again stressing the reported role of phosphate in inducing trimesate release from nanoMIL-100 MOFs 28 . In this study, we precise that it is the phosphate-to-metal molar ratio that decides the extent of nanoMIL-100 degradation. Figure 3-11 shows that the higher the P/M ratio, the higher the maximum proportion of trimesate released from nanoMIL-100(Fe) and nanoMIL-100(Al). Another noteworthy observation from Figure 3-11 is that the magnitude of % trimesate release, hence the magnitude of degradation of nanoMIL-100(Al), is higher than that of nanoMIL-100(Fe). For example, at a very low P/M ratio (0.3), trimesate release from nanoMIL-100(Al) attained 31%, nearly twice as much of that from nanoMIL-100(Fe) (15%). Also, the iron-based MIL-100 has a more progressive profile of trimesate formation than its aluminum analogues, especially at high degradation ratio P/M. (Al). Several phosphate-to-metal (P/M) molar ratios, calculated from PBS and nanoMOF concentrations, were used, namely: very low (0.3), low (0.5), moderate (5.5), and high (11). Information on the calculation of these ratios is in section 3.8.5

In addition, the constant phase in trimesate release could be due to insufficient phosphate to react with nanoMIL-100, since the higher the P/M ratio, the more trimesate was liberated. Also, PBS was consumed more in degradation with lower P/M. When the P/M ratio was raised from 0.3 to 0.5, the consumed amount of PBS dropped from 83% to 35%, as indicated in the 31 P NMR spectra of nanoMIL-100(Al)'s degraded supernatant (Figure 3 -12). Nonetheless, there is still PBS remaining while less than 100% trimesate was released, which argues for a more complex mechanism preventing the reaction from proceeding further. Briefly, the qualitative and quantitative liquid NMR spectroscopy data unambiguously evidence the connection between the presence of phosphate in the degradation medium and the trimesate release from nanoMIL-100 materials. Therefore, inorganic phosphate ions in PBS are directly involved in the nanoMIL-100 degradation.

Changes in crystallinity, porosity, and functional group

To further elucidate the mechanisms behind the formation of degraded MIL-100, characterization of the solid-phase reaction products was conducted, and the obtained data are described in the subsequent text. High degradation molar ratio P/M ( 11) was employed to ensure the reaction produces the maximum release of trimesate, hence the most degraded solid-phase products. The low and very low P/M ratios (0.5 and 0.3, respectively) were also used to study the intermediate status of the degradation reaction. Several reaction times were used to assess the time evolution of the reaction. The degraded was considered equilibrium after 48 hours of MOF incubation in PBS for all P/M ratios.

The long-range crystal structure of degraded nanoMOF was evaluated by powder XRD. In Figure 3-13, the XRD pattern of 48-hour degraded nanoMIL-100(Al) at a low P/Al ratio did not differ much from the pristine material. At this degradation ratio, XRD suggests that the degraded nanoMIL-100(Al) kept its long-range crystal structure despite a considerable loss of 42% initial trimesate content (see Figure 3-11B). In another work on nanoMIL-100(Fe) degradation, XRD showed that the degraded sample that experienced 100% trimesate loss became amorphous 30 .

Figure 3-13. XRD patterns of 48-hour-degraded nanoMIL-100(Al) at a low P/Al ratio (0.5) compared with the pristine material However, the same degraded nanoMIL-100 underwent a dramatic reduction in its porosity. When using the same P/M ratio as in the XRD experiment, the N2 uptake of degraded nanoMIL-100(Al) and nanoMIL-100(Fe) remarkably decreased (Figure 3-14). Zoom in the adsorption isotherm shows that the steps in gas uptake near P/P0 at 0.05 and 0.12 became hardly discernible. Possible reasons include restricted access to the MOF pores, assuming

Chemical environment and atomic composition at nanoMOF surface

To gain a more detailed understanding of the nanoMOF degradation, we used XPS to investigate the chemical environment at the nanoMOF surface, which is expectedly essential for any reaction of nanoparticles. Original and degraded nanoMIL-100 suspension of 10 mg.mL -1 in water were deposited on silicon wafers and left to dry overnight. Note that the stated concentration of degraded nanoMIL-100 suspension is the concentration of the starting suspension used for the degradation reaction. The remaining content of MOF in the sample after the degradation process was unknown.

Firstly, the original nanoMIL-100(Al) and nanoMIL-100(Fe) were studied. Their core-level XPS spectra in Figure 3-16 have the signatures of trimesate ligand and metal cations in nanoMIL-100 structure. Their deconvoluted C 1s spectra (Figure 3-16A, C) show the peaks of C carboxylate and C aromatic at 284.8 and 288.5 eV, respectively (the choice of deconvolution is explained in section 3.8.6). Also, the Al 2p and Fe 2p core levels spectra of MIL-100(Al) and MIL-100(Fe) in Figure 3-16B, D have their peak shape and binding energy position consistent with a (+III) oxidation state and a metal-oxygen bonding. More specifically, the Al 2p peak in original nanoMIL-100(Al) is centered at 74.6 eV, similar to the typical binding energy value for alumina (Al2O3) 42 . The Fe 2p spectrum of nanoMIL-100(Fe) has a doublet of Fe 2p1/2 and Fe 2p3/2 at 725 and 712 eV, respectively. This doublet's shape and binding energy are close to iron(III) oxide (Fe2O3) 43,44 . The deviation in the biding energy of Fe 2p3/2 is related to the difference in ligand electronegativity 44 . By comparison, a clear doublet was witnessed for Fe 2p in the iron MOF, but a broad signal was seen for Al 2p in nanoMIL-100(Al) because the spin-orbit splitting for Fe 2p is much larger than for Al 2p. Additionally, the survey spectrum of nanoMIL-100(Fe) contains a signal of Cl at 200 eV, and that of MIL-100(Al) contains a signal of N at 406 eV (annotated in Figure 3-17). These signals come from residual synthesis precursors, namely FeCl3 for nanoMIL-100(Fe) and Al(NO3)3 for nanoMIL-100(Al). Furthermore, XPS measurement of degraded nanoMIL-100 revealed considerable changes in the coordination sphere of nanoMIL-100 after degradation. First, the C 1s spectra of degraded nanoMIL-100 (Figure 3-18) illustrate a decrease of C 1s peaks representing trimesate carboxylate and sp 2 C after 48 hours of reaction with PBS, even total disappearance of these peaks in the case of degraded nanoMIL-100(Al). Second, the P 2p core level spectrum (Figure 3 The XPS data presented above clearly evidence the disappearance of trimesate ligands and the appearance of phosphate in the degraded MIL-100 pellets, which is entirely consistent with the increase of trimesate concentration and decrease of phosphate concentration in the degraded supernatant shown by lqNMR spectroscopy (section 3.3.1). Since XPS offers semi-quantitative atomic composition, we can correlate the C/M (carbon-to-metal) and P/M (phosphorous-to-metal) ratios in the degraded pellets' surface (XPS's results) and % trimesate release in the degraded supernatant (lqNMR's results). While the XPS technique provides atomic composition at the material's surface, one can also estimate the atomic composition in the bulk material by energy dispersive X-ray spectroscopy (EDS). EDS's penetration depth is 1-2 mm, high enough to probe the bulk composition of 100-nm-large MIL-100 particles. The P/M ratio in the degraded solid determined by EDS is plotted versus the % trimesate release in Figure 3-22. At first look, the magnitude of the P/M ratio determined by EDS is lower than that determined by XPS.

Recalling that XPS surveys only 10 nm depth of the material, which is 100 times shorter than EDS and corresponds to the surface region of nanoMIL-100. For nanoscale substances, it is expected that the reaction takes place more intensely at the surface than in bulk. Another important observation from Figure 3-22 is that with a rising proportion of trimesate release, the bulk P/Al increases gradually, but the bulk P/Fe ratio increases more slowly and seems to reach a limit when the % trimesate release approaches 100%. At a high level of trimesate release 60-80%, the bulk P/M ratio of degraded nanoMIL-100(Fe) is also lower than its counterpart in degraded aluminum MOF. Possibly, the degradation reaction propagates deeper and takes place more uniformly in nanoMIL-100(Al) than in nanoMIL-100(Fe) particles. However, the reaction at the surface experiences an opposite trend. The surface P/Fe ratio was higher and increased faster than the surface P/Al ratio, implying that the surface reaction was faster for the iron MOF than its aluminum analog. 3-22. Correlation between the P/M ratio determined by EDS (bulk) and by XPS (surface) and % trimesate release in degraded nanoMIL-100(Al) and nanoMIL-100(Fe)

Discussion on the ligand exchange process

All the experimental evidences described so far suggest that the degradation reaction of nanoMIL-100(Al)/nanoMIL-100(Fe) in PBS is driven by ligand exchange between their native trimesate ligand and phosphate inorganic ions. Generally, the coordination bond in MOF is 0 20 40 When using phosphate-containing degradation media like PBS pH 7.4 in our case, it is essential to specify the type of phosphate species and their proportion at equilibrium at a given pH, as phosphoric acid H3PO4 is a polyacid. Figure 3-23 displays the calculated distribution of phosphate species as a function of pH (Figure 3 -23). At pH 7.4, phosphate buffer contains 61% of dibasic phosphate ions (monohydrogen phosphate HPO4 2-) and 39% of monophosphate ions (dihydrogen phosphate H2PO4 -). The phosphate composition at equilibrium does not mimic the input ingredients, 85% Na2HPO4 and 15% KH2PO4 (media formulation [START_REF] Cierkowska | Building pharma pipelines using a socioeconomic lens[END_REF] ). A previous study has shown better stability of the MIL-100(Fe) at pH 5.4 than at 7.4 when the dihydrogen phosphate anion is the dominant species 9 (98% H2PO4 -and 2% HPO4 2-according to Figure 3-23). Therefore, the active species responsible for nanoMOF degradation at near neutral pH is more likely the monohydrogen phosphate anion HPO4 2-, which can act as a nucleophilic substituting agent towards coordination complexes 51 .

Figure 3-23. Distribution of phosphate species as a function of pH calculated with ChemEQL software using phosphoric acid dissociation constants pKa1 = 2.15, pKa2 = 7.20, and pKa3 = 12.35 The capability of phosphate anions to replace the polycarboxylate trimesate linkers in MOF is manifested in several aspects. First, their negative charge makes HPO4 2-and H2PO4 - attractive to electrophilic metal coordination centers in MOF. High affinity of phosphate towards various metal cations in living organisms has been reported 52 . Also, the charge distribution of phosphate is smaller than that of trimesate anions, thus making phosphate more advantageous for a nucleophilic agent. It was the same reason that makes I -anion a weaker metal complexing agent than F -anion 53 . Second, both pKa2 and pKa3 of phosphoric acid (7.20 and 12.35, respectively) are higher than all three pKa of trimesic acid (3.12, 3.89, and 3.40 54 ). Therefore, the corresponding conjugated base of phosphoric acid, HPO4 2-and H2PO4 -, have higher basicity, hence higher nucleophilicity towards metal 53 . The third reason is related to the stability of trimesate as a free anion. Generally, carboxylate is not the best leaving group for a nucleophilic substitution reaction. However, in trimesate anion, the negative charge of the carboxylate groups can be stabilized by the electron delocalization in Furthermore, the characterization of both liquid-and solid-degradation products consistently point out that the reactions of nanoMIL-100(Al) and nanoMIL-100(Fe) with phosphate exhibit different extent and rate. The release of trimesate during nanoMIL-100(Al)'s degradation is faster and has a higher magnitude, and the bulk P/M ratio in the degraded MIL-100(Al) pellet is also higher. It appears that nanoMIL-100(Al) is more prone to degradation by phosphate than its iron analogs. This difference in reactivity can be explained by the metal-ligand binding strength between Al/Fe(III) and carboxylate/phosphate. Several authors have reported that Fe(III) has a higher complexing affinity towards phosphate and several carboxylate ions than Al(III) 55,56 . With respect to aluminum, iron favors bonding with phosphate but, at the same time, does not favor departure/dissociation with trimesate. This push-and-pull affinity plausibly makes the release of trimesate, hence the degradation reaction of the iron MOF more progressive than that of the aluminum analog. Granted, gradual ligand release also occurred with nanoMIL-100(Al) upon its degradation at the P/Al ratio of 0.3, but that was when the amount of phosphate was insufficient for reacting with all aluminum. The aluminum MOF readily reacted with phosphate when the P/Al ratio was raised to 5.5. Also, the difference in reactivity towards phosphate of Fe(III) and Al(III) might explain the heterogeneity of the reaction in nanoMIL-100 material. As phosphate is more attracted to Fe(III) than Al(III), the degradation initiates at the nanoMIL-100 surface faster for the ironbased nanoMOF than the aluminum-based, leading to a higher surface P/Fe than P/Al ratio. Nevertheless, as the reaction continues and penetrates the NPs, the cleavage of trimesate was more preferred in nanoMIL-100(Al) than in nanoMIL-100(Fe). Consequently, the bulk of nanoMIL-100(Al) is degraded more easily than the bulk of nanoMIL-100(Fe), resulting in the gradual increase of the bulk P/Al ratio. Contrary to the homogenous degradation in nanoMIL-100(Al), the process is limited in the bulk nanoMIL-100(Fe), thereby slowing down and limiting the increase in the bulk P/Fe ratio.

Evolution of metal coordination sphere in nanoMIL-100(Al)

Although we have obtained substantial proofs of the ligand exchange/substitution underlying the nanoMIL-100 degradation. Nonetheless, we still do not know much about what MIL-100 becomes after the liberation of its trimesate constituting ligands. Apart from the hint of IR spectroscopy, there is no direct data about the coordination sphere of MOF's metal. For getting this information from the original and degraded MIL-100 materials, solidstate NMR (ssNMR) and X-ray near-edge structure (XANES) spectroscopy are relevant complementary techniques. Both provide non-destructive analysis of the local environment of MOF materials. SsNMR spectroscopy was used to study the diamagnetic nanoMIL-100(Al), and XANES spectroscopy was used for the paramagnetic nanoMIL-100(Fe), described in the two subsequent sections 3.3.5 and 3.3.6.

Firstly, the chemical structure and metal's coordination sphere in the original nanoMIL-100(Al) were characterized using ssNMR spectroscopy and the reference study on MIL-100(Al) microcrystal by Volkringer et al. 1 H and 13 C MAS NMR experiments show signatures of the trimesate ligand of nanoMIL-100(Al). 1 H MAS NMR spectrum (Figure 3 -25) shows two peaks at 9.1 and 8.4 ppm corresponding to aromatic protons of coordinated trimesate ligand. In this same region of aromatic protons, another signal at 8.0 ppm was assigned to the excess trimesic acid in the MIL-100 framework 6,35 . Likewise, the 13 C MAS NMR spectrum (Figure 3-26) consists of six peaks in the range 130-140 ppm and two peaks in the range 170-180 ppm attributed to the aromatic carbons and carboxylate carbons of trimesate, respectively. Notice that the trimesate ligands in nanoMIL-100(Al) have more nonequivalent NMR peaks than expected from the trimesic acid's chemical structure. In a solidphase and crystal MOF sample, ligands bonded with metal clusters and cannot freely rotate, thus creating several slightly different bond orientations. In the case of MIL-100(Al), the chemical environments of individual carbons are further modified and differentiated as there are different crystallographic aluminum sites 6 . As a result, nanoMIL-100(Al) with good crystallinity (Figure 3-6A) has distinctive carbon NMR signals like the microparticles 6 . Additionally, the 27 Al MAS NMR spectrum in Figure 3-27A shows signals in the range from 5 to -10 ppm, the signature chemical shift range of 6-coordinate aluminum 57 , which fits the reported crystal structure of nanoMIL-100 MOF (see Figure 3-1). In addition to two apparent signals at 3.2 and 1.6 ppm, there is a broad and elongated "tail" around -5 ppm, a feature of quadrupolar 27 Al NMR spectra. A way to remove the quadrupolar broadening and better resolve these peaks is the multi-quantum MAS (MQMAS) NMR measurement, which gives a pseudo-2D correlation map of these different sites with themselves according to their second-order quadrupolar parameters. The 27 Al MQMAS spectrum and its extractions in Figure 3-27B show four 6-coordinate aluminum species. Using the parameters obtained from fitting the MQMAS spectrum, these four 27 Al resonances were also deconvoluted from the 1D NMR spectrum (Figure 3-27A). The three lines at the isotropic position of 3.3, 1.6, and -8.6 ppm are consistent with the values from the study of Volkringer et al. about micro-MIL-100(Al) 6 , but they are less resolved in our MQMAS spectrum presumably due to the nano size of the MOF particles. These peaks were previously attributed to different nonequivalent crystallographic Al sites in MIL-100(Al) crystal structure. However, Figure 3-27B shows the fourth species at -3.8 ppm, which was not observed for MIL-100(Al) microparticles. Thus, it was assigned to the surface aluminum sites characteristics of MIL-100(Al) nanoparticles, as reported previously by some of us 26 . These surface species were reproduced in this study and are expected to play an active role in the degradation of nanoMIL-100(Al). The first observation is the broadening of the spectrum around -10 ppm, which in 27 Al NMR hints at the presence of a new peak. Also, the 27 Al MQMAS NMR spectrum witnessed a development in the region of surface Al species (Figure 3-29A, B), suggesting degradation at the nanoMIL-100(Al) surface creates new aluminum species. These species with the chemical shift range from -2 to -10 ppm could be 6-coordinate Al coordinated with at least one to two phosphate ions 58,59 created from the reaction between the MIL-100(Al) metal cluster and PBS's phosphate. The formation of the Al-O-P bond was further evidenced by the 31 P-27

Al MAS D-HMQC , which showed a through-space correlation peak between the 27 Al resonance at -2 ppm and the 31 P resonances around -10 ppm. This 31 P- 27 Al correlation peak indicates close spatial proximity within 5 Å between these two nuclei, which fits the Al-O-P bond distance 61 . H NMR peak at 2.9 ppm was found for gibbsite mineral of aluminum hydroxide 63 ). Then, due to an increase in interaction between the Al metal center and the newly bonded phosphate ligand, the interaction of the Al center with the carboxylic groups of trimesate ligands can weaken. As a result, phosphate can also attack and replace carboxylate's position, evidenced by lqNMR and XPS, as previously discussed in sections 3.3.1 and 3.3.3. Free trimesic acid is liberated when all 6 Al-O-C bonds between a trimesate linker and Al trimers are disconnected.

Furthermore, the 1D 27 Al MAS NMR spectrum of the degraded nanoMIL-100(Al) (Figure 3-28B) shows another newly formed Al species at 48 ppm, whose chemical shift lies between the range of 4-and 5-coordinate aluminum cations. A decrease in the coordination number of aluminum cations in MIL-100(Al) is a known phenomenon observed when heating the sample above 150 o C. As Haouas et al. 35 , we also found that the coordinated water ligands were removed by heating, thus creating 5-coordinate Al cations with a 27 Al chemical shift of 34 ppm (Figure 3 -28). The higher the temperature, the more 5-coordinate Al was transformed from 6-coordinate Al (Figure 3-33A). When the dehydrated MIL-100(Al) readsorbed water from ambient air, this 5-coordinate Al cation returned to 6-fold coordination without modifying the framework. This 5-coordinate Al was also created when heating the degraded MIL-100(Al), still detected at 34 ppm (Figure 3-33B). However, the chemical shift of the degraded low-coordinate Al remained at 48 ppm when the sample was exposed to air or heated above 150 o C, implying that this low-coordinate Al site no longer coordinates with water ligands. As Al(III) species bonded with phosphate usually have lower 27

Al chemical shifts than the same coordination number Al(III) species with no phosphate ligand due to a deshielding effect 57 , we hypothesize that these Al species at 48 ppm are 4coordinate Al with at least one phosphate ligand. The presence of tetracoordinate aluminum phosphate with a minor proportion alongside hexacoordinate aluminum phosphate has been reported in the literature 64,65 . Their corresponding 27 Al MAS NMR chemical shifts also match what we observed in the present study. Moreover, when degrading nanoMIL-100 with a higher P/Al ratio (0.5), the 4-coordinate 27 Al peak intensity increased slightly, and even a new 4-or 5-coordinate 27 Al peak was observed around 22-25 ppm (Figure 3-34). Those 4-and 5-coordinate aluminum phosphate may represent intermediate reaction products. These species with reducing aluminum's coordination number can be formed during the partial dissociation of the aluminum-trimesate coordination bond. So far, the ssNMR results of nanoMIL-100(Al) agree with the previous hypothesis about nanoMIL-100(Fe) 29 that aluminum phosphate/iron phosphate complex is the degradation product of nanoMIL-100 reacting with phosphate buffer PBS. While Li et al. could only hypothesize the formation of iron phosphate and iron oxide hematite and goethite 29 , the present study offers concrete evidence about the structure of the degraded nanoMIL-100(Al) by showing the coordination state of Al and the likely formation of Al-O-P bond as the main degradation product.

3.3.6 Evolution of metal coordination sphere in nanoMIL-100(Fe)

To study the paramagnetic MIL-100(Fe), which is not suitable for ssNMR spectroscopy, in the scope of a collaboration with synchrotron SOLEIL, we used X-ray near edge structure (XANES) spectroscopy to probe its coordination structure before and after the degradation reaction in PBS. To maximize our chance in synchrotron experiments to observe the evolution of iron's coordination sphere, nanoMIL-100(Fe) degradation was carried out with a high P/Fe ratio (11). Therefore, XANES spectroscopy reported here presumably detected species in the final stage of the nanoMOF degradation reaction.

The obtained X-ray absorption spectra were normalized so that the pre-and post-edge lines were at 0 and 1, respectively (Figure 3 -35). The following section focuses on the pre-edge and near-edge structures where we witnessed a more apparent evolution of XANES spectra. Likewise, the XANES spectrum of degraded nanoMIL-100(Fe) was also compared with other reference iron compounds (Figure 3 -37 and Figure 3-38). The first observation from Figure 3-37 is that the oxidation number of Fe in nanoMIL-100(Fe) does not change and remains as (+III), as the degraded and original nanoMIL-100(Fe) have the same edge position, and the edge position of nanoMIL-100(Fe) is close to the values of Fe(III) compounds. This preservation of iron's oxidation state agrees with the XPS results described above and a previous publication using Mossbauer spectroscopy 29 , which is consistent with our proposed degradation mechanism of ligand substitution. More significantly, in Figure 3-38, the XANES Fe K-edge spectrum of the degraded nanoMIL-100(Fe) is reminiscent of the spectrum of iron(III) phosphate dihydrate, a substance having 6-coordinate iron(III) bonds with phosphate through Fe-O-P bonds 67 . This similarity is also observed in the pre-edge region, where the intact MIL-100(Fe) shows two features at 7114 and 7115.4 eV, while the degraded MIL-100(Fe) and FePO4.2H2O only show a single, broad peak at 7114.9 eV. Moreover, this degraded nanoMIL-100(Fe) has a distinctive XANES shape from other probable degradation products such as iron(III) oxide and iron(III) oxyhydroxide 29,, thus excluding their presence in the degraded nanoMIL-100(Fe). In addition, the transformation from intact nanoMIL-100(Fe) to this 6-coordinate iron phosphate was observed when varying the input P/Fe ratios (Figure 3 -40). These XANES spectra of corresponding degraded MIL-100(Fe) were simulated by combining the spectra of original MOF and iron(III) phosphate dihydrate with fitted coefficients. The results of this linear combination analysis show that the XANES spectrum of the sample degraded at a moderate input P/Fe ratio (5.5) has a 42% contribution of the FePO4.2H2O (6-coordinate Fe-O-P). In contrast, the XANES spectrum of the sample degraded at a very low ratio (0.3), which looks like that of the original nanoMIL-100(Fe), is composed of 30%. Based on these XANES data, we propose that the 6-coordinate iron phosphate is the final principal degradation product of nanoMIL-100(Fe) when the P/Fe ratio is sufficiently high, e.g., 11. Its formation is likely to result from phosphate substitution of coordinated water and trimesate ligands in the same manner as nanoMIL-100(Al) degradation reaction. At the low P/M (0.5), nanoMIL-100(Fe), like its aluminum analog, also became less porous after degradation (Figure 3-14B) but with a smaller decrease (roughly 50%) in the gas uptake than in nanoMIL-100(Al) (roughly 90%). Reduction in porosity could suggest a passivation effect of iron phosphate on nanoMIL-100(Fe) surface like in degraded nanoMIL-100(Al) but to a lesser extent.

Discussion on the degradation mechanism of nanoMIL-100 in phosphate and possible degradation products

Based on all the results presented so far, we propose for the first time a reaction mechanism of nanoMIL-100(Al) and nanoMIL-100(Fe) degradation in the simulated biological medium phosphate buffer saline (PBS) as follows (Figure 3-41):

1) Phosphate ions in PBS, mostly the hydrophosphate dianinons HPO4 2-attack the nanoMIL-100 surface, substitute the labile water ligands, and form coordination M-O-P bond (M = Al, Fe). In consequence, the M-O-C coordination bonds between MIL-100(M) trimer and trimesate linker are weakened.

2) Phosphate ions substitute the carboxylate of trimesate linkers. As the substitution of water and trimesate continues, the formation of the 6-coordinate aluminum/iron bonded with 1 to 2 phosphates continues, producing the main degradation products 6-coordinate aluminum phosphate/iron phosphate. Free trimesic acid releases when all six M-O-C bonds between a trimesate linker and Al/Fe trimers are broken. During the breaking of metal-trimesate bonds and cleavage of trimesate ligands, some 4and/or 5-coordinate aluminum phosphate are created.

Then, depending on the input P/M (M = Fe or Al) ratio, the degradation reaction of nanoMIL-100 can follow two directions 3A) When the P/M is insufficient to cause 100% trimesate release, after 3-6 hours, the newly formed 6-coordinate aluminum phosphate/iron phosphate at the nanoMIL-100 particle's surface passivate the MOF pore and pore window. The intact metal trimers of nanoMIL-100 are no longer exposed to a high amount of phosphate ions. As a result, the degradation reaction, hence the release of free trimesate, slows down and reaches an equilibrium.

3B) In the extreme possibility when the P/M ratio is sufficiently high to disconnect all trimesate ligands from the MIL-100 framework, phosphate could react with the Al/Fe trimer faster than the passivation process. The final nanoMOF degradation products could be phosphate or hydroxyphosphate of six-and four-coordinate aluminum 64,69,70 /iron 67,71,72 . The two analogs MIL-100(Al) and MIL-100(Fe) are likely to have similar degradation reaction steps but react with phosphate at different rates and extents. At the molecular scale, the reaction between MIL-100 and phosphate is governed by the ligand exchange/ligand substitution process between phosphate and MIL-100's trimesate, hence the new bond formation between metal cations and phosphate anions and the bond breaking between metal cations and carboxylate anions. On the one hand, Fe cations have a stronger affinity with phosphate than Al cations; thus, the Fe-O-P coordination is thermodynamically more favored than the Al-O-P coordination. On the other hand, Fe cations have a stronger affinity with carboxylate than Al cations; therefore, it is more difficult to detach trimesate ligands from Fe cations than from Al cations. Moreover, the reactivity and kinetics of Fe(III) and Al(III) for nucleophilic substitution could relate to their d-orbital configuration. Generally, the ligand substitution of coordination metal complexes has three main mechanisms: dissociative, associative, or a combination of both 73 . If we apply this concept to MOF degradation, in dissociative substitution, nanoMIL-100 loses the trimesate ligand first, then reacts with the incoming phosphate. Meanwhile, in the associative mechanism, the incoming phosphate attacks the Al/Fe trimer, forming an intermediate. Afterward, this intermediate loses the trimesate ligand. Regarding nanoMIL-100(Fe), its Fe(III) with partly filled 3d 5 orbitals could resist the change from octahedral six-coordination to seven-coordination of any geometry 51 . Therefore, the iron MOF could favorably experience a dissociative ligand substitution mechanism. Contrastingly, nanoMIL-100(Al) with an empty 4d orbital may undergo both dissociative and associative substitution mechanisms. Having more pathways for ligand substitution could make the aluminum MOF react faster with phosphate than its iron analog.

Summary

In a nutshell, an arsenal of techniques has uncovered the degradation mechanism of nanoMOF in phosphate media driven by phosphate substitution of nanoMOF's native ligands, water and trimesate. The liberation of free trimesate and the exchange with phosphate were demonstrated by lqNMR spectroscopy and XPS. The difference in kinetics and extent of this ligand exchange process in nanoMIL-100(Al) and nanoMIL-100(Fe) particles were shown by lqNMR, XPS, and EDS. Furthermore, the coordination of metal with phosphate to form M-O-P bonds was first suggested by IR spectroscopy and later consolidated by ssNMR and XANES spectroscopy. 27 Al and 31 P NMR and their correlation measurement identified the most likely degradation products of nanoMIL-100(Al), 6-and 4coordinated aluminum phosphate. The similar transformation of nanoMIL-100(Fe) to 6coordinated iron phosphate was proved by Fe K-edge XANES spectroscopy.

Degradation and drug release mechanisms of modified nanoMIL-100(Al) and (Fe)

In this section, we investigated the impact of drug encapsulation and surface coating on the nanoMIL-100(Al) and nanoMIL-100(Fe) degradation processes. Literature has shown that the presence of loaded drug and surface coating can alter the physicochemical properties of nanoMOF 18,27 , thus potentially influencing nanoMOF's degradation reaction. NanoMIL-100(Al) and (Fe) were loaded with adenosine triphosphate (ATP) and adenosine monophosphate (AMP). Studying these two drugs belonging to the same family but having a different number of phosphate functional groups could help to reveal the effect of metalligand complexation strength on drug delivery processes such as drug encapsulation, drug release, and carrier degradation. Additionally, the surface coating with cyclodextrinphosphate (CD-P) was investigated to control the carrier stability/degradation without preventing drug release. The chemical structures of the two active molecules and the coating molecule are given in Figure 3-42. Firstly, we optimized the formulation for encapsulating ATP and AMP in nanoMIL-100(Al) and nanoMIL-100(Fe) to achieve the best drug payload wt% and encapsulation efficiency (EE%) while ensuring structure integrity of nanoMIL-100. We played with several values of input/theoretical drug loading in the range 20-35 wt% and two values of nanoMIL-100 suspension concentration: 0.5 and 5 mg.mL -1 . At the end of the drug loading process (details in experimental section), the suspension of nanoMOF and drug was separated into the supernatant containing unloaded drug and pellet of drug-loaded MOF. The amount of drug ATP/AMP and trimesate in the supernatant was measured by high-performance liquid chromatography (HPLC). The amount of drug encapsulated in nanoMOF was calculated by subtracting the amount of drug in the supernatant from the input quantity. Then, the drug payload wt% and encapsulation efficiency (EE%) were deduced from the ratio of the loaded quantity over the nanoMOF weight and the input quantity, respectively. The % trimesate loss (%BTC loss) is the percentage of trimesate measured in the supernatant out of the initial trimesate content in the drug loading suspension. These three figures of merits of ATP and AMP drug encapsulation in nanoMIL-100 are given in Table 3-4 andTable 3-5, respectively. Regarding ATP drug loading, using 5 mg.mL _1 MOF led to better ATP payload and EE% than 0.5 mg.mL -1 , especially in nanoMIL-100(Al) (Figure 3 -43). However, no matter of input ATP loading wt% and MOF concentration, nanoMIL-100(Fe) always gave higher ATP payload and EE% (≈ 100%) than nanoMIL-100(Al). Possibly because Fe 3+ cations create a more stable complex with triphosphate than Al 3+ cations 56 , MIL-100(Fe) forms a more stable complex with triphosphate than MIL-100(Al), hence higher affinity towards capturing ATP. In addition, among the studied input drug loading levels, high input drug quantity (30-35 wt%) causes a decrease in EE%, except for the formulation 5 mg.mL -1 nanoMIL-100(Fe). Notably, HPLC detects a non-negligible amount of trimesate in the supernatant recovered from the drug loading suspension of nanoMIL-100 and ATP (Figure 3 -44). This 10-20% leakage of trimesate after drug loading could be because ATP, a strong nucleophile with a triphosphate group, replaces trimesate during the drug loading process. The higher the ATP loading, the higher the %BTC loss is. Total maintenance of trimesate content, i.e., zero trimesate loss, likely requires very low input drug loading below 10 wt% 25 , which is not preferred in this case. Therefore, the optimal ATP loading formulation to compromise drug payload, EE%, and %BTC loss could be 25 wt% input ATP and 5 mg.mL -1 of nanoMIL-100. Regarding AMP drug loading (Table 345), raising MOF concentration from 0.5 to 5 mg.mL _1, like ATP drug loading, led to considerably better drug payload and EE%. However, unlike ATP drug loading, the two nanoMIL-100 experienced the same increase in AMP loading performance (Figure 3-45). They also have close values of AMP drug payload and EE%, hence similar affinity towards AMP. Taking the same compromise between drug payload, EE%, and %BTC loss, the AMP loading formulation could be 25 wt% input AMP and 5 mg.mL -1 of nanoMIL-100. It is also convenient to use one formulation of input drug loading and nanoMOF concentration for both ATP and AMP for comparing % drug release in the future. Looking again at the optimization results, we witness a leakage of trimesate ligand of nanoMIL-100 carrier upon both AMP and ATP encapsulation (Figure 3 -44). At a given MOF concentration, the % trimesate loss upon AMP encapsulation (8-16%) is lower than upon ATP encapsulation (9-20%) (Figure 3 -46). As phosphate is more nucleophilic than carboxylate group (see section 3.3.4), both ATP and AMP can act as nucleophilic substituting agents and replace trimesate ligands when mixing ATP/AMP with nanoMIL-100. However, AMP with one phosphate group has weaker interaction with the metal cluster than ATP with three consecutive phosphate groups, thus, at the same drug concentration, having a lower tendency to attack trimesate ligand 74 . Note that with the chosen formulation of 5 mg.mL -1 MOF, the drug payload values for AMP and ATP of each MIL-100 analog are similar. Chapter 3 Studying degradation mechanism of metal-organic framework nanocarriers

Drug-carrier interaction between ATP/AMP and nanoMIL-100(Fe)/(Al) carriers

This section explores the host-guest interaction between the ATP and AMP cargo and the nanoMIL-100 carrier by characterizing the physiochemical properties and the chemical structure of the drug-loaded nanoMIL-100.

Firstly, the presence of ATP and AMP inside nanoMOF influences their suspension's physicochemical properties differently. Their size distribution measured by DLS and zetapotential is given in Table 3-6. The two ATP-loaded nanoMIL-100 suspensions had an NP diameter of around 200-300 nm with a polydispersity index of around 0.2-0.3. The present study observed that the original nanoMIL-100(Al) always agglomerates in DI water, while it was previously published that the original nanoMIL-100(Fe) in water aggregates over time 27,75 . It appears that ATP encapsulation positively affects the carrier's colloidal stability and even significantly increases it in the case of nanoMIL-100(Al). However, AMP encapsulation seems to destabilize nanoMIL-100 suspension. Both AMP-loaded nanoMIL-100(Al) and nanoMIL-100(Fe) particles were larger than 1000 nm in diameter, according to DLS measurement. They probably exist as agglomerates of small NPs, as their TEM images (Figure 3-47) display particles with a similar size and shape to the empty particles. The difference in the interaction between the ATP/AMP drug and the nanoMIL-100 carrier was further investigated using ssNMR spectroscopy on nanoMIL-100(Al). We took advantage of 27 Al and 31 P NMR that give signals exclusively belonging to the MOF and the drug, respectively. 30 wt% input drug loading was used to increase the NMR measurement sensitivity.

27

Al NMR spectroscopy demonstrates that the six-fold coordination mode of Al centers is kept upon drug loading (Figure 3-48A). A new broad shoulder around -10 ppm can be seen in the spectrum of ATP-loaded-nanoMIL-100(Al). These new 6-coordinate Al species in drug-loaded MIL-100(Al) are better observed by MQMAS spectra (Figure 3 Notice that in the 1D 27 Al NMR spectrum (Figure 3-48A), the new Al-6-coordinate resonance is less intense in AMP-loaded than in ATP-loaded nanoMIL-100(Al), indicating that the aluminum cluster interacts more strongly with ATP than with AMP. 27 Al MQMAS and 31 P- 27 Al NMR spectra of AMP-loaded MIL-100(Al), respectively.

The input drug loading is 30 wt% for both ATP and AMP Also, the chemical coordination between ATP and nanoMIL-100(Al) is proved by 31 P NMR spectroscopy. Figure 3-49A shows a significant shift and broadening of three peaks compared to those in pure ATP's spectrum due to chemical coordination between ATP and Al(III) through the terminal phosphate group (γ-P). The assignment of α-, β-, γ-P signals here was based on previously published 31 P-31 P DQ-SQ MAS NMR spectroscopy data 26 .

Figure 3-49. 31 P NMR spectra of (A) ATP-loaded and (B) AMP-loaded nanoMIL-100(Al). The 1D 31 P NMR spectrum of AMP-loaded MOF was compared to the 31 P dimension extract from its 2D 31 P- 27 Al NMR spectrum

Contrasting to the triphosphate derivative ATP, pristine AMP comprises a single phosphate group, thus having only one 31 P NMR resonance at 1.9 ppm. However, three distinctive resonances were observed in the 31 P MAS NMR spectrum of AMP-loaded-nanoMIL-100(Al) (Figure 3-49B). The peak at 0.8 ppm is similar to pure AMP; thus, it is assigned to the excess AMP, which is not incorporated in MIL-100 pores but is assumably trapped at its external surface. This peak's assignment is also supported by 31 P- 27 Al NMR measurement, which shows that it does not correlate with any 27 Al resonance of MIL-100. Meanwhile, the two other phosphorous species, whose chemical shifts are at -5 and -12 ppm, exhibit correlations with 27 Al, proving their encapsulation in MOF cages. In the extraction from the 2D spectrum, the intensity of the 31 P peak at -12 ppm is much higher than the peak at -5 ppm (Figure 3-49B). The 31 P peak at lower NMR chemical shift is assigned to AMP coordinated to Al of MIL-100(Al) framework because the deshielding effect should be more pronounced on the P atom chemically bonded to Al. The peak at -5 ppm is associated with AMP adsorbed physically in the pores but not chemically bonded to the framework.

Additionally, 13 C MAS NMR spectra (Figure 3-50A) of both ATP-and AMP -loaded MIL-100(Al) samples contain the peaks corresponding to trimesate linker in the range 170-180 ppm and 130-140 ppm, suggesting that the MIL-100 framework was preserved after drug loading. 1 H MAS NMR spectra of drug-loaded MIL-100(Al) also have signature peaks of trimesate ligand like the original one (Figure 3-50B). In short, drug encapsulation of ATP and AMP in nanoMIL-100(Al) and nanoMIL-100(Fe) was successfully realized, and their formulations were optimized. The integrity of the MIL-100 structure after drug loading was also checked. We also learned different mechanisms underlying this successful drug incorporation that highlight the affinity between MIL-100's metal sites and the drug's phosphate groups. The triphosphate drug ATP shows a stronger affinity towards nanoMIL-100 than the monophosphate drug AMP. SsNMR spectroscopy proved that all encapsulated ATP molecules were coordinated to the MOF's aluminum while encapsulated AMP molecules exhibit several types of interaction with nanoMIL-100(Al), including chemical coordination and physical adsorption. In addition, the iron-based MIL-100 appears to have a higher drug loading capacity than MIL-100(Al), probably due to the stronger interaction of metal cations with phosphate groups of the active molecule.

Degradation of drug-loaded nanoMIL-100

Once we managed to load the drug ATP and AMP inside nanoMIL-100 and get some knowledge about their drug-carrier interaction, we proceeded to study the degradation mechanism of the drug-loaded nanoMIL-100. Following up the study in the last section, we continued using 27 Al and 31 P ssNMR spectroscopy to study the ATP-and AMP-drug-loaded nanoMIL-100(Al). In line with the degradation study of the original nanoMIL-100(Al), these samples were degraded with a very low P/Al ratio (0.3) for 48 hours to observe the reaction's intermediate phase. To simplify the terminology in the later discussion, the phosphate ions from PBS, 61% of HPO4 2-and 39% H2PO4 -, are called PBS. The degradation mechanism of AMP-loaded nanoMIL-100 could be rationalized as follows.

Both containing monophosphate groups, AMP and PBS possibly have comparable binding strength with Al cations. It was reported that AMP and HPO4 Finally, the notable absence of the low-coordinate aluminophosphate at the very low P/Al ratio could be explained by the chelating capability of triphosphate ATP 78,79 . ATP coordinates with MIL-100(Al) through the terminal γ-phosphorus. During the breaking of the Al-O-C bond, the β-phosphorus of ATP can compete with phosphate PBS ions and create a second coordination bond with Al. The bidentate Al-ATP coordination could prevent the formation of low-coordinate Al and the total release of ATP. Nevertheless, when degrading the ATP-loaded MIL-100(Al) with a P/Al ratio higher than the one used in this ssNMR study, 4-coordinate aluminum phosphate was observed (data not shown).

Surface coating of CD-P on nanoMIL-100

Since surface coating is an important part of a DDS formulation, its role in carrier degradation must be investigated. The surface coating is usually designed to enhance the carrier's colloidal stability without preventing drug release.

CD-P surface coating did improve drastically the colloidal stability of nanoMIL-100(Al) and nanoMIL-100(Fe). Table 3-7 summarizes the DLS and zeta-potential measurement results.

The CD-P-coated nanoMIL-100 suspensions had acceptable NP size and polydispersity, whereas the original MIL-100(Al) agglomerates in DI water and the original nanoMIL-100(Fe) aggregates over time 27,75 . The CD-P-coated nanoMIL-100(Fe) even keeps their size stable after 2 years storage at 4 o C (deviation 14% and PdI = 0.11). Zeta-potential measurement gave values in the range ±10-20 mV, considered relatively stable NP suspension 80 . TEM images of CD-P-coated MIL-100(Al) also showed NP size comparable to the original MOF (Figure 3 -53). The integrity of the MIL-100(Al) framework after surface modification was checked by ssNMR spectroscopy. 13 C MAS NMR spectrum of CD-P-coated nanoMIL-100(Al) (Figure 3-54) exhibits the peaks corresponding to trimesate linker in the range 170-180 ppm and 130-140 ppm. Also, it shows small peaks of CD-P in the range 60-110 ppm, indicating the presence of CD-P in the MOF structure.

In addition, the 27 Al MAS NMR spectrum of the coated-nanoMIL-100(Al) was modified only slightly with respect to that of the original one (Figure 3-55A); the small shoulder around -17 ppm is hardly noticed. The corresponding 31 P MAS NMR spectrum also contains three broad resonances in the range from -20 to 10 ppm like those observed in pristine CD-P (Figure 3 -55B). This modest interaction between CD-P with MIL-100(Al) is because bulky CD-P cannot pass through the small pore windows of MIL-100(Al) and only interacts with the NPs at its surface. The same conclusion has been made in published studies about CD-P surface coating of nanoMIL-100(Al) 26 and nanoMIL-100(Fe) 27 . The earlier study on nanoMIL-100(Al) also used 31 P- 27 Al MAS NMR experiment to prove that CD-P surface coating was realized by chemical coordination between the surface Al sites and the phosphate functional groups of cyclodextrin, resulting from the replacement of the framework's labile water ligands 26 .

Degradation mechanisms of nanoparticle-based drug delivery systems: the case of polymer conjugate and metal-organic framework D. L. Mai VUONG -October 2022 207 In short, as expected from the literature, the CD-P surface coating could help to improve the colloidal stability of nanoMIL-100 particles and stays exclusively at their surface.

Degradation of surface-coated nanoMIL-100

Once the surface coating quality was evaluated, the degradation of CD-P-coated nanoMIL-100 was investigated in the same manner used for the degradation of drug-loaded one. The degradation reaction of CD-P-nanoMIL-100(Al) with a very low P/Al ratio of 0.3 was carried out for 48 hours.

SsNMR spectroscopy shows that at a very low P/Al ratio, the degradation of the CD-P coated nanoMIL-100(Al) behaves in the same way as the degradation of the non-coated MOF. In Figure 3-56A and B, the 27 Al MAS NMR spectrum of the degraded CD-P-coated MIL-100(Al) shows both the 6-coordinate and 4-coordinate Al-O-PBS around -2 ppm and 48 ppm, respectively. Observing similar degradation products, we speculate that, as CD-P probably covers only 30% of the nanoMIL-100(Al) surface 27 , degradation reaction still occurs at the nanoMOF's surface not attached to CD-P, following the reaction mechanism Studying degradation mechanism of metal-organic framework nanocarriers proposed in section 3.3.7 (see Figure 3-41). Additionally, the 31 P MAS NMR spectrum in Figure 3-57 displays a shoulder around -10 ppm belonging to the remaining CD-P coating, and the dominant signal of PBS coordinated with MIL-100(Al) around -15 ppm. The presence of CD-P coating even after five days of degradation is consistent with CD-P being anchored to MIL-100's surface through 3-4 monophosphate groups 27 . Consequently, CD-P coordinated with nanoMIL-100(Al) was only partially replaced by PBS, as previously observed with CD-P-coated nanoMIL-100(Fe). Less than 10% coating detachment was detected after 24 hours of incubation 27 with a higher P/M ratio than in the present study (5.4 and 0.3, respectively).

Combining the effect of drug encapsulation and surface coating, Figure 3-56 shows that degradation of the ATP-loaded CD-P-surface-coated nanoMIL-100(Al) produced both 6coordinate aluminophosphate complex and 4-coordinate inorganic aluminophosphate, but with a slightly lower proportion of 4-coordinate degraded complex (1.0%) than that in the surface-coated empty nanoMIL-100(Al) (1.6%). Also, the 27 Al peak at 2.5 ppm belonging to the original nanoMIL-100(Al) is more preserved in the degraded CD-P-MOF in the presence of the loaded ATP. Essentially, the degradation mechanism of the CD-P-coated MIL-100 in PBS is likely to follow the degradation mechanism of the original MIL-100, as CD-P covering only 30% of nanoMIL-100 surface does not block the contact and penetration of phosphate to MOF. As PBS substitutes coordinated water ligands and then substitutes coordinated trimesate ligands, 6-coordinate aluminum phosphate/iron phosphate is formed, and free trimesate is released. CD-P strongly anchored to the surface metal sites is only partially detached and released. The presence of CD-P coating in the degraded nanoMIL-100 can be of interest because it can prevent aggregation/precipitation of the degraded MIL-100 NPs. Meanwhile, carrier degradation, hence the drug release process, is seemingly neither hindered nor promoted considerably by the CD-P surface coating.

Summary

To conclude the effect of drug loading and surface modification on nanoMIL-100 degradation, the determining factor is the binding strength with the nanoMOF's metal center of the drug/coating respecting to that of the native trimesate ligands. Degradation of drug-loaded and surface-coated nanoMIL-100 generally obeys the mechanism proposed for original non-modified nanoMIL-100. When the loaded drug has a comparable metal complexing affinity with the phosphate reactant, the substitutions of the coordinated drug and coordinated trimesate probably occur simultaneously. In the case of active substances being strong complexing agents, the trimesate ligands could be replaced by phosphate before the coordinated drug. Regarding the surface coating, both the metal binding strength and the surface density should be considered. The coating having strong interaction with MOF's metal compared to phosphate's is likely to remain attached to the nanoMOF surface during degradation reaction with phosphate. However, a certain degree of surface covering is needed to protect nanoMOF from being exposed to and degraded by phosphate.

Degradation process of nanoMIL-100(Al) and (Fe) at nanoscale

So far, we have seen from many techniques how the molecular structures of nanoMIL-100(Al) and nanoMIL-100(Fe) evolve during their reaction with phosphate. As we are studying NP suspension, the NPs' point of view is also a crucial aspect. This section will examine how the degradation occurs at the nanoscale, particularly how individual MOF NPs experience the degradation reaction. Recall from the previously presented results that nanoMIL-100(Al) and nanoMIL-100(Fe), regardless of having likely similar atomic-scale degradation mechanisms, have different colloidal stability in their aqueous suspensions.

Global size and morphology properties of nanoMIL-100(Al) and (Fe) suspensions

Small-angle X-ray scattering (SAXS) was selected to follow the evolution of the nanoscopic features of nanoMIL-100(Al) and nanoMIL-100(Fe) suspension during their degradation reaction. In our experiments conducted in synchrotron (details in experimental section), the distance scale is roughly from 30 to 300 nm, which is calculated from 𝜋 𝑞 ⁄ , where q is the scattering vector in Å -1 . In the present work, SAXS makes it possible to look into the colloidal state in the suspension phase, such as particle size and aggregation. More importantly, this technique can offer the average view of hundreds to thousands or more particles. These pieces of information cannot be attained using microscopy techniques, which take images of samples deposited on grids and the solvent is evaporated, and only on a limited number of spots subjectively selected by the experimentalist. Still, we also performed electronic microscopy measurement, for the visualization by microscopy is necessary to complement the X-ray scattering data.

A typical SAXS pattern of nanoMIL-100 suspension, shown in Figure 3-58, can be divided into several different regimes. The Guinier regime gives information about particle size and size distribution, while the power law regime contains clues about particle morphology, like geometry and surface roughness. The region included in the power law regime that reflects the particle's surface roughness is called the Porod regime. To start with, the original nanoMIL-100 suspensions in absolute ethanol (EtOH) and deionized (DI) water were studied. Looking at their SAXS curves in Figure 3-59B, the first difference lies in the Guinier regime, where the scattering pattern of iron MOF is smooth, but the one of aluminum MOF is bent around 𝑞 of 0.005 Å -1 . This so-called cross-over point is a sign of particle aggregation or agglomeration 81 , which is consistent with DLS and TEM experiments on nanoMIL-100(Al) in water. The second difference is observed in the power law regime, precisely the Porod regime of nanoMIL-100(Fe), which has a slight upward tail starting from 𝑞 of 0.05 Å -1 compared to the curve of its aluminum analog, which could be a sign of additional diffuse scattering.

Evolution of size and surface roughness of degraded MIL-100 nanoparticles

The differences between the two nanoMIL-100 are significantly noticeable when monitoring their degradation process over time. Figure 3-60 shows how the SAXS patterns evolve when incubating nanoMIL-100(Al) and nanoMIL-100(Fe) in PBS at different input P/M ratios, where P stands for phosphate and M stands for metal, e.g., Fe or Al. To be specific, SAXS experiments used three ratios: very low (0.3), low (0.5), and moderate (5.5).

Among the three chosen reaction ratios, the most dramatic changes were witnessed when the degradation reaction occurred at a moderate input P/M ratio of 5.5. At this ratio, the SAXS patterns of the corresponding degraded nanoMIL-100(Fe) (Figure 3-60C) had a gradually growing scattering tail starting from 0.02 Å -1 in the Porod regime, which indicates a highly roughened surface [START_REF] Uhrich | Polymeric Systems for Controlled Drug Release[END_REF] . This surface roughening was once observed by Li et al., who reported TEM images of degraded nanoMIL-100(Fe) particles with rounded edges 29 . In our study, SEM images show that the degraded MIL-100(Fe), even at a high P/M ratio of 11, are still distinguishable particles (Figure 3-61). In contrast, a different phenomenon was observed in the degradation of nanoMIL-100(Al). At the P/M ratio of 5.5 (Figure 3-60F), the degraded nanoMIL-100(Al) at all time points have similar SAXS patterns but different from those of the original. The position of the cross-over point shifted significantly to a higher 𝑞 around 0.04-0.05 Å -1 , which may correspond to a size decrease of the MIL-100(Al) agglomerates in water. Note that at degradation time points of 24-48 hours, this cross-over point seems to locate at a slightly lower 𝑞. 11) and duration of 1.5 h (B and E) and 48 h (C and F). The final concentration of SEM's sample was 0.5 mg.mL -1 At very low and low input P/M ratios (0.3 and 0.5, respectively), the morphological transformation during degradation of the two MIL-100 is less pronounced. At these ratios, there is no change in the scattering curves of degraded nanoMIL-100(Fe) at all time points compared to the original one (Figure 3-60A, B). This observation from SAXS is consistent with the corresponding SEM images. For the P/M of 0.3, SEM images still contain MIL-100(Fe) degraded particles with similar shape and size to those of intact particles, although some aggregation was observed (Figure 3-62E-F). In comparison, the SAXS data of aluminum MOF at these conditions offer a view of the intermediate stage of the degradation process. Looking at Figure 3-60D andE, the SAXS patterns of degraded MIL-100(Al) at very low (0.3) and low (0.5) P/Al ratios share the same Guinier regime and the exact position of cross-over point with that of intact MIL-100(Al), indicating that the agglomerate size was kept. However, the power law regime was considerably modified. The scattering curve there was increasingly bent with degradation time, which could be attributed to the surface roughening as seen when degrading MIL-100(Fe) (Figure 3-60C) and/or the size decrease seen in the degradation of MIL-100(Al) at a moderate ratio of 5.5 (Figure 3-60F). Either way, both the increase in surface roughness and the decrease in particle size are plausibly related to the appearance of small particles at the nanoMIL-100(Al) surface detected by high-magnification SEM and TEM measurement at low and very low input P/M ratios (Figure 3 -63 and Figure 3-64). 3 and 11, respectively). The sample concentration was 0.5 mg.mL -1 In this case, the X-ray scattering data could be an indicator for the extent and the rate of MOF degradation reaction, i.e., the percentage and the rate of organic ligand release, respectively. For instance, the scattering tail in SAXS patterns of degraded nanoMIL-100(Fe)

was not found at a P/Fe ratio of 0.3 and 0.5 but 5.5. It seems that the phosphate-to-metal ratio should be high enough to create sufficient degraded MOF particles that have detectable scattering. Likewise, the gradual surface roughening of nanoMIL-100(Fe) could correlate with the progressive release of trimesate at the P/Fe ratio of 5.5 (see Figure 3-11A).

Whereas degradation of MIL-100(Al) at comparable P/Al ratios causes non-controlled trimesate release (Figure 3-11B); thus, the impact of surface roughening on the X-ray scattering curve is likely dominated by that of aggregation.

To summarize, the combination of X-ray scattering and electron microscopy demonstrates that nanoMIL-100(Al) and nanoMIL-100(Fe) particles were transformed in different fashions during the reaction with PBS. NanoMIL-100(Fe) globally maintains its size and morphology during the degradation reaction, no matter the input P/Fe ratio. When the input P/Fe ratio is sufficiently high, like 5.5, the iron nanoMOF particles undergo a gradual surface roughening, presumably related to their progressive trimesate ligand release. Regarding nanoMIL-100(Al), the way the particles are transformed is distinctive for each P/Al level. No change was detected at the very low P/Al of 0.3, while the low P/Al of 0.5 caused surface roughening and the moderate P/Al of 5.5 led to a combination of aggregation and surface roughening. The agglomerate size is maintained at very low and low P/Al ratios. The surface of degraded nanoMIL-100(Al) roughened more quickly and was observed at a much lower P/M ratio (P/Al of 0.5) compared to the iron MOF (P/Fe of 5.5). This surface roughening could be explained by the formation of small aluminum phosphate particles on the NP surface visualized under electron microscopy. At a moderate P/Al ratio (5.5), nanoMIL-100(Al) degradation mainly produces aggregates of these aluminum phosphate particles.

Effect of drug loading and surface coating on morphological evolution of degraded MIL-100 nanoparticles

As in the previous section, the degradation of modified nanoMIL-100 was studied by X-ray scattering and electron microscopy. Firstly, Figure 3-65 shows that the SAXS curves of the intact CD-P-coated, ATP-and AMP-drug-loaded nanoMIL-100 have the same shape and slope to their corresponding non-modified MOF, even though the intensity is different in the case of drug-loaded nanoMIL-100(Fe). The SAXS patterns of the modified nanoMIL-100(Al) still have the same cross-over point as the original nanoMIL-100(Al), and those of the modified nanoMIL-100(Fe) are smooth like the original nanoMIL-100(Fe). Consistently, the corresponding TEM images of modified nanoMIL-100 also present particles with similar size and shape to the original ones (Figure 3 -47 and Figure 3-53). A slightly different scenario was witnessed for the degradation of modified nanoMIL-100(Al).

The SAXS patterns of degraded drug-loaded nanoMIL-100(Al) follow the general trend in those of degraded original ones (Figure 3-67A, B, C). However, the changes in the power law regime, i.e., the location of the cross-over point and the curved shape, were less apparent in the presence of encapsulated drugs. It seems that somehow the presence of ATP and AMP slowed down the formation and aggregation of small particles during the nanoMIL-100(Al) degradation. This easing aggregation was even more significant when the nanoMIL-100(Al) was coated with CD-P. The location of the cross-over point in the X-ray scattering curves, reflecting the size of the initial nanoMIL-100(Al) agglomerate, was roughly retained upon degradation of the CD-P-coated aluminum MOF, even at the moderate P/Al ratio of 5.5 (Figure 3-67D, E, F), in opposed to the significant shift of cross-over points observed with the degraded non-coated MOF. Overall, it seems that the encapsulated ATP and AMP and the CD-P surface coating only influence the degradation of nanoMIL-100(Al) but do not in the case of the iron analog.

Interaction between nanoMIL-100 and water in the nanoparticle suspension

Different behaviors at the nanoscale of degraded nanoMIL-100(Al) and nanoMIL-100(Fe) probed by X-ray scattering and electron microscopy studies above could be related to some specific interaction between water and MIL-100 NPs in their aqueous suspensions. As MIL-100 is microporous, we suspect that water molecules located inside the micropores, at the NP surface, and outside the NPs have different behaviors and interactions with the MIL-100 NPs (Figure 3 -68). During the nanoMOF degradation reaction, water molecules not only act as the solvent for the transport of reactants and products but also are involved in the initial reaction step themselves (discussed in section 3.3.7). Then, the behaviors of the surface and pore water molecules strongly interacting with the particles could be more influenced by the degradation process than the bulk water. Therefore, probing the behavior of water might be a way to follow the degradation process of nanoMIL-100. In the following section, we explored these plausible interactions and effects by NMR spectroscopy and relaxometry.

The NMR relaxometry was chosen to examine the exterior and interior surface of nanoMOF upon its degradation, as this technique is sensitive to changes in dynamic processes and molecular motions 83 . Firstly, the existence of interaction between water H2O and nanoMIL100 particles in their aqueous suspension is demonstrated by lqNMR spectroscopy. In this experiment, the NMR tubes contained the sample, the nanoMIL-100 aqueous suspension of a given concentration, and a coaxial insert containing a methylphosphonic acid solution in D2O. The coaxial insert serves for locking the magnetic field and reference purposes. Consequently, the NMR signals of the neighboring nuclei are heavily modified in terms of peak position and width. In contrast, the HDO inside the coaxial insert was not in direct contact with Fe(III). As a result, neither its 1 H linewidth nor its chemical shift was altered.

Meanwhile, the 1 H NMR spectrum of nanoMIL-100(Al) suspension consists of one main signal at 4.68 ppm (Figure 3 -69). This signal was assigned to the water in nanoMIL-100(Al) sample. Here, the peak of residual HDO could not be clearly distinguished because it was dominated by the intense peak of H2O. At this low concentration of 1 mg.mL -1 , the interaction between nanoMIL-100(Al) and water, hence the influence on the 1 H NMR signal of water, was less apparent than in the case of iron analog. To better assess the interaction between water and nanoMOF, the nanoMIL-100 concentration was varied from diluted (1 mg.mL -1 ) to concentrated (150 mg.mL -1 ). The resulting 1 H NMR spectra series are given in Figure 3-70. Starting with nanoMIL-100(Al) (Figure 3-70A), with higher MOF concentration, there was no apparent change in 1 H chemical shift but a gradual peak broadening of water, probably because more water molecules were confined in the nanoMIL-100(Al) pore and/or the viscosity increased, limiting the mobility of water. Rigid molecules have fast transverse relaxation rates 1/T2, thus having broad NMR peaks. Conversely, a more dramatic evolution of the water signal was witnessed with nanoMIL-100(Fe) samples (Figure 3-70B), as expected from the interference of paramagnetic Fe(III), which escalates with the iron MOF's concentration. The series of 1 H NMR spectra of nanoMIL-100(Fe) aqueous suspension with different concentrations always contains two peaks. The peak with an almost constant chemical shift at 4.68-4.69 ppm and line width of 2-3 Hz belongs to the residual HDO inside the coaxial insert, and the broad peak with changing position and increasing line width represents the water in nanoMIL-100(Fe) suspension. The water HDO inside the coaxial insert was not in close contact with Fe(III), thus not being affected by the elevating paramagnetic effect and maintaining its position and line width when changing the nanoMIL-100(Fe) concentration. On the opposite, the peak of water in nanoMIL-100(Fe) suspension was significantly modified upon varying the suspension's MOF concentration, evidencing that bulk water is exchanging with the pore and interface water. As shown in Figure 3-70A, the higher the concentration of nanoMIL-100(Fe), i.e., the higher the amount of paramagnetic Fe(III), the broader linewidth, and the higher chemical shift of water. Also, the paramagnetic influence is orchestrated in the 1 H relaxation rate (Figure 3 -71). That of water in the iron MOF sample is 200 times shorter than in the aluminum MOF. In short, the lqNMR data at room temperature highlight the interaction between water and nanoMIL-100 in its suspension. According to our hypothesis of bulk and pore water, they should have different 1 H NMR signals with distinctive features. However, at room temperature, water molecules undergo fast exchange between the bulk, surface, and pore environment. As a result, we could observe one representative NMR signal in both nanoMIL-100(Al) and nanoMIL-100(Fe) samples.

To prove the presence of different water environments in nanoMOF suspensions, NMR measurements at low temperature were performed. If existing, the pore and surface water are frozen at a much lower temperature than the bulk, whose freezing temperature is comparable to pure water. The presence of bulk and pore/surface water in nanoMIL-100(Al) was further evidenced by the fact that the proportion of pore water, calculated from its 1 H NMR peak area with respect to that of all water at room temperature, has a linear correlation with the concentration of nanoMIL-100(Al) (Figure 3 -73). The higher the NPs concentration, the more empty cages of nanoMOF that water can occupy.

Chapter 3 Studying degradation mechanism of metal-organic framework nanocarriers To sum up, lqNMR spectroscopy at low temperature proved that water molecules in nanoMIL-100 suspension experience different environments and display different behaviors: the bulk solvent and inside the MOF pore/close to the MOF surface. The bulk water has little interaction with the particles, whereas the pore/surface water strongly interacts with the metal cations in MIL-100. NMR spectroscopy at room temperature mainly detects the effect of space confinement in MIL-100(Al) and the paramagnetic effect of Fe(III) cations in MIL-100(Fe). 3.5.5 Behavior of water in degraded nanoMIL-100(Al) and (Fe)

Using the result from high-field NMR spectroscopy, low-field relaxometry measurements on nanoMIL-100 aqueous suspension were performed at cold and room temperature to examine the evolution of the water behaviors in the interior and exterior surface of nanoMIL-100 upon degradation.

We performed the NMR relaxometry measurement of nanoMIL-100(Al) at -15 o C, at which the bulk water in nanoMOF suspension was frozen, and there was no water motion among MOF particles. Therefore, the resulting 1 H NMR dispersion (NMRD) profile represents the relaxation of only pore/surface water in nanoMIL-100(Al) suspension, which is related to the amount of interior surface accessible to water inside MOF and not affected by the MOF quantity. Figure 3-77 presents the 1 H NMRD profile of degraded nanoMIL-100(Al) compared to the original one. Their NMRD profiles have a similar shape at the high-frequency range and start to look different from 0.2 MHz. Due to the very low amount of pore water in the degraded MOF, the obtained data at the low-frequency range was too noisy and not reliable enough to be considered, thus not being reported. Still, we could see an increase in the relaxation rate after degradation, probably because the pore/surface water becomes more rigid in MIL-100(Al) with decreasing particle size, found previously by SAXS and SEM. The effect of the encapsulated drug on water-nanoMOF interaction was also studied. The 1 H NMRD profiles of ATP-and AMP-loaded MIL-100(Al) nearly superimpose that of empty MIL-100(Al) at the high-frequency range but differ at the low-frequency range (Figure 3-78A).

The 1 H relaxation rate at low-frequency range decreases following the order: ATP-loaded > AMP-loaded > empty MIL-100(Al) samples. Also, their corresponding NMRD curves have different starting frequencies of the plateau in relaxation rate (data points were measured from high to low frequency). At cold temperatures, when only motions inside nanoMOF are possible, this frequency could associate with the time it takes water molecules to move from the pore to the surface. The lower the frequency, the longer the diffusion from pore to surface. Both the increase in water relaxation and diffusion rate suggests that the experience of water inside MOF cages was modified. Plausible explanations include alteration of pore structure, pore volume, and/or affinity with the species at/in the pore. ATP and AMP have more exchangeable protons than water; therefore, their encapsulation in MIL-100(Al) pores raises the chances of 1 H-1 H interaction in the pores, enhancing the surface diffusion and relaxation rate of pore water. In addition, Figure 3-78B shows that the NMRD profiles of degraded AMP-loaded MIL-100(Al) seem to follow the same trend as the degraded empty one, even though the former sample had noisy data and its full plot could not be acquired due to low amount of pore water there.

In brief, nanoMIL-100(Al) degradation seems to reduce its porosity or access to its pore network. The first observation is that the 1 H NMRD profile of degraded nanoMIL-100(Fe) is significantly altered compared to the original one and has a signature shape of an aqueous solution composed of paramagnetic ions, following the Solomon-Bloembergen-Morgan equation 84 . This development suggests that upon nanoMIL-100(Fe) degradation, the way Fe(III) is exposed at the nanoMOF's outer surface is changed. One of many reasons could be the existence of iron phosphate layer at the nanoMIL-100(Fe) surface formed after its degradation reaction with phosphate (see section 3.3.7). Also, one can notice that the degraded nanoMIL-100(Fe) has a lower relaxation rate than the intact NPs. Let us consider that the number of MOF particles did not change. This reduction in relaxation rate indicates a less intense paramagnetic effect of Fe(III), which could relate to a decrease in accessible iron sites to bulk water or different paramagnetic properties between iron trimesate and iron phosphate.

Furthermore, as observed with the aluminum MOF, the encapsulated drug does not have a substantial impact on the NMRD profiles of the degraded nanoMIL-100(Fe). The NMRD profiles of the empty and loaded nanoMIL-100(Fe) have similar shapes (Figure 3 -79). Note that the profile of ATP-loaded multiplied 1.4 times superimposes that of the empty. Given that both have equal MOF concentrations, this indicates a decrease of accessible iron sites to bulk water after drug encapsulation. As ATP occupies iron sites and space inside MOF cages, the exchange between bulk and pore water is less favored, thus decreasing the probability of interaction between bulk water and iron sites. After degradation, the NMRD profile of degraded ATP-loaded nanoMIL-100(Fe) matches that of degraded original MOF, except for a slight decline in intensity (Figure 3 -79). It seems that the presence of a loaded drug does not modify the MOF outer surface so much, as the drug is supposed to locate exclusively inside MOF cages and not at its surface. In short, nanoMIL-100(Fe) degradation results in a pronounced modification in the NMRD profile, which is related to the accessibility of water to surface iron sites. The paramagnetic interference from Fe(III) species makes the evolution of NMRD profiles of nanoMIL-100(Fe) more noticeable than its aluminum analog, as previously observed with high-field NMR spectroscopy.

To summarize, these results of high-field NMR spectroscopy and NMR relaxometry evidence the presence of internal and external environments for water in nanoMOF suspension and their changes upon nanoMOF degradation. Furthermore, NMR relaxometry experiments of nanoMIL-100(Al) at -15 o C and nanoMIL-100(Fe) at room temperature complementarily show that their degradation reaction with phosphate modifies the diffusion and molecular motions at the exterior and interior surfaces of MIL-100 NPs. Upon degradation, water molecules in the bulk and outer NP surface become less accessible to paramagnetic iron(III), thus significantly changing the shape of the degraded NP's NMRD profile. For diamagnetic nanoMIL-100(Al), water in their pores or the internal surface layer seems less mobile. Additionally, drug encapsulation probably interferes with the water mobility within and outside MOF particles but does not significantly affect the water behavior in degraded MOF.

Discussion on the erosion process of nanoMIL-100

Based on the ensemble of data presented in this study, some proposals could be made about the erosion process of nanoMIL-100 material. The two analogs nanoMIL-100(Al) and nanoMIL-100(Fe) are likely to experience the same degradation reaction steps with phosphate in PBS (ssNMR, XANES) but react with dissimilar rates and reactivity; thus, they are eroded in different ways.

Firstly, the two nanoMIL-100 MOFs have comparable porosity, cage size, and pore opening window. Also, phosphate is smaller than the MIL-100 pore window. Accordingly, phosphate ions can penetrate these porous nanoMOFs through the pore window and react with the metallic centers. This argument is true when there is no barrier for phosphate diffusion, but any pore-closing event could hinder the action of phosphate ions.

NanoMIL-100(Fe) seems to preferentially undergo surface erosion, judging from its high affinity towards phosphate and progressive trimesate release profile (lqNMR). When the P/Fe molar ratio is not high enough to cause total (100%) trimesate release, the degradation and erosion rate decrease with time as the accumulation of iron phosphate passivates the nanoMIL-100(Fe) surface and reduces its porosity (N2 adsorption), thus limiting phosphate from attacking nanoMOF. The surface erosion mechanism also fits the linear increase in surface P/Fe ratio (XPS) and non-linear increase in bulk P/Fe ratio (EDS) as a function of % trimesate release. Nevertheless, a hallmark of classic surface erosion, which is the decrease in DDS's size over time, is not observed in the case of nanoMIL-100(Fe). Note that the surface roughness of nanoMIL-100(Fe) did not modify significantly at very low and low degradation ratios (0.3 and 0.5, respectively) but a moderate value of 5.5 (SAXS). We speculated that degradation at low-range input P/Fe ratios produces an iron phosphate shell surrounding the intact MIL-100(Fe) core, as previously found with microMIL-100(Fe) 29 . Such a core-shell structure is consistent with the SAXS data showing the degraded nanoMIL-100(Fe) having the same NP size as the original one.

Furthermore, surface erosion appears to occur in nanoMIL-100(Al) particles at low input P/Al ratios (0.3 and 0.5), based on its progressive trimesate release profile and the presence of small species at nanoMIL-100(Al) surface (TEM). A decrease in aluminum MOF porosity also took place, but to a more significant extent than iron MOF. However, at increased input P/Al ratio that can induce total degradation of nanoMIL-100(Al), the phosphate penetration is fast enough to react with nanoMIL-100(Al) before the passivation occurs. One can imagine that nanoMIL-100(Al) quickly disintegrated and transformed into smaller aluminum phosphate particles (SEM, SAXS). That kind of bulk erosion could lead to non-controlled trimesate ligand release and concomitant increase in bulk and surface P/Al ratio. Note that the initial nanoMIL-100(Al) in water exists as agglomerates, unlike separate particles of nanoMIL-100(Fe).

All in all, the nanoMOF erosion process is governed by a combination of successive processes: phosphate penetration, phosphate substitution, surface passivation, and pore closure.

Conclusions and perspectives

To conclude, this study demonstrates the degradation mechanism of nanoMIL-100 drug carriers in phosphate buffer, a simulated physiological medium. The reaction was examined for two iron-and aluminum-based MIL-100 analogs whose degradation behaviors have both similarities and differences. The degradation of drug-loaded and surface-coated nanoMIL-100 was also studied. Moreover, multiple techniques have been involved in investigating the molecular and nanoscopic aspects of the degradation, providing a description of the whole process: degradation reaction and material erosion (recap scheme below).

At the molecular scale, several spectroscopy techniques, including lqNMR, ssNMR, XANES, XPS, IR, and EDS, helped to identify the nature, quantity, and composition of the degradation products both in liquid and solid phases. The chemical structure, crystallinity, and porosity of degraded nanoMIL-100 were determined.

Ligand substitution by phosphate ions is the main driving force for nanoMOF degradation in phosphate buffer. A stepwise reaction mechanism is proposed for both nanoMIL-100(Al) and nanoMIL-100(Fe). First, upon substitution of labile water ligands by phosphate, M-O-P (M = Fe, Al) coordination bonds are formed between metal sites at the nanoMOF surface and phosphate from the medium, creating 6-coordinate aluminum phosphate or iron phosphate in the case of MIL-100(Al) or MIL-100(Fe), respectively. Second, the substitution of trimesate by phosphate leads to the release of free trimesate and creates more 6coordinate metal phosphate complexes. During the dissociation of the metal-trimesate M-O-C bond, another degradation product, 4-coordinate aluminum phosphate, might form. Afterward, the degradation reaction could progress following two possibilities depending on the input P/M ratio. If the input P/M ratio is not high enough to induce the full release of trimesate ligands, the build-up of the 6-and 4-coordinate metal phosphate complexes at the nanoMOF surface partially blocks access of phosphate to nanoMOF and therefore slows down the degradation reaction. When the input P/M ratio is sufficient so that the coordination between phosphate and the Al/Fe metal centers is faster than the surface passivation, nanoMIL-100 could be totally transformed into metal phosphate or metal hydroxyphosphate.

Moreover, our study demonstrates that chemical modification such as drug encapsulation and surface coating can influence the nanoMOF degradation, hence its drug release process. The main affecting factor is the metal-ligand bond strength of new coordinating agents relative to the existing ones. The triphosphate-bearing drug ATP, with a stronger interaction with nanoMIL-100 than its monophosphate derivative AMP, could compete with the degrading agent, phosphate ions, and alter the MIL-100 carrier degradation and drug release. Surface coating with CD-P ameliorates the nanoMOF's colloidal stability in water but has no significant effect on its degradation, probably related to its surface covering density.

At the nanoscale, the behaviors of MOF nanoparticles and the surrounding water in their suspension change during the degradation process, which could be rationalized based on the degradation mechanism at the molecular scale. X-ray scattering and electronic microscopy illustrated the size and morphological evolution of MIL-100 NPs during degradation. On the one hand, MIL-100(Fe) NPs seem to react as individual particles and maintain their global size and morphology. Starting from a certain P/Fe ratio that causes a high degree of degradation, the degraded nanoMIL-100(Fe) undergoes gradual surface roughening due to the formation of iron phosphate. On the other hand, the growth of small aluminum phosphate particles on MIL-100(Al) initial agglomerate results in a more abrupt surface roughening. The aggregation of these aluminum phosphate particles was detected when nanoMIL-100(Al) was close to being fully degraded.

Additionally, preliminary data of NMR relaxometry illustrate different types of interaction between water and nanoMIL-100(Al)/nanoMIL-100(Fe) in their aqueous suspension. NMR spectroscopy at cold temperature made it possible to distinguish the water confined in the MOF cages or near the internal surface versus water in bulk outside MOF. Also, the NMR relaxation of water in nanoMIL-100(Fe) is heavily affected by the paramagnetism of iron(III).

Later, the NMR relaxometry study of degraded and drug-loaded MOF systems suggests reduced water mobility inside MOF pores and altered accessibility to surface metal sites.

Here, the flexible use of NMR spectroscopy is helpful in investigating nanoMOF degradation. Quantitative lqNMR spectroscopy determines the reaction kinetics and the nature of the degradation products in the liquid phase. SsNMR spectroscopy provides direct evidence about the chemical structure and the metal coordination sphere of the degraded nanoMIL-100. NMR relaxation and relaxometry give information on the dynamic and mobility of water to probe modification of nanoMOF's pore and surface environments Nevertheless, further experiments will be highly appreciated to understand more about the degradation and drug release mechanism of nanoMIL-100 material. For example, kinetics experiments for drug loading and drug release processes of ATP and AMP are necessary to analyze the influence of the drug-carrier interaction on the drug delivery performance of nanoMIL-100. In addition, modeling the NMR relaxometry data and calculating the dynamic parameters of pore/surface and bulk water would help to verify our hypothesis about the behaviors of water in degraded nanoMIL-100 suspension.

A possible outlook of this work is transferring the multiscale approach presented here to study the degradation process of other MOF families and porous systems. It is also beneficial to investigate MOF degradation under more complex conditions, e.g., a combination of coordinating anions including phosphate, sulfate, and carbonate, as well as conditions closer to the actual physiological environment in the human body. Expanding our knowledge about the nanoMIL-100 chemistry can help to rationalize their drug delivery performance.

Chapter 3 Studying degradation mechanism of metal-organic framework nanocarriers 3.8 Experimental methods

Chemicals

All reagents and solvents were used without any purification. The chemicals and their sources are listed as follows.

For nanoMOF synthesis: aluminum nitrate nonahydrate Al(NO3)3.9H2O, 98% (Alfa Aesar, Thermo Fisher), trimethyl 1,3,5-benzenetricarboxylate, i.e., trimethyl trimesate 98% (Alfa Aesar, Thermo Fisher), benzene 1,3,5-tricarboxylic acid, i.e., trimesic acid (Aldrich Chemistry), nitric acid 65% (Fluka Analytical of Sigma Aldrich) Solvent: methanol HPLC grade 99.9% (Fisher Chemical, Sigma Aldrich), ethanol absolute (Analar NORMAPUR of VWR Chemicals). Deionized water was taken from the PURELAB Ultra ELGA purification system.

For drug loading: adenosine 5'-triphosphate disodium salt hydrate and adenosine monophosphate sodium salt with purity ≥ 99% (Sigma Aldrich). For surface modification: phosphated β-cyclodextrin (degree of substitution DS ≈ 2-6) sodium salt from CYCLOLAB R&D For nanoMOF degradation: Dulbecco's Phosphate Buffered Saline (PBS) 1X, i.e., 10 mM, pH 7.0-7.3 (Gibco by life of Thermo Fisher), comprised of 8.1 mM Na2HPO4, 1.5 mM KH2PO4, 138 mM NaCl, and 2.7 mM KCl.

For HPLC experiments: monoammonium phosphate NH4H2PO4 and tetrabutylammonium perchlorate (TBAP) 98% are from ACROS Organics, acetone nitrile 99.8% from Fluka Analytical. For liquid NMR experiment: 3-(Trimethylsilyl)propionic-2,2,3,3-d4 acid sodium salt 98 atom% D (Sigma Aldrich), deuterated water D2O (Innova-Chem). For ICP: nitric acid 69% of analytical grade, certified reference material aluminum and iron standard are all from Sigma Aldrich.

Synthesis of nanoMIL-100

NanoMIL-100(Al) and nanoMIl-100(Fe) were synthesized by microwave synthesis technique, adapting and following published procedure 8,34 .

NanoMIL-100(Al) synthesis: In a Teflon reactor, 1.425 g of aluminum(III) nitrate nonahydrate and 1.215 g trimethyl benzene-1,3,5-tricarboxylate were mixed in 20 mL DI water under vigorous stirring for 5 minutes at room temperature. Next, 4 mL of nitric acid 4 M was added, and the reaction mixture was stirred for 5 minutes at room temperature. Then, in the microwave system, two reactors were heated from 30 to 210 o C for 10 minutes, then held at 210 O C for 30 minutes. The microwave system (CEM Microwave Accelerated Reaction System MARS 5) optimized the applied power automatically. The maximum power output was 1600 W, 50%. The reactors were taken out at 80 o C and cooled down in an ice bath for 15 minutes. The resulting yellow wet solid was recovered by centrifugation at 10500 RCF (relative centrifugal force) for 30 minutes and then was stirred in 50 mL methanol overnight to remove excess reactant. The activated product was recovered by centrifugation at 12000 RCF for 15 minutes and then dispersed in absolute ethanol (EtOH abs). In the size separation step, the largest particles were removed from this suspension by a short centrifuge of 6000 RCF for 1 minute; then, the obtained liquid fraction was centrifuged at 6000 RCF for 1 minute. The final nanoMOF suspension was stored at room temperature in 25-30 mL EtOH abs.

NanoMIL-100(Fe) synthesis: In a Teflon reactor, 1.622 g of iron(III) chloride hexahydrate and 0.845 g benzene 1,3,5-tricarboxylic acid were added, followed by adding 20 mL of deionized water and heated for 6 min at 130 o C under stirring. The applied power was optimized automatically by the microwave system. The maximum power output was 800 W, 50%. The reactors were taken out at 80 o C and cooled down in an ice bath for 15 minutes. The resulting brown wet solid was recovered by centrifugation at 10000 RCF for 10 minutes and then washed with EtOH abs six times to remove excess reagents. The product was then recovered by centrifugation at 11000 RCF for 12 minutes and then dispersed in EtOH abs. The largest particles were separated from this suspension by a short centrifuge of 6000 RCF for 1 minute. The final supernatant was stored at room temperature at room temperature in 25-30 mL EtOH abs.

The concentration of the final nanoMOF suspension was determined by drying a 500 µL sample at 50 o C for 3-5 hours. The fully hydrated nanoMOF was prepared by drying the sample in the oven at 100 o C for 2 days and then left exposed to an ambient atmosphere for 2-3 days. The as-synthesized MOF was routinely characterized by DLS, XRD, and gas adsorption to check the NP size, porosity, and crystallinity.

Drug encapsulation

Following a reported soaking procedure 14 , 5 mg.mL -1 MOF suspension was mixed with 1.5 mg.mL -1 drug aqueous solution, corresponding to 30 wt% input drug loading. The concentration of drug solution is 1.25 and 1.0 mg.mL -1 when the input drug loading is 25 and 20 wt%, respectively. The volume of the drug solution was determined from the needed weight of MOF. The drugs are adenosine triphosphate (ATP) or adenosine monophosphate (AMP). Their solutions were prepared from ATP disodium hydrate salt or AMP sodium hydrate salt.

First, in a round-bottom 50 mL centrifuge tube, the wet pellet of nanoMIL-100(Al) or nanoMIL-100(Fe) was extracted from the stock suspension in EtOH abs by centrifuge 10000 RCF for 15 minutes. To remove the residual ethanol, this pellet was washed two times in deionized water (DI water). The pellet was gently mixed in 100 µL DI water first, then 1 mL DI water was added, and the suspension was remixed. The volume of added water could be doubled when a higher amount of MOF was prepared. Water was removed by centrifuging 10 000 RCF for 10-15 minutes. After the washing step, a small volume of 1.5 mg.mL -1 ATP/AMP solution was added to the wet MOF pellet such that the wet pellet was immersed in the liquid, and the mixture was dispersed by gentle shaking and vortex. When this suspension was well dispersed, the remaining drug solution was added, and the mixture was subjected to continuous shaking overnight at room temperature. The drug-loaded nanoMOF was recovered by centrifugation at 10000 RCF for 10 minutes, which was suspended in DI water to get a 5 mg.mL -1 aqueous suspension or dried under atmosphere at room temperature 20 o C or under vacuum at 25 o C.

The actual drug payload wt% was calculated as:

𝑚 𝑑𝑟𝑢𝑔 𝑖𝑛𝑝𝑢𝑡 -𝑚 𝑑𝑟𝑢𝑔 𝑟𝑒𝑚𝑎𝑖𝑛𝑖𝑛𝑔 𝑚 𝑀𝑂𝐹 × 100, where 𝑚 𝑑𝑟𝑢𝑔 𝑖𝑛𝑝𝑢𝑡 is the weight of the drug added to nanoMOF. The amount of unloaded drug remaining in the supernatant after drug loading, 𝑚 𝑑𝑟𝑢𝑔 𝑟𝑒𝑚𝑎𝑖𝑛𝑖𝑛𝑔 , was determined by HPLC. The encapsulation efficiency (EE%) was calculated by the ratio between the actual loaded drug quantity and the input quantity.

Surface modification

Adapting the method disclosed by Agostoni et al 27 , surface coating of cyclodextrinphosphate (CD-P) on nanoMIL-100 surface was carried out by mixing MOF suspension 10 mg.mL -1 with CD-P aqueous solution of 3 mg.mL -1 . CD-P solution was prepared from β-CD-P sodium salt (average 3-4 phosphate groups per cyclodextrin unit). The volume of the coating solution was determined from the needed weight of MOF.

The wet pellet of nanoMOF was extracted and washed two times in DI water like in the drug loading procedure. Then, CD-P solution was added to this obtained pellet. The resulting suspension was mixed overnight at room temperature. The surface-coated-nanoMOF was recovered by centrifuging at 10000 RCF for 10 minutes, then washed three times in DI water to remove excess CD-P. It was stored similarly to drug-loaded-MOF.

The actual wt% of CD-P coating was calculated as:

𝑚 𝑐𝑜𝑎𝑡𝑖𝑛𝑔 𝑖𝑛𝑝𝑢𝑡 -𝑚 𝑐𝑜𝑎𝑡𝑖𝑛𝑔 𝑟𝑒𝑚𝑎𝑖𝑛𝑖𝑛𝑔 𝑚 𝑀𝑂𝐹 × 100,
where 𝑚 𝑐𝑜𝑎𝑡𝑖𝑛𝑔 𝑖𝑛𝑝𝑢𝑡 is the weight of CD-P added to nanoMOF. The amount of CD-P remaining in the supernatant after surface coating and washing (𝑚 𝑐𝑜𝑎𝑡𝑖𝑛𝑔 𝑟𝑒𝑚𝑎𝑖𝑛𝑖𝑛𝑔 ) was determined by quantitative 31 P NMR spectroscopy.

Degradation of nanoMIL-100

NanoMOF was mixed with phosphate buffer saline (PBS) pH 7.4 at 37 o C for 2-5 days under several conditions defined by the nanoMOF and the PBS concentration (Table 345678). The mole of metal was calculated from the weight percentage of metal in the original nanoMOF, which was determined by ICP and TGA (see 3.8.6.). The mole of phosphorous was determined from the sum of phosphate salts in PBS (Na2HPO4 and KH2PO4). The incubation of nanoMIL-100(Al) or nanoMIL-100(Fe) in phosphate buffer was carried out as follows. First, a suspension containing a determined mg of nanoMOF in EtOH abs was centrifuged at 11000 RCF for 15 minutes. The taken volume of EtOH abs suspension was weighed when preparing samples for quantitative studies. After removing EtOH abs, the obtained pellet was washed 2 times with 1-3 mL DI water, depending on its milligrams, and then centrifuged at 10000 RCF for 10 minutes. The wet pellet was redispersed in 100-200 µL DI water first before adding the remaining DI water, as described in the drug loading procedure. After the removal of water from the second water washing, PBS was added to the wet pellet. The concentration of PBS was dependent on the chosen reaction conditions, and its volume was based on the MOF concentration and MOF weight used. The degradation reaction was timed, starting from the addition of PBS. Suspension of nanoMOF in PBS was shaken vigorously for a few seconds, then placed immediately in a rotator inside an oven set at 37 o C. The nanoMOF was incubated in PBS for a given time, namely, 0.5, 1.5, 3, 6, 18, and 48 hours. When the time was over, the sample was removed from the oven and then immediately centrifuged at 13000 RCF for 15 min. For the time point "0 h", the sample was mixed with PBS for 5-10 seconds and then centrifuged with the same force. The recovered supernatant, called the degraded supernatant, was stored in the fridge at 4 o C until further characterization. The amount of trimesate in the degraded supernatant was determined by HPLC or quantitative 1 H NMR spectroscopy. The recovered pellet, called the degraded pellet, was washed once with at least 1 mL DI water, and then centrifuged at 10000 RCF for 10 minutes. In the case of XPS characterization, to wash the excess PBS after the incubation, the degraded pellet prepared from a high P/M ratio was washed two times with DI water of the same volume as the added PBS. This degraded pellet was finally dried at room temperature 25 o C or suspended in a specific volume of DI water to get a desired concentration, depending on the later characterization technique. Apart from XPS, TEM, SEM, EDS, SAXS, and NMR relaxometry, which measure degraded nanoMOF suspension, all other techniques measure dried powder of degraded MOF. Note that the "MOF concentration" of the degraded MOF suspension is based on the initial weight of MOF. The remaining MOF content after the degradation reaction was not known.

For the drug-loaded and surface-coated nanoMIL-100, PBS was added to the wet MOF pellet recovered right after the surface coating or drug encapsulation process or to dried powder.

Characterization techniques for nanoMIL-100

Particle size distribution and zeta-potential of nanoMOF suspension were analyzed by dynamic light scattering (DLS) and electrophoretic light scattering in a Zetasizer Nano ZS90 instrument from Malvern Panalytical. For DLS, 0.5 mg.mL -1 nanoMOF suspension in EtOH abs or DI water was prepared. 5 mg of MOF was centrifuged from stock suspension in EtOH abs and then washed twice in DI water as described above. After the second washing, the obtained wet pellet was redispersed in 1 mL of DI water by vortex. The obtained suspension is then diluted with DI water to achieve the final nanoMOF concentration of 0.5 mg.mL -1 . For ZP, 5 mg.mL -1 nanoMOF suspension was diluted with a KCl solution of 10 mmol.mL -1 and DI water. Their corresponding volumes were adjusted so that the final suspension consists of 0.5 mg.mL -1 nanoMOF and KCl 1 mmol.mL -1 .

From the wt% extra trimesate and the MW of MIL-100, the number of extra trimesate molecules was calculated. Finally, the mass of non-coordinated water was estimated by subtracting that of the SBU M3O(H2O)2OH(C6H3(CO2)3)2 and extra trimesic acid from the MW of MIL-100. Dividing the obtained mass by the MW of H2O, one get the number of noncoordinated water. For ICP, 3-4 mg of MOF dried powder was digested in 1 mL HNO3 69% inside an oven set at 100 o C overnight. The following day, immediately after taking the sample out of the oven, the sample was diluted with DI water and 2% HNO3 solution of calculated volume so that the final HNO3 concentration was 2%. ICP measurement of this sample was carried out on an Agilent 720 Series ICP-OES apparatus.

High-performance liquid chromatography (HPLC) experiments were done to determine the amount of trimesate formed during nanoMIL-100 degradation and the amount of ATP/AMP. For trimesic acid quantification, as previously reported 29 , the mobile phase consisted of 90% v/v ammonium dihydrophosphate NH4H2PO4 (5.75 g.L -1 in DI water) and 10% v/v tetrabutylammonium perchlorate TBAP (5 mmol.L -1 in acetonitrile). The detection wavelength was 220 nm. The flow rate was 1 mL.min -1 , and the run time was 10 minutes. Standard solutions of trimesic acid in DI water were prepared to determine the calibration curve and the degraded supernatant, were diluted in DI water. All measurements were done in duplicate. For ATP and AMP quantification, following a reported procedure 85 , the mobile phase comprised of 0.1 mol.L -1 ammonium dihydrogen phosphate NH4H2PO4, 1% v/v methanol, pH adjusted to 6.0 by adding KOH. The detection wavelength was 254 nm. The flow rate was 0.8 mL.min -1, and the run time was 10 minutes for ATP-containing samples and 12 minutes for AMP-containing samples. Standard solutions of ATP/AMP in DI water were prepared to determine the calibration curve, and samples were diluted in DI water.

Liquid-state NMR spectroscopy

Liquid-state NMR (lqNMR) spectroscopy was performed at 9.4 T on a Bruker Avance spectrometer equipped with a probe head of 5 mm, and at 7.05 T on a Bruker Avance1 spectrometer with a 5 mm probe head and autosampler. The chemical shifts were referenced to tetramethylsilane for 1 H and to H3PO4 85 wt% solution for 31 P. Table 3-9 lists the main parameters for lqNMR measurements. For trimesate quantification by 1 H NMR, the waiting time between two consecutive scans, or the sum of acquisition and recycle delay, is 5 times longer than the T1 relaxation time of trimesic acid (5 s). For the quantitative measurement of degraded supernatants, 3-(Trimethylsilyl)propionic-2,2,3,3-d4 acid sodium salt, shortly TMSP, was used as an internal standard. Each NMR sample contains 300 µL of the degraded supernatant, 300 µL of D2O, and 20 µL of TMSP standard solution 1.835 mg.mL -1 . The added volumes of degraded supernatant and standard were weighed. Subsequently, the 1 H NMR spectra of these NMR samples were recorded at 7.05 T. The trimesate molar in the NMR tube was calculated using the 1 H NMR peak area of the three aromatic protons of trimesate and of nine protons of TMSP standard. The aromatic 1 H singlet peak of trimesate is around 8.3-8.5 ppm and the methyl 1 H singlet peak of TMSP is around 0 ppm. The formula used is as follows:

𝑛 𝑡𝑟𝑖𝑚𝑒𝑠𝑎𝑡𝑒 = 𝐼 𝐵𝑇𝐶 𝐼 𝑇𝑀𝑆𝑃 × 9𝑛 𝑇𝑀𝑆𝑃 3
, where 𝐼 is the peak area determined by integration using Bruker TopSpin software. From this trimesate content in the NMR tube, the amount of trimesate in the degraded supernatant was deduced.

Quantitative 1 H NMR spectroscopy was also used to work out the initial amount of trimesate in MIL-100. 10 mg of MOF was extracted from stock suspension by replicating the exact washing steps in the nanoMOF degradation procedure. For the iron MOF, the liquid part containing trimesate was separated by centrifugation at 12000 RCF for 15 minutes. Subsequently, the digested solution from MIL-100(Al) or in the digested liquid from MIL-100(Fe) was analyzed by 1 H NMR at 9.4 T. A coaxial insert containing solution of 10.6 mmol.L -1 methyl phosphonic acid (MPA) was put inside the NMR tube (Figure 3-81), thereby, the signal of MPA was recorded in the same NMR spectrum with the signal of trimesate. Taking the molar of MPA as the reference, the analyte's molar was calculated by the following expression: In a similar fashion, quantitative 31 P lqNMR measurement was also done using MPA coaxial insert. The reference 31 P peak of MPA confined in the coaxial insert locates at 28.2 ppm and associates to one phosphorous.

𝑛 𝑡𝑟𝑖𝑚𝑒𝑠𝑎𝑡𝑒 = 𝐼 𝑡𝑟𝑖𝑚𝑒𝑠𝑎𝑡𝑒 𝐼 𝑀𝑃𝐴 × 𝑛 𝑀𝑃𝐴 × 𝐻 𝑀𝑃𝐴 𝐻 𝑡𝑟𝑖𝑚𝑒𝑠𝑎𝑡𝑒
Regarding the NMR experiments at low temperature, 1 H NMR spectra of nanoMOF aqueous suspension were acquired at varying temperature at 9.4 T using single pulse sequence. First, one decreased the temperature by 5-10 o C step from room temperature to less than 0 o C, until observing a sudden significant drop in NMR signal intensity, indicating the freezing of water. Subsequently, the temperature was raised slowly by 1-2 o C, until observing a significant increase of NMR intensity, indicating the melting of water. The NMR spectrum was recorded after waiting for 1-2 minute until the temperature value displayed in the spectrometer was stable. The cooling and heating were done by a constant flow of liquid N2 and heater resistor. From the spectrometer's software, the temperature in the space close to the NMR tube's bottom detected by a thermocouple was set. Temperature was decreased by constantly flowing cold N2 gas (from liquid N2 tank) at 535 L.h -1 inside the NMR spectrometer. The heating level needed to achieve the desired temperature were regulated by the spectrometer's software, which adjusted the heating power of a heater resistor. The temperature inside the NMR tube was calibrated from the temperature displayed by the spectrometer later. A thermocouple probe covered by a thin glass capillary was placed in 1.7 cm height from the bottom of the NMR tube, which contained non-deuterated methanol.

The temperature was read from an external thermometer connected with the thermocouple probe.

Solid-state NMR spectroscopy

Magic-angle spinning (MAS) NMR spectroscopy was performed at 11.7 T on a Bruker NEO NMR WB spectrometer equipped with probe heads of 4, 3.2, and 2.5 mm, except for the 31 P-27 Al 2D NMR measurement, which was recorded at 17.6 T with a 4 mm probe head. The chemical shifts were referenced to tetramethylsilane for 1 H and 13 C, to Al(NO3)3 1 M solution for 27 Al, and to H3PO4 85 wt% solution for 31 P. Fine powder was packed firmly in zirconia rotors. Analysis of ssNMR was done by the Dmfit software 87 . Table 3-10 provides the main acquisition parameters for solid-state NMR (ssNMR) experiments.

X-ray photoelectron spectroscopy

X-ray photoelectron spectroscopy (XPS) experiments were conducted on a Thermofisher Nexsa instrument that has X-ray source from a monochromated Al-Kα anode (1486.6 eV) and a dual flood gun (low energy electron and ion). Droplets (40-50 μl) of 10 mg/ml nanoMOF aqueous suspensions were deposited on silicon wafer. Survey and high energy resolution core level spectra were recorded with 1 and 0.1 eV energy step, respectively. The photoelectron detection used a constant analyzer energy (CAE) mode with 200 or 20 eV pass energy for survey and core level acquisition, respectively. The spot size was 400 µm. Shirley type background subtraction and peak deconvolution were done using the Thermofisher Scientific Avantage software and its "ALTHERMO1" library as sensitivity factor collection.

Deconvolution of core-level C 1s of original nanoMIL-100 is reasoned as follows. The structure of trimesate ligands in nanoMIL-100 has two main types of C that are detectable by XPS C 1s measurement: C sp 2 of the aromatic ring and carboxylate COO. However, the position of C sp 2 coincides with that of adventitious carbon, a layer of 1-2 nm thickness on sample surface composed of carbon-containing organic and inorganic compounds 45 . The sample contamination of adventitious carbon is due to sample exposure to hydrogencarbons and other gases in the atmosphere. The contribution of these adventitious carbons to the peak around 286 eV of C sp 2 must not be neglected. In our case, from the chemical structure of trimesate ligands [C6H5(COO)3] 3-, we estimated that the XPS peak area of aromatic C should be twice as much as that of carboxylate C. All the presented C 1s corelevel spectra were analyzed when fixing this peak area ratio. In addition, the deconvoluted peak area of adventitious carbon are similar among different nanoMOF samples, which is reasonable for contamination species. For the Fe quantification in MIL-100(Fe), area of the Fe 2p peak must be used 88 .

X-ray adsorption near-edge (XANES) spectroscopy

Coordination structure of nanoMIL-100(Fe) was studied by XANES spectroscopy conducted at the SAMBA beamline in Synchrotron SOLEIL, France. The signals were detected at the Kedge of Fe at 7112 eV. The electron storage ring was operated at 2.75 GeV. A Si(220) double crystal monochromator was used, which was calibrated by assigning the first inflection of a Fe foil XANES spectrum to 7112 eV. Depending on the Fe concentration in the sample, the XANES spectra were collected in transmission or fluorescence mode using a Ge 36-pixel fluorescence detector (Canberra). Spectra were collected from 7000 to 8000 eV using a continuous scan acquisition mode, with a 5 eV.s -1 velocity and 0.1 s.point -1 integration time.

Multiple spectra were repeatedly collected for each sample until their average spectrum had a satisfying signal-to-noise ratio. The samples were prepared as dried powders mixed with graphite and pressed into 6 mm diameter pellets. The fluorescence detector was set at 90 o with respect to the incident beam and the sample was set at 45 o with respect to the incident beam. X-ray absorption data processing was done using the Athena program of the Demeter software 89 .

Small-angle X-ray scattering

Small-angle X-ray scattering (SAXS) experiments were conducted at the SWING beamline (Synchrotron SOLEIL, France) using a fixed energy of 12 keV. With two sample-to-detector distances of 0.5 m and 6 m, the corresponding measured range of 𝑞, the scattering vector, is from 1.9E-02 to 2.32E+00 Å -1 and from 9.68E-04 to 1.84E-01 Å -1 . NanoMOF suspension of 10 mg.mL -1 was inserted in cylindrical glass capillaries of 1 mm outer diameter, which was then sealed by flame. For each sample, 10 frames of X-ray scattering image were measured. Data processing, including conversion from scattering image to scattering pattern and curve average was done using the beamline's Foxtrot software. The background SAXS pattern of pure solvent and empty capillary were subtracted from sample's SAXS pattern.

NMR relaxation and relaxometry

The longitudinal relaxation time T1 of 1 H was measured using an inversion-recovery sequence. The signal intensity is a function of recovery delay t:

𝐼 𝐼 0 = 1 -2 exp ( -𝑡 𝑇 1
). The transverse relaxation time T2 was measured using a Hahn echo train for which the recorded signal delay can be expressed as:

𝐼 𝐼 0 = exp ( -𝑡 𝑇 2
).

NMR relaxometry measurement was conducted using a fast-field-cycling (FFC) relaxometer from Stelar, which measures the 1 H longitudinal relaxation rates 𝑅 1 = 4 Developing in situ NMR spectroscopy method for studying the degradation of nanocarriers

Abstract

Using nanoparticles to carry and transport drugs for drug administration is a way to create effective treatments for challenging diseases. For designing a suitable drug delivery system, accurate characterization is indispensable. However, there is a lack of in situ characterization techniques that can analyze the reactions of nanoparticles in their suspension form without separating the suspension and interrupting the reaction. This chapter presents an in situ technique based on NMR spectroscopy that can provide both qualitative and quantitative information about the degradation of metal-organic framework nanocarrier. The colloidal suspension is contained in a hermetic insert fitting a standard 4 mm NMR rotor, which is suitable for magic-angle spinning (MAS) measurement without liquid leakage. During the degradation of nanoMIL-100(Al) in phosphate buffer, in situ 1 H NMR recorded the kinetics of trimesate ligand loss in the solid phase and its release in the liquid phase. Simultaneously in the same measurement, 27 Al NMR observed the ligand replacement process of aluminum and the formation of new aluminum species. Using the same strategy, we also compared the degradation of nanoMIL-100(Al) loaded with two drug analogs, indicating an effect of metal-ligand complexing strength. Furthermore, this technique is also applicable to studying the reaction of paramagnetic nanoMIL-100(Fe) in the liquid phase. Overall, the established in situ NMR technique offers an alternative to ex situ techniques for studying the degradation mechanism of metal-organic framework nanocarriers and possibly other nanoscale drug delivery systems.

Introduction

The development of accurate characterization methods is a vital axe in the field of drug delivery, including characterization on a single particle basis 1 , in-line characterization 2 , and in situ characterization 3 . The need for in situ techniques is especially critical for nanoparticlebased drug delivery systems (DDSs), as many characterization aspects of nanocarrier need proper separation of NPs suspension. For example, determining drug payload requires extraction of free drug molecules into supernatant and bound drug molecules into precipitate. Common separation techniques such as filtration or (ultra)centrifugation risk disturbing the state and the dynamics of the investigating suspension, potentially decreasing the measurement reliability. Therefore, it will be advantageous to study the suspension as it is and acquire results closer to reality than using ex situ techniques.

While chapters 2 and 3 of this thesis explore the reaction mechanisms of nanocarriers using standard NMR spectroscopy techniques, the present chapter is dedicated to developing in situ NMR spectroscopy, a lesser-used method to study nanocarrier degradation.

We want to establish a method to characterize the reactions of nanoparticles (NPs) in their original suspension state, i.e., without separation. Several examples have been reported in the literature using label fluorescent-related techniques or label-free spectroscopy. For instance, the degradation of PEG-dextran hydrogel in different shapes was monitored in vivo and in vitro by following the fluorescent intensity of the polymer conjugated with fluorescent tag 4 . In another example, the kinetics of protein release from polymer hydrogel and the polymer degradation were determined from the evolution of IR peak intensity during in situ measurement using a special IR probe 3 . Such a method that can follow the chemical fingerprint of NPs is not usual in the context of drug delivery applications. In this context, the prospect of in situ NMR spectroscopy is worth investigating because it can provide both qualitative and quantitative information without altering the samples. As demonstrated in the second and third chapters, NMR spectroscopy can provide valuable information about the NP degradation process, including the evolution of molecular structure, the nature and quantity of degradation products, and the reaction rate.

There are also existing in situ studies using NMR spectroscopy. Alonso et al. investigated the lipid NPs colloidal suspension and obtained indirect information about the lipid component existing in the suspension's solid phase, whose signal was invisible in the 1 H lqNMR spectra of the NP suspension 5 (see chapter 1.3.3.1). Kenneth and co-workers attempted to record the formation of a Ni-based metal-organic framework (MOF) in both solid and liquid phases using a solid-state NMR (ssNMR) spectrometer 6 . In situ 31 P NMR showed the disappearance over time of the organic ligands in the liquid phase. In situ 1 H NMR data demonstrated organic linkers' deprotonation and metal coordination during MOF nucleation and growth.

From the 1 H NMR intensity, the crystallization kinetics and the process's activation energies were calculated. Nevertheless, only experiments giving liquid-phase NMR data were helpful for understanding the reaction mechanism, whereas the experiments designated for solidphase species were less informative. To our knowledge, there is no in situ NMR studies of drug carrier yet that can monitor species from different phases simultaneously.

In order to establish the in situ NMR method in this study, nanoMIL-100(Al) was used as the model DDS for its degradation reaction forms products in both solid and liquid phases, as we learned in chapter 3. In addition, the degradation products, i.e., trimesic acid and aluminophosphate, can be detected by 1 H and 27 Al NMR, both of whom are highly sensitive NMR nuclei.

The following chapter consists of three parts:

1) Setting up the in situ NMR method that can monitor species at liquid and solid phases at the same time: NMR acquisition parameters, degradation conditions

2) Validating the concept of in situ NMR for studying the degradation of the model DDSs, nanoMIL-100(Al) and nanoMIL-100(Fe)

3) Applying the established in situ NMR method for studying the degradation of the drugloaded nanoMIL-100(Al)

Set up in situ NMR method

The ultimate goal of the present study is to create an in situ NMR spectroscopy method that can detect species at both solid and liquid phases. The chosen case study, degradation of nanoMIL-100, as studied in chapter 3, produces solvated trimesate in the liquid phase and the remaining coordinated trimesate ligand in solid form. Here comes the first problem, as the solid-and liquid-state substances have distinctive NMR signals and properties, and they are typically measured by different NMR techniques. In Figure 4-1, the trimesic acid solution has a narrow NMR signature peak of 2.5 Hz width, while trimesic acid powder has a broad peak of 50 Hz width under a MAS rate of 10 kHz. To deal with this issue, we learned the strategy of high-resolution magic-angle spinning (HR-MAS) NMR spectroscopy that employs sample spinning around an axis tilted by 54.7 o from the static magnetic field axis.

We still used the probe of solid-state NMR because we wanted to keep the resolution for detecting the solid-phase species. The effect of the spinning rate on the NMR spectra quality is shown in Figure 4-2. We measured MAS NMR spectra of nanoMIL-100(Al) suspension in phosphate buffer at different MAS rates. The higher the spinning rate, the more resolved the NMR peaks and the more reduced the spinning sidebands are. When there is no spinning (static), the signals are broad, and no information can be deduced. Generally, static NMR measurement of a pure liquid or a diluted suspension yields signals of species at the liquid phase, which should be narrow (line width of a few Hz). However, the investigating suspension of nanoMIL-100 is highly concentrated and heterogeneous. Due to the gradient of magnetic susceptibility (explained in section 1.3.3.3), the obtained NMR peaks of liquid-phase species are also broad, like those of a pure solid-state sample 7 . Therefore, a certain MAS rate is required to average this magnetic susceptibility and to increase the resolution of solid species' signals.

At a very slow spinning speed of 60 Hz, the presence of multiple spinning sidebands plagues and crowd the NMR spectrum, making the distinction of overlapping signals impossible. Raising the MAS rate to 0.5 or 2 kHz significantly improved the NMR spectrum quality; narrow lines originating from the liquid phase became visible. Still, at 0.5 kHz, the intensity of the isotropic signal is not maximized, and at 2 kHz, a few spinning sidebands are still close to the peaks of interest. We found that the minimum MAS rate needed to get acceptable NMR spectra quality is 5 kHz. We decided on the MAS rate of 10 kHz to have sufficient 27 Al NMR resolution for solid-state species and to get a temperature inside the sample rotor of 34 o C, close to the human body temperature of 37 o C. Note that a similar effect of MAS on the NMR spectra of nanoMIL-100(Fe) suspension was observed (data not shown), even though only signals of liquid species were detectable. The second issue to be solved is choosing the MOF degradation condition for in situ NMR study. The NMR rotor-insert combo (Figure 4-3) has a maximum volume capacity of 25 µL, which is dramatically smaller than the usual volume used for ex situ NMR study (see chapter 3's experimental method). Other constraints are to get decent NMR sensitivity for both solid and liquid trimesate. The input MIL-100(Al) weight should be at least a few milligrams, while the input PBS/Al ratio must be high enough to release sufficient liquid trimesate. Therefore, to minimize the laboratory-scale degradation condition, concentrated phosphate buffer was used to match the input MOF reactant. In addition, phosphate buffer (PB) was prepared in The optimization was done on a simple solid-liquid mixture. MIL-100(Al) powder was mixed with trimesic acid solution and aluminum nitrate solution to optimize 1 H or 27 Al NMR measurement parameters, respectively. The used values are reported in the experimental section.

In situ NMR study of nanoMOF degradation

There are three ways to conduct in situ NMR measurement reported in the literature: simultaneous, alternating, and sequential NMR. Simultaneous NMR 8 , which can record NMR signals up to four NMR nuclei simultaneously, offers the most accurate time-resolved NMR measurement. However, this acquisition method is challenging to set up because one needs to choose a single set of NMR measurement parameters for different nuclei based on a single parent nuclear. All child nuclei share the same NMR parameters as the parent. Alternating acquisition 9 offers independent implementation of solid-selective and liquidselective pulse sequences, yet, only one nucleus can be recorded once at a time. In comparison, sequential NMR measurement records a series of normal 1D NMR experiments 6 , making it relatively easy to set up. The disadvantage is less accurate kinetics measurement, as there is a time lag between different nuclei's kinetics NMR measurement.

To choose the best compromise between NMR resolution and accuracy, we tested these two acquisition modes to measure nanoMOF degradation kinetics. 3). The two in situ NMR methods result in acceptable quality for 1 H NMR spectra and, more importantly, conservation of the total liquid and solid trimesate peak area after 48 hours, a critical requirement for quantitative NMR. The 1 H NMR peak area fluctuation is 3% and 7% for sequential and simultaneous acquisition, respectively. Nevertheless, the in situ 27 Al NMR spectra sequentially recorded have a better signal-to-noise ratio and peak resolution than those recorded with simultaneous mode (Figure 4-4C and D, respectively). The decrease of the peak representing the original Al species connected with trimesate ligand Al-O-C and the increase of the peak representing degraded Al species connected with phosphate Al-O-P were better observed. For this study of MIL-100(Al), simultaneous acquisition parameters imposed by 1 H NMR are not the most suitable for 27 Al NMR due to different properties between ½-spin and quadrupolar spin nuclei. Meanwhile, for the same experimental time in the sequential acquisition mode, the waiting delay between two scans, also known as relaxation or repetition delay, can be shortened by 1-2 orders of magnitude for 27 Al, which permits to increase the number of scans by 1-2 orders of magnitude. We evaluated that the benefits of real-time measurement that this mode offered do not outweigh the compromise of spectra quality. Therefore, the sequential acquisition was chosen to continue the study to get a better NMR quality and a more flexible choice of NMR parameters in exchange for slightly less accurate kinetics time synchronization. Chapter 4 Developing in situ NMR spectroscopy method for studying the degradation of nanocarriers As a comparison, in situ liquid NMR measurement was also performed on the nanoMIL-100(Al) suspension in PB with the same P/Al ratio as in situ MAS NMR. The sample volume increased from 25 to 600 µL, and the MOF amount was accordingly scaled up from 5 to 120 mg. Figure 4-5 displays the obtained 1 H NMR series in stacked and superimposed representations. In the beginning, there were two peaks, a narrow peak representing released trimesate and a small broad peak, signaling suspension precipitation. As degradation time went on, these two peaks gradually converged into one moderately broad peak. Due to this precipitation and the fact that lqNMR misses the signal from solid species, the total trimesate peak area was not well conserved (peak area variation of 35%). In situ liquid NMR neither observed trimesate in the solid phase nor allowed quantification of trimesate formation. In addition, it could not detect any substantial signal of 27 Al (data not shown). This experiment demonstrates that MAS NMR is necessary for the in situ NMR study of MOF degradation, as it alone can observe both solid-and liquid-phase trimesate. From the sequential in situ NMR measurement, one can plot the kinetics of the MIL-100 degradation reaction (Figure 4-6) by deconvoluting the peak area of solid-phase and liquidphase trimesate species. As their total peak area is nearly conserved over time (3% fluctuation), it represents 100% of trimesate in the sample. The proportion of trimesate in the solid and liquid phases is then calculated by the relative ratio to the total trimesate peak area. Notably, the trimesate release profile obtained from in situ NMR has a similar shape to the one obtained from ex situ NMR and HPLC experiments reported in chapter 3 (see Figure 3-11). Under the degradation condition here, 45% of free trimesate was formed after 48 h, corresponding to 55% of trimesate loss of original MIL-100(Al). Less than 6% standard deviation was obtained from two different synthesis batches of MOF, demonstrating that the in situ measurement was reproducible. Additionally, in situ degradation with pure deuterated water was performed, leading to only a 6% release of trimesate after 48 hours. The time resolution of the trimesate release kinetics is 3 minutes, equal to the 1 H measurement time. Overall, it is feasible to use in situ NMR for the quantitative study of NPs suspension, which could save labor and time for the experimenter compared to ex situ studies when one extracts samples after each time point. The degradation kinetics of nanoMIL-100(Fe) was also studied. The resulting time-resolved NMR spectra were reported in Figure 4-8A. Contrasting to the in situ 1 H NMR spectra of MIL-100(Al), those of MIL-100(Fe) only consist of an increasingly narrow peak representing the liquid-phase trimesate. The peak of solid-phase trimesate was not visible as it directly bonds to paramagnetic Fe(III). That the peak of liquid trimesate from nanoMIL-100(Fe) degradation is observable and has a similar line width to that from nanoMIL-100(Al) indicates that neglectable or no free Fe 3+ cations formed during MIL-100(Fe) degradation. In summary, a proof-of-concept for in situ NMR method for NPs suspension was obtained. An example case study of nanoMOF degradation was demonstrated. The presented in situ NMR method can monitor the reaction products both at liquid and solid phases and offer both qualitative and quantitative analysis.

In situ NMR study of drug-loaded nanoMOF degradation

Here, we demonstrate that the established in situ NMR technique could acquire information about drug release and degradation mechanisms. As a proof-of-concept, nanoMIL-100(Al) were loaded with the previously studied drugs ATP and AMP. Then, these obtained drugloaded MOFs, ATP@nanoMIL-100(Al) and AMP@nanoMIL-100(Al), were degraded in PB (2 mg MOF + 25 uL PB 1 mol.L -1 ).

In situ NMR kinetics experiments were performed. 1 H NMR presented in Figure 4-9 gave information about the drug release process. First, in both 1 H NMR spectra of ATP-and AMPloaded nanoMIL-100(Al), the development of released liquid trimesate and coordinated solid trimesate was similar to what has been observed for the degradation of empty nanoMIL-100(Al) described in the previous section 4.3. Furthermore, in the spectra of degraded AMP@nanoMIL-100(Al) in Figure 4-9A, a signal of AMP was located at 6 ppm, corresponding to the proton of the sugar ring's CH-group connecting to the adenosine 10 . This peak has a line width of 50-70 Hz, which is relatively small to the peak of solid trimesate. Therefore, it was assigned to the liquid-state AMP released from the AMP@MOF. On the opposite, there was no peak at 6 ppm 11 in the 1 H NMR spectra of degraded ATP@nanoMIL-100(Al) (Figure 4-9B), consistent with our previous findings that ATP@nanoMIL-100(Al) is less degraded than the AMP-loaded counterpart due to its stronger drug-carrier connection (see chapter 3.4). Note that other signals of AMP and ATP in the aromatic region overlapped with the signal of nanoMIL-100(Al)'s trimesate ligands, making it difficult to calculate the ligand release and drug release kinetics without additional experiments.

Conclusions and perspectives

To conclude, the present study establishes a new in situ NMR spectroscopy method for monitoring drug delivery processes of drug nanocarriers. Without separating NP suspension, qualitative and quantitative information about the mechanism and kinetics of carrier degradation could be achieved (recap scheme below).

First, the setup of the MAS NMR method was validated by comparison with other fashions of in situ NMR measurements. Furthermore, as a case study, we examined the degradation of an aluminum-based nanoMOF, nanoMIL-100(Al), and were able to determine the reaction kinetics of the ligand release, hence the carrier degradation. Changes in nanoMIL-100(Al)'s coordination environment during its degradation reaction with phosphate could also be detected. In addition, we demonstrated that this method could be useful for MOFs made of paramagnetic iron like nanoMIL-100(Fe), which is usually not adapted for NMR spectroscopy. It is also helpful for comparing two drugs' release processes from nanoMIL-100(Al).

What could be done more includes calibrating the absolute amount of trimesic acid to determine the degradation kinetics of the iron-based MOF and calibrating the absolute amount of AMP to determine the drug release kinetics.

This study could be used as a framework to develop in situ NMR methods for other drug delivery systems.

higher MOF weight was prepared. After the washing step, 100 µL of 1.5 mg.mL -1 ATP/AMP solution was added to the wet MOF pellet first, and the mixture was dispersed by handshaking. Then, the remaining drug solution was added, and the mixture was dispersed by handshaking and pipette. Finally, the suspension was subjected to continuous shaking overnight at room temperature.

For the drug release experiment the next day, 2 mg of drug-loaded MOF was extracted from the drug loading suspension by centrifuging 10 000 RCF for 10 minutes. The obtained pellet was washed with 1 mL D2O, which was later removed by centrifuge 10 000 RCF for 10 minutes. The supernatant should be discarded as much as possible. Finally, the wet MOF pellet was mixed with 25 µL of PB-D2O 1 mol.L -1, and the obtained mixture was placed in an NMR insert as in the degradation experiment. In situ NMR measurement immediately follows.

In situ NMR spectroscopy

Magic-angle spinning (MAS) NMR spectroscopy was performed at 11.7 T on a Bruker NEO NMR WB spectrometer equipped with a 4 mm probe head. The chemical shifts were referenced to tetramethylsilane for 1 H and to Al(NO3)3 1 M solution for 27 Al.

For spinning samples, the suspension was contained in a Bruker disposable insert purchased from CortecNet. For reuse, the sample inside the insert was removed by adsorption on tissue paper. Then, the insert was sonicated separately in ethanol, acetone, and water. Finally, the insert was dried in the oven at 30 o C.

Sequential NMR measurement was performed by direct excitation. At each time point, the 1 H signal was acquired first, followed immediately by 27 Al. Kinetics measurement was programmed using a parameter d2 placed after data acquisition in the 27 Al pulse sequence. The time steps were 30 minutes for the first six measurements from "0" to 3 h, 1 hour for the next three measurements from 3 to 6 h, 2 hours for the next three measurements from 6 to 12 h, 4 hours from 12 to 24 h, and 8 for the measurement from 24 h to 48 h. The value of d2 was calculated by subtracting the 1 H measurement time from the kinetic time step. The acquisition parameters for 1 H and 27 Al NMR were reported in Table 4-1.

General conclusion and perspectives

Prompted by the curiosity about the interaction of nanomaterials in living bodies which has not been fully understood yet, this PhD project has set out to discover the drug delivery mechanisms of nanocarriers. The reported findings about the degradation mechanism of nanocarriers and experimental methods could interest the research community. It has dealt with two different types of nanoparticles (NPs), those made of polymer (chapter 2) and those made of metal-organic framework (chapter 3). NMR spectroscopy, a key tool in this thesis, has been used to identify the degradation and release products of the two studied nanoparticulate systems, as well as to monitor in situ the degradation process of metalorganic framework nanoparticles (chapter 4).

Chapter 2 describes PLA-ITA NPs, a polymer-drug conjugate composed of itaconic acid (ITA) and lactic acid (LAC). We elaborated a new synthesis method to maximize the amount of conjugated drug (ITA) into a polymer (PLA) and obtained knowledge about the resulting NPs' pH-dependent degradation and drug release. Liquid NMR spectroscopy and ESI and MALDI mass spectrometry afforded evidence about PLA-ITA's water-soluble degradation products and residual degraded polymer. We propose that the degradation of the PLA-ITA NPs at pH 7.4 is driven by a pH gradient and occurs by a base-catalyzed ester hydrolysis mechanism. Random scission of ester bonds forms a mixture of oligomers made of ITA and LAC and free ITA and LAC. Very slow degradation of PLA-ITA (less than 20% after 2 months) leads to a low release of ITA drug, presumably due to the absence of hydroxyl end groups.

In addition, TEM was used to follow the morphology of the original and degraded polymer NPs for up to 1 month. Based on characterization at the molecular level by NMR and mass spectrometry and the microscopic level by TEM, bulk erosion was suggested for PLA-ITA NPs.

This study highlights the complex degradation of the polymer NPs chemically conjugate with the drug, which depends on the reactivity of hydrolyzable bonds in the polymer backbone and between the drug and polymer chain. Also, the presented procedures for synthesizing polymer-drug conjugate followed by NP formulation can be helpful for other hydrophilic active molecules that do not suit physical loading in polymer NPs. Qualitative and quantitative analysis of NP's degradation and drug release provided by NMR spectroscopy gives another method to check the indented function in the early stage of DDS development.

The next study should employ a polymer that degrades faster than d,l-PLA, such as the copolymer of lactic acid and glycolic acid PLGA with a low ratio of lactic acid over glycolic acid, or plays with the crystallinity of PLA by decreasing the proportion of l-lactide. Furthermore, the design of two-drug-terminal and one-drug-terminal polymers should be evaluated in terms of hydrolysis rate using the same NMR method established in chapter 2. This NMR method could also be applied to other types of short-chain polymers. Some preliminary experiments were performed on a cyclodextrin-citrate oligomer, showing that NMR spectroscopy could monitor and distinguish the release of free citrate and citrate attached to cyclodextrin.

Chapter 3 dives into the degradation of nanoMOFs (metal-organic framework NPs), highly porous drug carriers comprised of a coordination network between metal and organic ligands. The degradation behaviors of two structural analogs, nanoMIL-100(Al) and nanoMIL-100(Fe), in phosphate buffer saline at pH 7.4 were analyzed at multiple scales. The governing factors, the presence of phosphate and the ratio of phosphate/metal (P/M), were proved by quantitative liquid NMR spectroscopy and XPS. The higher the P/M, the more released free trimesate ligands and the higher degree of nanoMIL-1000 degradation. In addition, solid-state NMR and XANES spectroscopy gave solid evidence of the change in the coordination sphere of degraded MOF. Based on these results, we reported for the first time a stepwise reaction mechanism of the degradation of nanoMIL-100(Al) and (Fe) in a phosphate medium. It starts with the replacement of water ligands and coordination between metal and phosphate, followed by the replacement of trimesate ligands by phosphate and the release of free trimesate. At high P/M ratio, when almost all constituting trimesate ligands were removed, the final MOF degradation products are likely 6-coordinate and 4-coordinate iron/aluminum phosphate. A reaction-limiting step was suggested, in which the newly formed iron/aluminum phosphate could passivate the nanoMIL-100 surface, preventing the degradation reaction's progress.

Furthermore, the degradation process of drug-loaded and surface-coated nanoMIL-100 occurs as a function of complexation strength and affinity between the MOF's metal cations and the modifying agent, e.g., the active substances and the coating molecules. NanoMIL-100(Al) and nanoMIL-100(Fe) have stronger affinity, thus higher loading capacity towards adenosine triphosphate (ATP) than its monophosphate derivative (AMP). However, it renders the degradation of ATP-loaded MOF and the release of ATP more difficult. Likewise, the phosphate-functionalized cyclodextrin (CD-P) was held firmly on the nanoMIL-100 surface but did not help to slow down nanoMIL-100 degradation due to insufficient surface protection.

Having similar degradation behaviors at the molecular scale, the two iron-and aluminumbased MOFs, nevertheless, exhibit different kinds of evolution at the nanoscopic scale. SAXS, TEM, and SEM complementarily follow the size and morphology of MIL-100 NPs during their degradation. Individual MIL-100(Fe) NPs keep their global size and experience gradual surface roughening due to the progressive formation of iron phosphate. Meanwhile, agglomerates of MIL-100(Al) NPs were transformed more suddenly into aggregates of small aluminum phosphate particles. These behaviors were also expressed when examining the interaction between water and nanoMIL-100 in their aqueous suspension by NMR relaxometry. Different mobility of water confined in the MOF cages compared to water in bulk outside MOF was shown by liquid NMR spectroscopy at cold temperature. The paramagnetic Fe(III) in the nanoMIL-100(Fe) dramatically impacts the NMR signature and relaxation of water. The NMR relaxation of water also significantly changed upon degradation of iron MOF, suggesting altered accessibility to surface iron sites. The low-temperature NMR relaxometry study of nanoMIL-100(Al) hints at reduced mobility of water inside MOF.

The differences between nanoMIL-100(Al) and nanoMIL-100(Fe) degradation and erosion processes were argued based on their reactivity towards ligand substitution and the strength of their bonding with trimesate and phosphate coordinating anions. NanoMIL-100(Fe) reacts progressively with phosphate and undergoes "surface" erosion at all P/Fe ratios, while nanoMIL-100(Al) generally reacts faster than its iron analog and, at a high P/Al ratio, undergoes bulk erosion.

The work in chapter 3 opens several perspectives for further investigation on the nanoMOF drug carriers. One should follow the kinetics of drug loading and drug release to understand better how the encapsulated ATP and AMP interact with nanoMOF. This information is critical as loading phosphate-functionalized molecules is often seen in the formulation of nanoMOF. Selective drug encapsulation in small and large cages of nanoMIL-100 or coencapsulation of multiple drugs is also an interesting direction to pursue. In addition, an ongoing effort has been spent on constructing a model to describe reactions in porous media and simulate the obtained NMR relaxometry profiles. More importantly, the target for the subsequent degradation study should be degradation in multiple coordinating anions. It would be challenging to predict the degradability of different MOFs by different anions, which requires exhaustive research on other biodegradable MOFs such as UiO and ZIF.

Chapter 4 raises the question of how to detect both liquid and solid-phase species by a single NMR experiment. We demonstrated a proof-of-concept for an in situ NMR method to monitor a reaction occurring in an NP suspension. The data acquisition mode was chosen from existing in situ NMR measurements. Sequential MAS NMR measurement can detect and quantify released and remaining trimesate during nanoMIL-100(Al) degradation without separating NPs from suspension. Information about the aluminum coordination sphere was also acquired, which agrees with our proposed nanoMOF degradation mechanism in chapter 3. Likewise, this in situ NMR method could be used to study the degradation kinetics of paramagnetic nanoMIL-100(Fe) and the drug release process from nanoMIL-100(Al).

Follow-up research should include determining the calibration for quantitative in situ NMR measurement and drug release kinetics using an internal or external standard. Attempts should be made to apply in situ NMR methods for other drug delivery-related processes, such as drug loading and surface coating. Different active molecules and NMR nuclei like 19 F and 13 C can be considered. More demanding goals include monitoring degradation and drug release from stimuli-responsive drug delivery systems that will need to figure out how to change pH or temperature in situ.

From all the research carried out in the scope of this thesis, there are several common learning points.

First, we learned different applications of NMR spectroscopy for studying drug delivery mechanisms. It has provided insights into two types of drug nanocarriers with different chemical properties. In chapter 2, liquid NMR spectroscopy allowed us to identify and quantify the PLA-ITA's water-soluble degradation products. NMR correlation experiments were crucial to distinguish the mixture of ITA and LAC oligomers. In chapter 3, solid and liquid NMR measurements on phase-separated degradation products proved the reaction mechanism between nanoMIL-100 and phosphate ions. In addition, NMR relaxometry and low-temperature liquid NMR of MIL-100 NP suspension gave complementary methods to follow its degradation process through the dynamics of solvent molecules. In chapter 4, using in situ NMR spectroscopy enabled one to measure NP's degradation products in their native suspension without phase separation.

Nevertheless, the caveat of using NMR spectroscopy for DDS characterization has been reflected. Liquid NMR spectroscopy and in situ NMR spectroscopy measurement needs to use deuterated solvent. Therefore, the effect of deuterated water on the NMR result should be checked when possible. Also, the somewhat modest detection limit of quantitative NMR spectroscopy should be kept in mind when preparing samples for NMR measurement.

Other experimental techniques have also been beneficial for getting a complete characterization of DDSs. Specific spectroscopic techniques such as XPS and XANES provided additional valuable data that would not be easily achieved by NMR spectroscopy, such as when investigating thin surfaces or paramagnetic systems. Nanoscopic properties, e.g., particle size and morphology, assessed by SAXS and TEM/SEM microscopy, are also essential to draw a global picture of the DSS's function from the sub-nano to the micron scale.

On the whole, examples of characterization techniques and varied NMR spectroscopy methods demonstrated in this thesis offer alternative analytical tools for biodegradable nanocarriers. They may serve as a framework for other researchers to study other DDSs.

Second, we learned the different manifestations of degradation in organic and inorganicorganic hybrid materials whose mechanisms are directly related to their innate chemical reactivity. Regarding the study of polymer-drug conjugate NPs, PLA-ITA hydrolyzes in water when changing the external pH; the susceptibility of ester bonds to hydrolysis dictates the degradation rate and drug release. This behavior should apply to degradable polymers consisting of ester bonds or other types of hydrolyzable bonds. Regarding the study of metal-organic framework NPs, nanoMIL-100 materials degrade in phosphate buffer because of their coordination chemistry with phosphate. The degradation products are the free trimesate ligand and metal-phosphate compounds whose coordination number can differ from the original value. The relation between drug encapsulation and carrier degradation was found in both polymer and metal-organic framework nanocarriers. These findings, in agreement with the literature, emphasize the essential role of studying carrier degradation mechanisms at the early stage of the DDS development process.

The general outlook of this thesis is to study the nanocarrier degradation in real and complex biological media, which will require experimental methods to separate nanoparticles from biomolecules.

Each DDS made of unique materials needs its own analysis of degradation and drug release mechanism. We hope that the approach and methodology developed in this thesis may help any scientist interested in designing and characterizing nanocarriers or using NMR spectroscopy to understand their chemistry.

  Discussion on the erosion process of nanoMIL-100 ................................................................... 3.7 Conclusions and perspectives ............................................................................................................. 3.8 Experimental methods............................................................................................................................

Figure 1 - 1 .Figure 1 - 2 .

 1112 Figure 1-1. Schematic illustration of how a drug is delivered to a body by a DDS

Figure 1 - 3 .

 13 Figure 1-3. Schematic drug release kinetics from immediate vs. controlled release DDS. Adapted from 22

Figure 1 - 4 .

 14 Figure 1-4. Journey of a DDS from its preparation to actions

Figure 1 - 6 .

 16 Figure 1-6. Schematic representation of several drug release mechanisms from DDS

Figure 1 - 8 .

 18 Figure 1-8. The effect of the increased surface area provided by nanomaterials: Illustration of the surface area change of solid cubes with the same total volume but different sizes. The starting cube dimension is 1 cm 30

Figure 1 - 9 .

 19 Figure 1-9. Different classes of NPs for drug delivery64

Figure 1 -

 1 Figure 1-10. Chemical formula of several polymers used for drug delivery

Figure 1 - 11 .

 111 Figure 1-11. Examples of biodegradable polyesters

Figure 1 -

 1 Figure 1-12. Impact of PEG surface coverage and conformation on protein adsorption. Protein adsorption (A) happens at low PEG surface density, (B) does not happen when the PEG density is high, and (C) when PEG chains lie on NPs surface 138

Figure 1 -

 1 Figure 1-13. (A) SEM images of porous and pore-closed microspheres and (B) their kinetics of protein release. (C) Scheme for protein sustained delivery system based on the pore-closing treatment of porous PLGA microspheres

Figure 1 - 14 .

 114 Figure 1-14. Ester hydrolysis by acid-catalyzed (AAC2) and base-catalyzed (BAC2) mechanisms 141

Figure 1 -

 1 Figure 1-15. Representative MOF families used as DDSs. (A) Schematic assembly of MOF from secondary building unit (SBU) and organic ligands. (B-D) SBU, organic ligands, and the cages of MIL-100(Fe), ZIF-8(Zn), and UiO-66(Zr), respectively. Illustrations of MIL-100 cages and ZIF-8 cages are from reference 207 and ChemTube3D

Figure 1 - 16 .

 116 Figure 1-16. Size compatibility of drugs and nanoMOF carriers: (A) Encapsulation of two enzymes in nanoPCN-333(Al): superoxide dismutase (SOD) can fit in MOF's medium cages, while catalase (CAT) can fit in only large cages 259 . (B) Encapsulation of two antibiotics in nanoMIL-100(Fe): amoxicillin can cross only the hexagonal window of MOF's large cage, while potassium clavulanate can cross both hexagonal and pentagonal windows 260

Figure 1 -

 1 Figure 1-17. Scheme of the formulation of re-dispersible drug-loaded nanoMOFs and photographs of nanoPCN-222(Zr) coated with mPEG-PO3 287

Figure 1 -

 1 Figure 1-18. Monitoring the degradation of ZIF-8(Zn) microparticles in PBS: (A) EDX elemental map of the original and 24-hour-degraded microparticles, and (B) IR spectra of ZIF-8(Zn) before and after reaction with PBS 306

Figure 1 -

 1 Figure1-19. Degradation studies ofMIL-100(Fe) in phosphate-containing media 1) byLi et al. 308 : TEM images of nanoMIL-100(Fe) (A) before and (B) after incubation in PBS 1.19 mmol.L -1 for 6 h, Raman microscopy images of microMIL-100(Fe) (C) before and (D) after incubation in PBS 11.9 mmol.L -1 for 8 days, and 2) byKoo et al. 311 : TEM images of MIL-100(Fe) (E) as synthesized and (F) after treatment with H3PO4 80 mmol.L-1

Figure 1 -

 1 Figure 1-20. Degradation of drug-loaded nanoMOF: (A-B) Drug release and trimesate release from drug-loaded nanoMIL-100(Fe) in PBS, (C-D) Drug release and trimesate release from drug-loaded nanoMIL-100(Fe) in sulfate. The three drugs are prednisolone (P), prednisolone phosphate (PP), and prednisolone sulfate (PS) 297

X

  -ray photoelectron spectroscopy (XPS) works based on the photoelectric phenomenon. Xray irradiation of 1-5 keV excites electrons from the top layer 1-10 nm deep of the material, making XPS a surface-sensitive technique (Figure1-21). XPS detects the kinetic energy of those emitted electrons, which depends on their binding energy unique for each chemical state of the element (kinetic energy = excitation energy -binding energy). XPS has been used to investigate the surface chemistry of various materials, including polymer and MOF. For example, the presence of the coating mPEG-PO3 on nanoPCN-222(Zr) and its coordination with Zr(IV) is suggested by P 2p and Zr 3d XPS spectra 287 . The atomic composition and C 1s spectra confirmed the presence of CD-P on nanoMIL-100(Fe) 279 .

Figure 1 -

 1 Figure 1-21. Schematic illustration of the working principle of the XPS technique 326

Figure 1 -

 1 Figure 1-22. An illustrative X-ray absorption spectrum at the K-edge of transition metal 328

Figure 1 -

 1 Figure 1-23. Simplified illustration of X-ray scattering by a sample in a SAXS experiment 329

Figure 1 -

 1 Figure 1-24. An illustrative SAXS pattern/profile for a cylindrical particle represented by the blue rectangle in the picture 332

Figure 1 -

 1 Figure 1-25. Illustration of Zeeman splitting of an atomic nucleus with spin ½ 333

Figure 1 -

 1 Figure 1-26. 1 H NMR chemical shift range for representative functional groups of organic compounds 336

Figure 1 -

 1 Figure 1-28. Frequency-domain 1D NMR spectrum obtained by Fourier transformation of time-domain free induction decay (FID) signal

Figure 1 - 30 . 1 H

 1301 Figure 1-30. Temperature evolution of 1 H NMR spectra of isolated liquid NPs based on tripalmitin, lecithin, and PEG-stearate lipid suspended in D2O 351

Figure 1 -

 1 Figure 1-31. (A) Illustration of microemulsion encapsulating the chloramphenicol in the regime of surfactants Tween 20 and Span 20. (B-D) Chemical formula of chloramphenicol, Span 20, and Tween 20, respectively. The moieties in surfactant molecules that are the most sensitive to the benzene group of the drug, are annotated.

Figure 1 -

 1 Figure 1-32. Percentage of lactide units (PLA %) remaining compared to the initial PLA content in the PLA-PMPC copolymer at different degradation times, analyzed by 1 H NMR 359

Figure 1 - 33 .

 133 Figure 1-33. Effect of anisotropy on solid-state NMR spectra 337

Figure 1 -

 1 Figure 1-34. MAS in ssNMR spectroscopy. (A) Schematic illustration of MAS 334 . (B) Effect of MAS on the NMR line shape of the spin-1/2 13 C 367 . (C) Photographs of rotors manufactured by Bruker whose outer diameters are 1.3, 2.5, 3.2, 4, and 7 mm (from left to right) (photo by Thomas Kress under CC BY-SA 4.0)

Figure 1 - 1 H

 11 Figure 1-35. The use of ssNMR spectroscopy to probe the physical state of the drug loaded in DDS. (A) 1 H MAS NMR spectra of ibuprofen crystalline and encapsulated in mesoporous silica MCM-41 with two different pore sizes Ibu-116 and Ibu-35 374 . (B) 1 H MAS NMR spectra of phenylphosphonic acid (ФP) crystalline and loaded in original and functionalized mesoporous silica SBA and SBA-NH 375

Figure 1 -

 1 Figure1-36. 27 Al MQMAS and 31 P-27 Al D-HMQC of nanoMIL-100(Al) (A) pristine, (B) coated with CD-P, (C) loaded with ATP, and (D) loaded with ATP and coated with CD-P. The 1D27 Al spectrum (red line) is on top of the27 Al dimension projection (blue or black) to distinguish surface/bulk Al-O-P contribution264

Figure 1 -

 1 Figure 1-37. Tools and inserts used for HR-MAS NMR experiments, manufactured by Bruker. (A) tool for inserting screw cap insertion, (B) the screw cap of the insert, (C) the upper Teflon® insert, (D) the lower Teflon® insert for 30 μL volume, (E) tool for inserting/extracting the top cap of the insert, (F) the top Kel-F® insert cap, (G) the bottom Kel-F® insert for 12 μL sample volume, (H) plug for disposable insert, (I) disposable 30 μL Kel-F® insert, (J) cap of the 4 mm NMR rotor, (K) disposable insert fit in a 4 mm rotor, (L) 4 mm zirconia rotor. 378

Figure 1 - 38 .

 138 Figure 1-38. Schematic illustration of NMR T1 relaxation process: (A) magnitude of magnetization versus time, (B) macroscopic picture 386

Figure 1 -

 1 Figure 1-40. In situ tracking of degradation of PEG-dextran hydrogel in different shapes. (A) Fluorescent images at different times of disked-shaped hydrogel implanted in mice (B-C) Erosion kinetics of hydrogel converted from time evolution of fluorescent signal intensity, (D) Correlation between in vitroin vivo erosion kinetics

Figure 1 - 41 .

 141 Figure 1-41. The use of FTIR spectroscopy for in situ monitoring drug release. (A) Set up of IR measurement using a fiber optic probe placed inside a three-neck round bottom flask containing NPs suspension. (B) Time evolution of IR peak intensity of the amide C=O stretching band of the protein, reflecting its release over time from acid-degradable (black) and non-degradable (red) hydrogel

Figure 1 -

 1 Figure 1-42. The proportion of m-aminobenzoic acid (m-ABA) in the liquid (+) and solid (x) phases as a function of time, determined from NMR integration

Figure 2 - 1 .

 21 Figure 2-1. Synthetic scheme for the two-directional itaconate PLA polymer

  1 gives the signals in MALDI MS assigned to the PLA-ITA structure. The 13 C NMR, IR, and MALDI mass spectra, the list of 1 H and 13 C NMR chemical shifts, and the IR vibration bands are given in Appendix A.1.

Figure 2 - 2 . 1 H

 221 Figure 2-2. 1 H NMR spectra of the synthesized PLA-ITA, compared with itaconic acid (ITA). Polymer was dissolved in CDCl3, and ITA was dissolved in D2O

Figure 2 -

 2 3 is that the signals of interest of lactic acid (LAC) in the regions 1.1-1.8 ppm and 4.0-5.2 ppm strongly overlap with the resonances of PVA at 1.7 and 3.9 ppm. The signals of PVA are very intense compared to those of ITA and LAC due to the high amounts (4 mg.mL -1 ) of PVA in the NP suspension 39 .

Figure 2 -

 2 Figure 2-3. 1 H NMR spectra of water-soluble degradation products of PLA-ITA single emulsion degraded in pH 7.4 and 5.3 for 1 month 11 days

Figure 2 - 4 . 1 H

 241 Figure 2-4. 1 H NMR spectra of water-soluble degradation products of PLA-ITA nanoprecipitation degraded at pH 7.4 for 2 months and the zooms in the signals of (A, B) ITA and (C, D) LAC. Notation (q) and (d) in the NMR spectrum of LAC means quadruplet and doublet peak, respectively

Figure 2 -

 2 Figure 2-6. 1 H-1 H COSY NMR spectra of water-soluble degradation products of PLA-ITA nanoprecipitation degraded at pH 7.4 for 2 months

Figure 2 - 9 .

 29 Figure 2-9. Positive-mode MALDI mass spectrum of degraded PLA-ITA pellet at pH 5.3 Table 2-4. Peak assignment of positive mode MALDI mass spectra of original and degraded PLA-ITA pellets. Incubation in PBS pH 5.3 was done for 1 month

Figure 2 - 3 Figure 2 -

 232 Figure 2-10. TEM images of PLA-ITA NPs in a 1-month course: (A) original, (B) degraded at pH 7.4, and (C) degraded at pH 5.3

Figure 2 -

 2 Figure2-12 describes the proposed degradation process of PLA-ITA polymer-drug conjugate NPs at physiological pH, as revealed at the microscopic level by TEM and atomic scale by NMR and mass spectrometry analyses. A bulk erosion is suggested because of the following reason. First, at pH 7.4, there was no significant change in particle size and morphology for up to 1 month. The second reason is the limited release of oligomeric species made of LAC and ITA of varied chain lengths resulting from random ester bond cleavage. Finally, the residual polymer had decreasing molecular weight, which would have been more or less maintained if surface erosion had occurred47 .

Figure 2 -

 2 Figure 2-12. Schematic illustration of degradation and drug release at pH 7.4 of PLA-ITA polymer-drug conjugate NPs following bulk erosion mechanism with limited species release

Figure 3 - 1 .

 31 Figure 3-1. Structure of the MIL-100 MOF 9

Figure 3 - 2 .

 32 Figure 3-2. Size distribution of nanoMIL-100(Fe) in absolute ethanol (EtOH). The size distribution is determined based on the scattered light intensity, measured by DLS. Each sample's measurement is recorded three times by the DLS instrument

Figure 3 - 3 .Figure 3 - 4 .

 3334 Figure 3-3. Size distribution of nanoMIL-100(Al) in EtOH (A) before and (B) after size separation. The size distribution is determined based on the scattered light intensity measured by DLS. Each sample's measurement is recorded three times by the DLS instrument

Figure 3 -

 3 5A), and each agglomerate is roughly 1 µm diameter (Figure 3-5B). The individual particles still have faceted shapes and sharp edges like those in nanoMIL-100(Fe) aqueous suspension (see Figure 3-5B, C and E, F of nanoMIL-100(Al) and nanoMIL-100(Fe), respectively). Likewise, zeta-potential (ZP) measurement gave consistent data with TEM and DLS. The absolute ZP value of nanoMIL-100(Al) is nearly 10-time lower than that of nanoMIL-100(Fe) (ZP mean values are -2.5 ± 1.1 mV and +23.5 ± 0.6 mV, respectively).

Figure 3 - 5 .

 35 Figure 3-5. TEM images recorded at different magnifications of (A to C) nanoMIL-100(Al) and (D to F) nanoMIL-100(Fe) in DI water, with MOF concentration of 0.5 mg.mL -1 . The magnification is 600 for (A, D), 10k for (B, E), and 30k for (C, F) images.

Figure 3 - 6 .

 36 Figure 3-6. XRD patterns of original (A) nanoMIL-100(Al) and (B) nanoMIL-100(Fe). The theoretical patterns were calculated using Mercury software and the crystal structure .cif files from reference5,6

Figure 3 - 7 .

 37 Figure 3-7. Nitrogen gas adsorption isotherm of original (A) nanoMIL-100(Al) and B) nanoMIL-100(Fe)

Figure 3 - 8 .

 38 Figure 3-8. IR spectrum of original nanoMIL-100(Al) and nanoMIL-100(Fe). The arrows point to the vibration bands similar to non-coordinated trimesic acid

  Figure 3-9 displays typical TGA curves of MIL-100 structure with three thermal decomposition steps. The first weight loss from room to 100 o C (region 1)corresponds to the removal of water and solvent molecules adsorbed in the framework, followed by gradual evaporation of coordinated water and solvent from 100 to 200 o C (region 2). The most noticeable weight loss (region 3), occurring from roughly 200 to 550 o C for aluminum MOF and to 500 o C for iron MOF, is due to the decomposition of organic trimesate. The plateau at the end of the TGA curve corresponds to the wt% of remaining metal (III) oxide. From the wt% decrease after each step of MOF decomposition, the weight proportions of metal, trimesate ligand, and water in the initial MIL-100 were determined (see details in experimental section 3.8). Moreover, the content of trimesate and metal determined from TGA were crossed-checked with %C and %H determined by elemental analysis and %Al and %Fe measured by ICP, respectively. Carbon is considered to come from only trimesate ligand, and hydrogen is considered to come from trimesate and water. The chemical composition of the initial nanoMIL-100 is summarized in Table3-3.

Figure 3 - 9 .

 39 Figure 3-9. TGA curves of fully hydrated (A) nanoMIL-100(Al), and (B) nanoMIL-100(Fe)

Figure 3 -

 3 Figure 3-10. Zoom in the trimesate peaks in the 1 H NMR spectra of the degraded supernatant of nanoMIL-100(Al) and nanoMIL-100(Fe), recovered after degradation for 48 hours with a P/M ratio of 0.3

Figure 3 -

 3 Figure 3-11. Kinetics of trimesate released from the degradation of (A) nanoMIL-100(Fe) and (B) nanoMIL-100(Al). Several phosphate-to-metal (P/M) molar ratios, calculated from PBS and nanoMOF concentrations, were used, namely: very low (0.3), low (0.5), moderate(5.5), and high(11). Information on the calculation of these ratios is in section 3.8.5

Figure 3 -

 3  P NMR spectra of48-hour degraded supernatant of MIL-100(Al) at very low and low P/Al ratios (0.3 and 0.5, respectively), compared with the starting PBS solution. The signal at 29 ppm belongs to the quantitative reference, methyl phosphonic acid (MPA), in the coaxial insert placed inside the NMR tube.

  structure is intact, and/or reducing pore volume and internal surface area induced by the presence of adsorbed species inside the pore or by structural alteration.

Figure 3 -

 3 Figure 3-14. N2 adsorption isotherms and zoom from 0 to 0.15 p/p0 of (A) degraded nanoMIL-100(Al) and (B) degraded nanoMIL-100(Fe) versus their corresponding original MOFs. NanoMOF samples were incubated in PBS with a low P/M ratio (0.5) for 48 hours. Legend: (Al) and (Fe) stand for MIL-100(Al) and MIL-100(Fe), respectively

Figure 3 -

 3 Figure 3-15. IR spectra of (A) degraded nanoMIL-100(Al) and (B) degraded nanoMIL-100(Fe) as a function of degradation time, in comparison with the intact MOF. NanoMOF were incubated in PBS with a very low P/M ratio (0.3) for 48 hours

Figure 3 -

 3 Figure 3-16. XPS core level spectra of original nanoMIL-100(Al): (A) C 1s, (B) Al 2p, and of original nanoMIL-100(Fe): (C) C 1s, (D) Fe 2p. The adventitious carbon (C adv) identified in the C 1s spectra is due to air-borne contamination 45

Figure 3 - 18 .Figure 3 - 19 .

 318319 Figure 3-18. Time evolution of XPS C 1s core level spectra of (A) degraded nanoMIL-100(Al) and (B) degraded nanoMIL-100(Fe) at a high P/M ratio of 11

Figure 3 -

 3 21A shows a linear relation between the decrease of the C/M ratio in the degraded solid and the increase of the C/M ratio in the degraded liquid, evidencing the cleavage of trimesate ligand due to MOF degradation. Likewise, Figure3-21 suggests a linear tendency between the increase of the P/M ratio and the increase in % of trimesate release, indicating the incorporation/coordination of phosphate in the MIL-100 framework upon its incubation in PBS. Both these two trends suggest that ligand exchange between trimesate and phosphate happens during nanoMIL-100's degradation reaction with PBS.

Figure 3 - 21 .

 321 Figure 3-21. Correlation between XPS atomic composition and % trimesate release in degraded nanoMIL-100(Al) and nanoMIL-100(Fe): (A) C/M ratio, (B) P/M ratio

Figure

  Figure 3-22. Correlation between the P/M ratio determined by EDS (bulk) and by XPS (surface) and % trimesate release in degraded nanoMIL-100(Al) and nanoMIL-100(Fe)

  ring. This charge stabilization enhances the leaving ability of trimesate. The deprotonation of trimesate is also favored at pH 7.4 above its pKa.In short, trimesate linkers in nanoMIL-100 structure are susceptible to replacement by the monohydrogen phosphate anions HPO4 2-in PBS. Phosphate ions can attack the metal centers in MIL-100, causing the cleavage of the M-O-C bond (M = Al or Fe). Free trimesate can be liberated after breaking all of its connections with metal sites, as schematically shown inFigure 3-24.

Figure 3 -

 3 Figure 3-24. Schematic illustration of successive phosphate replacement of coordinated trimesate ligands in MIL-100

Figure 3 -

 3 Figure 3-25. 1 H MAS NMR spectrum of original nanoMIL-100(Al)

Figure 3 -

 3 Figure 3-27. Coordination sphere of Al(III) in MIL-100(Al) studied by high-field ssNMR spectroscopy: (A) 27 Al MAS 1D, and (B) 27 Al MQMAS NMR spectra of original nanoMIL-100(Al) compared with 1D extraction from 2D MQMAS

Figure 3 -

 3 Figure 3-28. (A) 27Al MAS NMR of the degraded nanoMIL-100(Al) compared to the original one, and (B) zoom on the 4-coordinate Al region of degraded, dehydrated, and original nanoMIL-100(Al). Degradation was done using a P/Al ratio of 0.3 (very low ratio) for 48 hours

Figure 3 -

 3 Figure 3-31. (A) 1 H-31 P CP-HETCOR MAS NMR spectrum of the degraded nanoMIL-100(Al), and (B) extract of 1 H dimension, in comparison with 1D 1 H MAS NMR spectrum of the degraded nanoMIL-100(Al). Degradation was done using a P/Al ratio of 0.3 (very low ratio) for 48 hours

Figure 3 - 3 Figure 3 -

 333 Figure 3-33. 27 Al MAS NMR of original and degraded nanoMIL-100(Al) heated until 200 o C. The 48-hour degraded nanoMIL-100(Al) was prepared using a P/Al ratio of 0.3

Figure 3 -

 3 Figure 3-35. Normalized XANES spectra of several representative samples measured for studying nanoMIL-100(Fe) at full scale

Figure 3 -

 3 Figure 3-36. (A) Normalized Fe K-edge XANES spectra and (B) zoom on the pre-edge region of original nanoMIL-100(Fe) and iron(III) acetate. (C, D) Coordination unit structure of MIL-100(Fe) and iron(III) acetate

Figure 3 -

 3 Figure 3-37. Normalized Fe K-edge XANES spectra of intact and degraded nanoMIL-100(Fe) compared with several reference iron compounds. NanoMIL-100(Fe) was degraded at P/Fe of 11 for 48 hours

Figure 3 -

 3 Figure 3-38. (A) Normalized Fe K-edge XANES spectra and (B) zoom on the pre-edge region of degraded nanoMIL-100(Fe) compared with original nanoMIL-100(Fe) and iron(III) phosphate dihydrate. NanoMIL-100(Fe) was degraded at a P/Fe ratio of 11 for 48 hours. (C) Coordination unit structure of iron(III) phosphate dihydrate and its unit cell from 67 plotted with VESTA software.

Figure 3 -

 3 Figure 3-39. (A) Normalized Fe K-edge XANES spectra, and (B) zoom on the pre-edge region of 48hour-degraded nanoMIL-100(Fe), denoted as "degraded(Fe)" compared with several Fe(III) possible degradation products, iron phosphate, iron oxides, and iron oxyhydroxide. The used P/Fe ratio is 11

Figure 3 -

 3 Figure 3-40. (A) Normalized Fe K-edge XANES spectra, and (B) zoom on the pre-edge region of degraded nanoMIL-100(Fe) at different input PBS/Fe ratios (0.3 and 11)

Figure 3 -

 3 Figure 3-41. Proposed degradation mechanism of nanoMIL-100 MOF in phosphate buffer solution. For simplicity, MIL-100(Al) is taken as an example. The same mechanism could be applied toMIL-100(Fe)

Figure 3 - 42 .

 342 Figure 3-42. Chemical structure of (A) adenosine triphosphate (ATP, (B) adenosine monophosphate (AMP, and (C) cyclodextrin phosphate (CD-P)

Figure 3 -

 3 Figure 3-43. Comparison of two MOF concentrations, 5 and 0.5 mg.mL -1 , for loading ATP in nanoMIL-100 carrier (A) Drug payload wt% and (B) encapsulation efficiency (EE%)

Figure 3 - 44 .

 344 Figure 3-44. Liquid chromatogram of supernatant loading ATP and AMP in nanoMIL-100(Fe), compared to trimesic acid solution, indicating the presence of free trimesate and drug

Figure 3 -

 3 Figure 3-45. Comparison of two MOF concentrations, 5 and 0.5 mg.mL -1 , for loading AMP in nanoMIL-100 carrier (A) Drug payload wt% and (B) encapsulation efficiency (EE%)

Figure 3 -

 3 Figure 3-46. Trimesate loss after loading ATP and AMP drugs in nanoMIL-100(Al) and nanoMIL-100(Fe) carriers using two MOF concentrations (A) 5 mg.mL-1 , and (B) 0.5 mg.mL-1

Table 3 - 6 .Figure 3 -

 363 Figure 3-47. TEM images of (A, B) ATP-drug-loaded nanoMIL-100(Al) and nanoMIL-100(Fe), respectively, and (C) AMP-drug-loaded nanoMIL-100(Al)

  -48C). Porcino et al. proved by 31 P-27 Al D-HMQC correlation NMR experiments that this new peak represents the Al coordinating bonded to the terminal γ -phosphate group of ATP through an Al-O-P coordination bond, resulting from the substitution of water ligand 26 . 31 P-27 Al D-HMQC NMR verified that a similar Al-O-P bonding also exists in AMP-loaded-MIL-100(Al) (Figure 3-48D).

Figure 3 -

 3 Figure 3-48. (A)27 Al MAS NMR spectra of MIL-100(Al)-loaded with ATP and AMP. (B) Chemical structure of ATP and AMP. (C-D)27 Al MQMAS and 31 P-27 Al NMR spectra of AMP-loaded MIL-100(Al), respectively.The input drug loading is 30 wt% for both ATP and AMP

Figure 3 -

 3 Figure 3-50. (A) 13 C CPMAS NMR spectra and (B) 1 H MAS NMR spectra of nanoMIL-100(Al) loaded with ATP and AMP, in comparison with the original nanoMIL-100(Al)

  Degradation of the AMP-loaded nanoMIL-100(Al) behaves like the degradation of the original one. Its27 Al MAS NMR spectrum (Figure3-51A) displays the signal around -5 ppm characteristic of 6-coordinate Al-O-P resulting from coordination between MIL-100(Al) and PBS. It also shows the peak of 4-coordinate at 48 ppm (Figure3-51B). Also, the 31 P MAS NMR spectrum (Figure3-52B) indicates an increased signal intensity at -11 ppm, a similar position to the phosphate from PBS coordinated to the framework Al. Still, the NMR peaks belonging to encapsulated AMP were present, indicating the drug was not 100% released. The 31 P MAS NMR spectrum of the degraded AMP-loaded MIL-100(Al) can be reconstructed from a linear combination of normalized spectra of intact AMP-loaded and the degraded original MOF with coefficients 0.27 and 0.73, respectively, implying a high contribution of PBS-O-Al.

Figure 3 -

 3 Figure 3-51. (A) 27 Al MAS NMR spectra of degraded original and drug-loaded MIL-100(Al), and (B) zoom on chemical shift region of 4-and 5-coordinate Al. Degradation P/Al ratio was 0.3, and duration was 48 hours

Figure 3 -

 3 Figure 3-53. TEM images of CD-P-surface-coated nanoMIL-100(Al)

Figure 3 -

 3 Figure 3-54. 13 C CPMAS NMR of CD-P-coated MIL-100(Al), original one and CD-P

Figure 3 -

 3 Figure 3-56. (A) 27 Al MAS NMR spectra of degraded CD-P-coated MIL-100(Al) without and with the loaded ATP, and (B) zoom on chemical shift region of low-coordinate Al species. NanoMIL-100(Al) was degraded at a very low P/Al ratio of 0.3 for 48 hours

Figure 3 -

 3 Figure 3-58. A typical SAXS pattern of nanoMIL-100 suspension of 10 mg.mL -1 in EtOH

Figure 3 -

 3 Figure 3-59. SAXS patterns of original nanoMIL-100(Al) and (Fe) suspension in (A) EtOH and (B) DI water.

Figure 3 - 60 .

 360 Figure 3-60. Time evolution of SAXS patterns of (A, B, C) degraded nanoMIL-100(Fe) and of (D, E, F) degraded nanoMIL-100(Al) suspensions under different input P/M ratios: very low (0.3), low (0.5), and moderate(5.5). For the sample of "0 h" time point, PBS was mixed with nanoMOF in a few seconds, followed by centrifugation to separate the degraded MOF from PBS (details see experimental section)

Figure 3 -

 3 Figure 3-61. Time evolution of SEM images of (A-C) original and degraded nanoMIL-100(Al), (D to F) original and degraded nanoMIL-100(Fe). NanoMOF degradation was done with high input P/M (11) and duration of1.5 h (B and E) and 48 h (C and F). The final concentration of SEM's sample was 0.5 mg.mL-1

Figure 3 - 1 Figure 3 -

 313 Figure 3-62. Time evolution of SEM images of (A-C) original and degraded nanoMIL-100(Al), (D to F) original and degraded nanoMIL-100(Fe). NanoMOF degradation was done with very low input P/M (0.3) and duration of 1.5 h (B and E) and 48 h (C and F). The SEM's sample concentration was 0.5 mg.mL -1

Figure 3 -

 3 Figure 3-65. SAXS patterns of (A) CD-P surface-coated, (B) ATP-drug loaded, and (C) AMP drug-loaded nanoMIL-100(Al) and nanoMIL-100(Fe)

Figure 3 -

 3 Figure 3-66. SAXS patterns of 48-hour-degraded modified versus original nanoMIL-100(Fe) at (A) very low, (B) low, and (C) moderate P/M ratios, which equals 0.3, 0.5, and 5.5, respectively. The curve of the degraded original MIL-100(Fe) was identical to that of the degraded AMP-loaded MIL-100(Fe)

Figure 3 -

 3 Figure 3-67. (A to C) SAXS patterns of 48-hour-degraded ATP-and AMP-loaded nanoMIL-100(Al) at P/Al ratios of 0.3, 0.5, and 5.5, respectively. (D to F) SAXS patterns of 48-hour-degraded CD-P-coated nanoMIL-100(Al) at the same series of P/Al ratios. The patterns of degraded original and intact MOFs were plotted for comparison

Figure 3 -

 3 Figure 3-68. Schematic representation of different types of water existing in nanoMIL-100 suspension

Figure 3 -

 3 69 gives the 1 H NMR spectra of nanoMIL-100(Al) and (Fe) suspension of 1 mg.mL -1 , which have distinctive features. The spectrum of iron MOF consists of a small sharp peak at 4.69 ppm, which was attributed to the residual water HDO inside the coaxial insert solution. A similar narrow peak at 4.71 ppm was observed in the NMR spectrum of pure D2O. The other 1 H peak of iron MOF located at 4.76 ppm corresponds to the water in the nanoMIL-100(Fe) sample, which contains paramagnetic Fe(III) in its structure. Due to the paramagnetic effect, the water peak in iron MOF has a distinguished chemical shift and line width from HDO and pure water. Lone electrons in Fe(III) create an electron-induced magnetic field inside the sample, and the interaction between these local magnetic fields and the externally induced magnetic field of neighboring nuclei creates additional energy transition for nuclear spins.

Figure 3 -

 3 Figure and nanoMIL-100(Fe) suspensions of 1 mg.mL-1 compared with the spectra of water H2O and deuterated water D2O. The NMR tubes had the same coaxial insert containing methylphosphonic acid (MPA) solution in D2O

Figure 3 - 1 Figure 3 -

 313 Figure 3-70. 1 H NMR spectra at room temperature of (A) nanoMIL-100(Al) and (B) nanoMIL-100(Fe) suspension of varying concentrations from 1 to 150 mg.mL -1

Figure 3 -

 3 Figure3-72A reports the modification of 1 H NMR spectra of MIL-100(Al) aqueous suspension when decreasing the sample temperature. As soon as the temperature changed from -9.5 o C to -12 o C, a sudden and significant drop in 1 H intensity took place, signifying that the major proportion of water in the sample was frozen. After this sudden decrease in NMR intensity, we observed a much lower intensity and much broader peak than the one recorded at room temperature. This remaining peak should correspond to water molecules that were not frozen at -12 o C, which should be the water residing at nanoMOF pore and/or close to the internal surface because of the following reasons. First, the low pore volume available for water compared to the suspension's total volume explains the low NMR intensity. Second, the protection of nanoMOF makes these pore/surface water molecules have lower freezing temperatures than freely diffusing bulk water. Most importantly, space confinement of water in MOF cages reduces their mobility, which causes peak broadening and peak shift. Such impact of reduced mobility was already observed for pure water. When the temperature gradually raised from -12 to -2 o C (Figure3-72B), 1 H NMR recorded two signals, one signal of liquid water formed by ice melting, which is narrow and located at 4.9 ppm, and one signal coming from "half-melted" water (photo in Figure3-72B), which is broad and located around 5.8 ppm.

Figure 3 -

 3 Figure 3-72. (A) 1 H NMR spectra of 25 mg.mL -1 nanoMIL-100(Al) aqueous suspension when decreasing the temperature inside the NMR tube. (B) 1 H NMR spectrum of deionized water at -2 o C when elevating temperature. The inset photograph was taken immediately after the tube was removed from the spectrometer

Figure 3 -

 3 Figure 3-73. Proportion % of pore water versus nanoMIL-100(Al) concentration, showing a good fit for linear fitting

Figure 3 -

 3 Figure 3-74. Variable temperature 1 H NMR spectra of 25 mg.mL -1 nanoMIL-100(Al) aqueous suspension when increasing temperature (A) from -12.0 to -2.6 o C, and (B) from 0.1 to 4.3 o C

Figure 3 -

 3 Figure 3-75. Temperature dependence of 1 H NMR intensity of pure water and nanoMIL-100(Al) water suspension at different concentrations (A) temperature range below 0 o C, and (B) the full studied range Likewise, the 1 H NMR spectra of nanoMIL-100(Fe) suspension in water with lowering temperature are reported in Figure 3-76. A marginal peak broadening and a slight peak shift were observed when decreasing temperature from 25 o C. A few minutes after the temperature reached -9.5 o C, the 1 H peak intensity at around 7 ppm dropped to zero, indicating that the bulk water in nanoMIL-100(Fe) was frozen (Figure 3-76A). If observable, the peak of pore water expectedly has a very different chemical shift and broad linewidth due to its close location to paramagnetic atoms. When measuring NMR with an enlarged spectra window (Figure 3-76B), a dramatically broad peak was detected at 46 ppm, a very different chemical shift from pure water, which might correspond to the peak of pore water in nanoMIL-100(Fe). When increasing the temperature, there was no gradual change in the NMR peak area until an abrupt jump was detected at 5 o C. Therefore, we could not plot an intensity versus temperature evolution for MIL-100(Fe) like for MIL-100(Al).

Figure 3 -

 3 Figure 3-76. 1 H NMR spectra of nanoMIL-100(Fe) suspension in water of 50 mg.mL -1 (A) when decreasing temperature from 25 to -9.5 o C, and (B) when increasing temperature from -9.5 to -4 o C with extended chemical shift range

Figure 3 -

 3 Figure 3-77. 1 H NMRD profile of degraded nanoMIL-100(Al) in comparison with original MOF, recorded at -15 o C. The original nanoMIL-100 concentration was 25 mg.mL -1 , which was also used to prepare the degraded suspension

Figure 3 -

 3 Figure 3-78. (A) 1 H NMRD profiles of ATP and AMP-drug loaded nanoMIL-100(Al), in comparison with original MOF, recorded at -15 o C. (B) 1 H NMRD profile of degraded AMP-loaded nanoMIL-100(Al), in comparison with those of original, AMP-loaded, and degraded original MOF, recorded at -15 o C. The initial MOF concentration of all samples was 25 mg.mL -1

Figure 3 -

 3 Figure 3-79. 1 H NMRD profile of degraded nanoMIL-100(Fe) without or with the loaded ATP in comparison with the samples before degradation, recorded at 25 o C. The starting nanoMIL-100 concentration was 25 mg.mL -1 for all samples

Figure 3 - 81 .

 381 Figure 3-81. Illustration of a coaxial insert positioned inside a NMR tube for quantitative NMR measurement of liquid samples 86

1 𝑇 1 when varying the 1 H

 111 Larmor frequency. The outcome of measurement, called nuclear magnetic relaxation dispersion (NMRD) profile, plots the relaxation rate R1 versus the NMR frequency. The SpinMaster relaxometer probes the low frequency domain from 10 kHz to 30 MHz. T1 has been measured using a pre-polarized sequence from 10 kHz to 10 MHz and a non-polarized sequence from 10 MHz to 30 MHz. Polarization was done at 20 MHz for 0.25 s. The recycle delay was set as 0.6 s for nanoMIL-100(Al) at -15 o C and 1 s for nanoMIL-100(Fe) at 25 o C. Glass tube of 5 mm diameter was filled with 1-1.2 mL of nanoMOF suspension, which was then positioned inside the relaxometer. The sample temperature controlled by regulated air flux was set at -15 o C for nanoMIL-100(Al), and at 25 o C for nanoMIL-100(Fe). One waited for temperature stability(15-30 minutes) before starting the acquisition of NMRD profile.

Figure 4 - 1 . 1 H

 411 Figure 4-1. 1 H NMR spectra of trimesic acid solution and trimesic acid powder, measured by liquid NMR and solid NMR spectroscopy, respectively

Figure 4 - 2 . 1 H

 421 Figure 4-2. 1 H MAS NMR spectra of nanoMIL-100(Al) suspension in PB, which were recorded at different MAS rates

  to avoid the signal of water which is especially problematic when measuring micro-scale volume. Also, no salt was added to PB as the required salt concentration of concentrated PBS exceeds their solubility in D2O.

Figure 4 - 3 .

 43 Figure 4-3. Photo of insert-rotor combo for in situ NMR experiments

Figure 4 -

 4 Figure 4-4 presents the in situ NMR spectra acquired by these two fashions as a function of the degradation time of nanoMIL-100(Al) suspension in PB. It was possible to observe the signature of both solid-state and liquid-state trimesate by, assigned to the peaks with line widths larger than 200 Hz and the peaks with line widths 50-70 Hz, respectively. Both 1 H NMR spectra series recorded by simultaneous and sequential acquisition show an increase in liquid trimesate peaks and a decrease in solid trimesate peaks. This phenomenon well fits the proposed trimesate release due to the cleavage of coordinated trimesate ligands (see chapter3.3). The two in situ NMR methods result in acceptable quality for 1 H NMR spectra and, more importantly, conservation of the total liquid and solid trimesate peak area after 48 hours, a critical requirement for quantitative NMR. The 1 H NMR peak area fluctuation is 3% and 7% for sequential and simultaneous acquisition, respectively. Nevertheless, the in situ27 Al NMR spectra sequentially recorded have a better signal-to-noise ratio and peak resolution than those recorded with simultaneous mode (Figure4-4C and D, respectively). The decrease of the peak representing the original Al species connected with trimesate ligand Al-O-C and the increase of the peak representing degraded Al species connected with phosphate Al-O-P were better observed. For this study of MIL-100(Al), simultaneous acquisition parameters imposed by 1 H NMR are not the most suitable for27 Al NMR due to different properties between ½-spin and quadrupolar spin nuclei. Meanwhile, for the same experimental time in the sequential acquisition mode, the waiting delay between two scans, also known as relaxation or repetition delay, can be shortened by 1-2 orders of magnitude for27 Al, which permits to increase the number of scans by 1-2 orders of magnitude. We evaluated that the benefits of real-time measurement that this mode offered do not outweigh the compromise of spectra quality. Therefore, the sequential acquisition was chosen to continue the study to get a better NMR quality and a more flexible choice of NMR parameters in exchange for slightly less accurate kinetics time synchronization.

Figure 4 - 4 .

 44 Figure 4-4. In situ 1 H and 27 Al MAS NMR spectra recorded as a function of time during nanoMIL-100(Al) 48-hour-degradation in PB by sequential (A and C) and simultaneous acquisition (B and D). Condition: 5 mg MOF + 25 µL PB-D2O 400 mmol.L -1

Figure 4 -

 4 Figure 4-5. (A) Stacked and (B) superimposed view of in situ 1 H NMR spectra recorded as a function of time during nanoMIL-100(Al) 48-hour-degradation in PB by sequential acquisition in liquid NMR spectrometer. Condition: 120 mg MOF + 600 µL PB-D2O 400 mmol.L -1

Figure 4 - 6 .

 46 Figure 4-6. Kinetics of nanoMIL-100(Al) degradation in phosphate (PB-D2O), determined by in situ NMR, represented by the kinetics of trimesate ligand release in liquid phase and ligand loss in solid phase, in comparison with the kinetics of ligand release when incubating nanoMIL-100(Al) in D2O. Reaction condition: 5 mg nanoMOF + 25 μL PB-D2O 400 mmol.L -1 or 25 μL D2O. The standard deviation of degradation kinetics in PB-D2O was calculated from two different synthesis batches of MOF

Figure 4 - 7 .

 47 Figure 4-7. Zoom in the Al-4-coordinate chemical shift region from the in situ 27 Al MAS NMR spectra recorded as a function of time during nanoMIL-100(Al) 48-hour-degradation in PB by sequential acquisition (i.e., zoom in Figure 4-4C). Condition: 5 mg MOF + 25 µL PB-D2O 400 mmol.L -1

The 1 H

 1 peak area as a function of time is demonstrated in Figure4-8B, which shows a trend similar to the curve of liquid trimesate release percentage in Figure4-6. The time evolution of peak area could be converted to trimesate release kinetics if one calibrates trimesic acid concentration in the in situ NMR condition. For investigating the reaction kinetics of paramagnetic MOF like MIL-100(Fe), this in situ NMR method could be an alternative to time-and labor-consuming ex situ experiments.

Figure 4 -

 4 Figure 4-8. (A) In situ 1 H MAS NMR spectra recorded and (B) peak area as a function of time of nanoMIL-100(Fe) 48-hour-degradation in PB. Condition: 5 mg MOF + 25 µL PB-D2O 400 mmol.L -1

Figure 4 - 9 .

 49 Figure 4-9. In situ 1 H MAS NMR spectra recorded as a function of time during the 48-hour-degradation of nanoMIL-100(Al) loaded with (A) AMP and (B) ATP. Condition: 2 mg MOF + 25 µL PB-D2O 1 mol.L -1

Figure 4 - 1 Figure 4 - 11 .

 41411 Figure 4-10. In situ 27 Al MAS NMR spectra recorded as a function of time of 48-hour-degradation of nanoMIL-100(Al) loaded with (A) AMP and (B) ATP. Condition: 2 mg MOF + 25 µL PB-D2O 1 mol.L -1

Figure A1. 13 C

 13 Figure A1. of the PLA-ITA

  Figure A2. IR spectrum of the synthesized PLA-ITA

Figure A3 .

 A3 Figure A3. MALDI mass spectrum of the synthesized PLA-ITA recorded without and with supplemented Na + ions

  interaction between ATP/AMP and nanoMIL-100(Fe)/(Al) carriers .......3.4.3 Degradation of drug-loaded nanoMIL-100 ............................................................................. 3.4.4 Surface coating of CD-P on nanoMIL-100 ............................................................................... 3.4.5 Degradation of surface-coated nanoMIL-100 ........................................................................ 3.5 Degradation process of nanoMIL-100(Al) and (Fe) at nanoscale........................................... 3.5.1 Global size and morphology properties of nanoMIL-100(Al) and (Fe) suspensions ..

3.5.2

Evolution of size and surface roughness of degraded

MIL-100 nanoparticles.............

  Drug delivery ................................................................................................................................................

1 INTRODUCTION -LITERATURE REVIEW ........................................................................... 1.1

  Nanocarriers for drug delivery .............................................................................................................. 1.2.1 Generalities about nanoscale drug delivery systems.

  Characterization techniques for drug delivery systems ............................................................... 1.3.1 Routine characterization techniques for nanocarriers ........................................................... 1.3.2 Starting material characterization techniques .........................................................................1.3.3 NMR spectroscopy .............................................................................................................................. 1.3.4 In situ NMR spectroscopy .................................................................................................................

  5 312 . The authors could not detect zirconium in the degradation products by ICP, probably because of the fast hydrolysis of Zr 4+ to amorphous zirconium hydroxide. Another study also concluded similar stability of this MOF in acid, providing no cationic electrophiles are present 313 . DeCoste et al. explained different degradation behaviors of UiO-66 at different pH. Its degradation in NaOH is due to the dissolution of the terephthalate linkers, whereas in HCl, terephthalic acid is protonated, but this process is reversible, thus causing little degradation 203 . 0.1 M bicarbonate, 0.1 M potassium phosphate (KP), 0.1 M PBS, 0.05 Tris, and DMEM (Dulbecco's modified eagle medium) cell media 315 . They found that serum dissolves ZIF-8 and makes holes in the ZIF-8 crystal, causing ligand leaching. Meanwhile, PBS and KP, despite having similar phosphate concentrations, reacted to different extents with ZIF-8. The high NaCl concentration in PBS seems to prevent the release of free zinc into the solution, as shown by ICP measurement.

	In the presence of additional coordinating species, pH can still influence MOF degradation.
	Several authors observed that nano and microMIL-100(Fe), in the presence of phosphate, is
	more stable in acidic pH than in neutral one 297,314 . Christodoulou et al. explained this
	phenomenon by the lower affinity for Fe(III) of H2PO4 -anions existing at pH 5.4 than a
	mixture of H2PO4 -and HPO4 2-anions existing at pH 7.4, as well as by the higher stability of
	Fe in an acidic environment than neutral 297 .

Different degradation behaviors of nanoMOF carriers have been found with various degradation media, emphasizing the need for systematic studies of individual MOFs. For instance, some authors tested ZIF-8 NPs stability at room temperature in several laboratory media, including

Table 2 -

 2

	MS peak			DP
	mode	m/z	Assignment	= x + y
		3132.01		39
		3204.05		40
	[M + Na] +	3276.09		41
		3164.05		41
		3308.13

1. MALDI mass spectrum assignment of the polymer-drug conjugate PLA-ITA. DP refers to the degree of polymerization

Table 2 -

 2

				δ 13 C of ITA
		δ 1 H of ITA	δ 1 H of LAC	(ppm)	δ 13 C of LAC
		(ppm)	(ppm)	C1 / C2 / C3 /	(ppm)
	Substances (with diffusion coefficients)	H3 / H4	H2 / H3	C4 / C5	C1 / C2 / C3
				173.9 / 137.4 /
		6.24 & 5.81		128.2 & 138.1
	ITA (457 μm 2 .s -1 )	/ 3.43		/ 40.8 / 179.1

2. Summary of NMR chemical shift (δ)

and tentative assignment of water-soluble degradation products of PLA-ITA NPs. For LAC fragments, δ is listed from left to right of the polymer chain, i.e., the hydroxyl end group, the main chain unit, and the carboxylic end group, in the order given

Table 2 -

 2 2 (continued).Summary of NMR chemical shift (δ) and tentative assignment of water-soluble degradation products of PLA-ITA NPs. For LAC fragments, δ is listed from left to right of the polymer chain, i.e., the hydroxyl end group, the main chain unit, and the carboxylic end group, in the order given

				δ 13 C of ITA
		δ 1 H of ITA	δ 1 H of LAC		(ppm)	δ 13 C of LAC
		(ppm)	(ppm)	C1 / C2 / C3 /	(ppm)
	Substances (with diffusion coefficients)	H3 / H4	H2 / H3	C4 / C5	C1 / C2 / C3
			4.32 / 1.465
						177.2 & 182.7
						/ 68.9 & 68.3 /
	d,l-LAC (524 μm 2 .s -1 )		4.33 / 1.47			21.1
				179.0 / 138.1 /	172.7 & 179.5
		6.23 & 5.80		128.4 & 138.1	& 177.1 / 67.6
	ITA-LAC	/ 3.43	5.0 / 1.57	/ 40.8 / 174.0	& 74.5 / 17.9
						171.1 & 179.6
						& 177.7 / 73.8
						& 73.2 / 19.7
						& 13.7
			4.56 / 1.54			172.7 & 179.6 & 178.8 / 73.6
	(LAC)2		4.97 / 1.53			& 74.5 / 17.8
						171.1 / 68.3 /
	LAC-propane-diol		4.59 / 1.58			20.33
				173.9 / 137.5 /
		6.25 & 5.81		128.4 / 40.8 /
	β-ITA-propane-diol	/ 3.57		174.8
				169.2 / 138.1 /
		6.25 & 5.81		(a)	/ 39.2 /
	α-ITA-propane-diol	/ 3.57		173.9

Table 2 -

 2 2 (continued).Summary of NMR chemical shift (δ) and tentative assignment of water-soluble degradation products of PLA-ITA NPs. For LAC fragments, δ is listed from left to right of the polymer chain, i.e., the hydroxyl end group, the main chain unit, and the carboxylic end group, in the order given

				δ 13 C of ITA
		δ 1 H of ITA	δ 1 H of LAC	(ppm)	δ 13 C of LAC
		(ppm)	(ppm)	C1 / C2 / C3 /	(ppm)
	Substances (with diffusion coefficients)	H3 / H4	H2 / H3	C4 / C5	C1 / C2 / C3
		6.52 & 6.00	5.4 / 1.65		main unit: (a)
		/ 3.47	5.0 / 1.57		end unit: 172.7
		6.51 & 5.99	5.3 / 1.7	169.2 / 135.37 / 135.49 / 40.4	& 179.5 & 177.1 / 66.9 &
	ITA-d,l-(LAC)2	/ 3.46	5.0 / 1.57	/ 174.9	74.5 / 17.95

(a)

non-identified signal due to spectral overlapping or low intensity

Table 2 -

 2

			Negative
	Species	Chemical structure and formula	ion mode	m/z
	LAC	C3H6O3	[M-H] -	89.3
			[M-H] -	129.1
	ITA	C5H6O4	[M-H-H2O] -	111.2
	LAC3	C9H14O7	[M-3H] 3-	79.3
	ITA-LAC	C8H10O6	[M-H] -	201.1

3. Assignment of ESI(-) mass spectra of water-soluble degradation products of PLA-ITA NPs after 2-month incubation in PBS pH

7.4

Table 2 -

 2 3 (continued). Assignment of ESI(-) mass spectra of water-soluble degradation products of PLA-ITA NPs after 2-month incubation in PBS pH7.4

			Negative
	Species	Chemical structure and formula	ion mode	m/z
	ITA-(LAC)2	C11H14O8	[M-H-H2O] -	255.3

Table 2 -

 2

	PLA-ITA NPs	% ITA release		% LAC release
	degradation	1 month 2 months 1 month 2 months
	pH 7.4	18.4	17.4	8.7	8.8
	pH 5.3	7.7	16.0	2.8	6.7
	2.4.3 Size, morphology, and stability of degraded PLA-ITA NPs

5. Itaconate (ITA) and lactate (LAC) release after PLA-ITA NPs degradation analyzed by quantitative 1 H NMR spectroscopy

Table 2 -

 2 1 H and13 C NMR were measured with a Bruker Avance 400 MHz spectrometer equipped with a 5 mm BBI probe head and operated at a magnetic field strength of 9.4 T, respectively. D2O or CDCl3 were used as deuterated solvents. According to conventional standards, chemical shifts are reported relative to 1% Me4Si in CDCl3 for both 1 H and 13 C. Apparent diffusion coefficients were obtained using an adapted algorithm based on the inverse Laplace transform stabilized by maximum entropy. The experimental parameters of lqNMR spectroscopy are given inTable 2-6.

			6. Parameters for acquiring lqNMR spectra
	NMR	Pulse		Measurement	Information
	experiment	sequence	Main parameters	time	acquired
		single pulse	pulse length 2.8 μs,		group of chemical
		with 30 o flip	recycle delay 2 s,		shifts for H3C-, HC-,
	1 H	angle	acquisition time 3 s a few minutes	H2C=, H2C-
		power-gated			group of chemical
		decoupling			shifts for -CH3, -CH2, -
	13 C	pulse	recycle delay 1 s	overnight	C=CH2, -CH, -COOH
		Bruker			measure translational
		sequence	bipolar and spoil		diffusion coefficient of
	1 H DOSY	"ledbpgs2s"	gradients	1.5 h	released products
					distinguish protons
		Bruker			belonging to the
	1 H{ 1 H} COSY	sequence	recycle delay 1 s	overnight	same species
					identify carbons
			recycle delay 1 s,		coupling with
	1 H{ 13 C}		mixing time 133 ms		different protons up
	HMBC	J-spin echo	= ½ JH-C (7.5 Hz)	overnight	to 3-bond distance
	Quantitative

Degradation mechanisms of nanoparticle-based drug delivery systems: the case of polymer conjugate and metal-organic framework D. L. Mai VUONG -October 2022 151 1 H NMR spectroscopy was based on the calibration of itaconic acid and lactic acid solution in D2O with known concentration.

  Introduction ................................................................................................................................................ 3.2 Synthesis and characterization of nanoMIL-100(Al) and (Fe) ..................................................3.2.1 Size, shape, and water stability .................................................................................................... 3.2.2 Crystallinity, porosity, and functional group ........................................................................... 3.2.3 Chemical composition .................................................................................................................... 3.3 Degradation mechanisms of original nanoMIL-100(Al) and (Fe) at molecular scale ..... 3.3.1 Reaction kinetics at liquid phase ................................................................................................. 3.3.2 Changes in crystallinity, porosity, and functional group .................................................... 3.3.3 Chemical environment and atomic composition at nanoMOF surface ......................... 3.3.4 Discussion on the ligand exchange process............................................................................. 3.3.5 Evolution of metal coordination sphere in nanoMIL-100(Al) ............................................ 3.3.6 Evolution of metal coordination sphere in nanoMIL-100(Fe) ............................................ Degradation and drug release mechanisms of modified nanoMIL-100(Al) and (Fe) ..... Degradation process of nanoMIL-100(Al) and (Fe) at nanoscale........................................... Discussion on the erosion process of nanoMIL-100 ................................................................... 3.7 Conclusions and perspectives ............................................................................................................. 3.8 Experimental methods............................................................................................................................

3 STUDYING DEGRADATION MECHANISM OF METAL

-ORGANIC FRAMEWORK NANOCARRIERS ................................................................................................................... 3.1 3.3.7 Discussion on the degradation mechanism of nanoMIL-100 in phosphate and possible degradation products .

............................................................................................................... 3.4 3.4

.1 Optimization of drug loading formulation for ATP/AMP drug and nanoMIL-100(Fe)/(Al) carriers .................................................................................................................................... 3.4.2 Drug-carrier interaction between ATP/AMP and nanoMIL-100(Fe)/(Al) carriers ....... 3.4.3 Degradation of drug-loaded nanoMIL-100 ............................................................................. 3.4.4 Surface coating of CD-P on nanoMIL-100 ............................................................................... 3.4.5 Degradation of surface-coated nanoMIL-100 ........................................................................ 3.5 3.5.1 Global size and morphology properties of nanoMIL-100(Al) and (Fe) suspensions .. 3.5.2 Evolution of size and surface roughness of degraded MIL-100 nanoparticles............. 3.5.3 Effect of drug loading and surface coating on morphological evolution of degraded MIL-100 nanoparticles ...............................................................................................................................

3.5.4

Interaction between nanoMIL-100 and water in the nanoparticle suspension ...........

3.5.5 Behavior of water in degraded nanoMIL-100(Al) and (Fe) ................................................. 3.6 3.8.1 Chemicals ............................................................................................................................................ 3.8.2

Synthesis of nanoMIL-100 ............................................................................................................. 3.8.3 Drug encapsulation.......................................................................................................................... 3.8.4 Surface modification ........................................................................................................................ 3.8.5 Degradation of nanoMIL-100 .......................................................................................................

3.8.6 Characterization techniques for nanoMIL-100 .......................................................................

Table 3 -

 3

1. Assignment of IR spectra of nanoMIL-100(Al) and nanoMIL-100(Fe)

Wave number (cm -1 ) Assignment Reference

	(a.u)
	Transmittance
	4000	3200	1600	800
	3250	O-H stretching		35,36
	3096	Aromatic C-H stretching	35
	1715	Carboxylic C=O stretching	37
	1625-1630	Carboxylate C=O stretching	36,38
	1450-1460	Aromatic C-C stretching	35,37
	1380-1400	Carboxylate COO stretching	35
	1276	Carboxylic C-OH stretching	37
	760-770 and 700-710	Aromatic C-H out-of-plane	36

Table 3 -

 3 2. Particle size in suspension, zeta-potential, and BET surface area of nanoMIL-100(Al) and nanoMIL-100(Fe). Z-average, which represents the particle's mean hydrodynamic diameter, is calculated by the DLS instrument from the particle's scattered light time evolution39 . PdI means polydispersity index

		Size in EtOH	Size in DI water		BET surface
		z_average		z_average			area
		(nm)	PdI	(nm)	PdI	ZP (mV)	(m 2 .g -1 )
	nanoMIL-100(Al)	147 ± 14 0.10 ± 0.02 2622 ± 680 0.27 ± 0.11	-2.5 ± 1.1 1632 ± 325
	nanoMIL-100(Fe)	226 ± 15 0.13 ± 0.03	271 ± 34 0.18 ± 0.04 +23.5 ± 0.6 1621 ± 327

Table 3 -

 3

	3. Chemical composition of nanoMIL-100(Al) and nanoMIL-100(Fe), whose general formula is
	[M3O(H2O)2OH(C6H3(CO2)3)2].nH2O.xC6H3(COOH)3, determined from TGA, elemental analysis, and ICP
			wt%		n_H2O non-	x_extra
			trimesate wt% metal	coordinated	trimesic acid
	nanoMIL-100(Fe)	dried from EtOH 52.4 ± 4.1	18.5 ± 2.8	7.3 ± 0.4	0.2
	fully rehydrated	dried from H2O	53.5 ± 5.0	19.6 ± 2.9	6.0 ± 1.0	0.1
	nanoMIL-100(Al)	dried from EtOH 61.4 ± 7.2	9.7 ± 0.6	5.7 ± 1.1	0.5
	fully rehydrated	dried from H2O	62.0 ± 6.3	9.9 ± 0.3	5.8 ± 1.0	0.5

Table 3 -

 3

	4. Drug payload and encapsulation efficiency (EE%) of ATP in (A, B) nanoMIL-100(Al) and (C, D)
	nanoMIL-100(Fe) using different input drug loadings and MOF concentrations. The input MOF
		concentration is 5 mg.mL -1 in (A, C) and 0.5 mg.mL -1 in (B, D)
				nanoMIL-100(Al)
	(A) 5 mg/mL MOF			(B) 0.5 mg/mL MOF
	Input ATP	Payload		%BTC	Input ATP	Payload		%BTC
	loading wt%	wt%	EE%	loss	loading wt%	wt%	EE%	loss
	30	19	65	17	35	17	47	20
	25	20	80	9	27	13	50	15
	22	18	81	9	26	13	51	14
	19	15	81	10	19	9	48	10
				nanoMIL100(Fe)
	(C) 5 mg/mL MOF			(D) 0.5 mg/mL MOF
	Input ATP	Payload		%BTC	Input ATP	Payload		%BTC
	loading wt%	wt%	EE%	loss	loading wt%	wt%	EE%	loss
	34	33	98	17	36	26	72	12
	31	27	87	16	35	30	87	17
	26	26	98	15	28	27	97	15
	22	22	100	11	22	22	99	13

Table 3 -

 3 5. Drug payload and encapsulation efficiency (EE%) of AMP in (A, B) nanoMIL-100(Al) and (C, D) nanoMIL-100(Fe) using different input drug loadings and MOF concentrations. The input MOFconcentration is 5 mg.mL -1 in (A, C) and 0.5 mg.mL-1 in (B, D)

	nanoMIL-100(Al)

  2-have similar values of metal complex formation constant76 . Therefore, they compete to coordinate the MIL-100 framework's Al cations. Once PBS replaces the labile water ligands, the Al-O-C bonds with trimesate are weakened and probably broken first, as trimesate ligand has weaker binding strength with Al than AMP and PBS. When more and more 6-coordinate Al-O-PBS are formed, the coordination bond Al-O-AMP is weakened. Then PBS can easily replace the coordinated AMP, leading to AMP release. The drug will be totally released when phosphate ions substitute all coordinated AMP. The main degradation products of AMP-loaded nanoMIL-100(Al) are 6-coordinate aluminum phosphate, free trimesate, and free AMP. The 4-coordinate Al species could be formed as a side product during the dissociation of Al-Otrimesate and Al-O-AMP bonds.Contrary to AMP-loaded, the degradation reaction of the ATP-loaded nanoMIL-100(Al) exhibits distinctive features from the original one. No sign of low-coordinate Al was detected (Figure3-51B).27 Al NMR (Figure3-51A) again shows the 6-coordinate Al coordinated with PBS around -2 ppm as the main degradation product, whereas the signal of Al bonded with ATP can still be observed. Likewise, the 31 P MAS NMR spectrum in Figure3-52A shows both a peak of coordinated PBS phosphate and a main contribution of ATP. Both27 Al and 31 P NMR data suggest that nanoMIL-100(Al) loaded with ATP is less degraded than the empty one, and only a part of ATP is liberated from the framework. The triphosphate ATP is a much stronger complexing ligand towards Al(III) than the hydrogen phosphate and dihydrogen phosphate ions77 , meaning that MIL-100(Al) forms a more stable coordination complex with ATP than with phosphate ions76 . Therefore, it is difficult to break the coordination between ATP and Al(III) by solely PBS. Nonetheless, the robust Al-O-P bond between Al(III) and ATP makes the Al-O-C bond between Al(III) and trimesate linker more easily attacked by PBS. We hypothesize that PBS preferentially attacks the Al-O-C bond with trimesate first, forming 6-coordinate Al-O-P in the region -2 to -5 ppm as observed in the27 Al NMR spectrum. Once the Al center forms the Al-O-P bond with PBS, its coordination bond with ATP is weakened. As a result, ATP could be displaced by PBS and released from MIL-100(Al). Also, the steric effect could play a role. Bulky ATP possibly hinders phosphate's diffusion in the MIL-100(Al) cage and protects the trimer from phosphate's attack, thereby limiting or preventing trimesate substitution.

Table 3 -

 3

	7. Particle size and zeta-potential of surface-coated nanoMIL-100(Al) and nanoMIL-100(Fe)
		Size in DI water
		z_average (nm)	PdI	ZP (mV)
	nanoMIL-100(Al)	2622 ± 680	0.27 ± 0.11	-2.5 ± 1.1
	nanoMIL-100(Fe)	271 ± 34	0.18 ± 0.04	+23.5 ± 0.6
	CD-P-coated nanoMIL-100(Al)	291 ± 58	0.12 ± 0.04	-16.3 ± 3.3
	CD-P-coated nanoMIL-100(Fe)	237 ± 6	0.14 ± 0.01	-36.9 ± 0.9

Table 3 -

 3 8. Conditions of various P/M (M = metal) ratios used for degradation experiments of nanoMIL-100(Al) and nanoMIL-100(Fe) in phosphate. P/Al and P/Fe ratios are the molar ratio of PBS's phosphate salts over MOF's Al and Fe, respectively

		P/Al	P/Fe	MOF concentration	PBS concentration	MOF weight
	Name	ratio	ratio	(mg.mL -1 )	(mmol.mL -1 )	(mg)
	very low	0.3	0.3	0.25	10	3
	low	0.5	0.6	0.5	10	0.5 or 5
	medium	5.5	5.8	0.5	1	0.5 or 5
	high	11.0	11.7	10	10	10

Table 3 -

 3 9. Parameters of quantitative lqNMR experiments

	NMR	Pulse		Experiment	Information
	nucleus	sequence	Main parameters	time	acquired
		Single pulse	Recycle delay 25 s, acquisition time		Presence and
		with solvent	1.6 s, pulse length 8 μs (90 o ), pulse		quantity of
	H (9.4 T)	suppression	power 30 W	≤ 15 min	trimesate
					Presence and
			Recycle delay 6 s, acquisition time 1.6		quantity of PBS,
		Proton	s, pulse length 16 μs (90 o ), pulse		CD-P, ATP, and
	P (9.4 T)	decoupling	power 150 W	30 min-4 h	AMP
			Recycle delay 6.5 s, delay before
		Single pulse	signal acquisition 18.5 s, acquisition		Presence and
		with solvent	time 1.95 s, pulse length 14.7 μs,		quantity of
	H (7.05 T)	suppression	pulse power 27 W	≤ 15 min	trimesate

  To the obtained MOF pellet, 60 µL of 14.3 mol.L -1 NaOH solution in D2O was added. The resulting mixture was sonicated for 15-30 minutes at 50 o C. Then, 540 µL D2O was added, and the digested sample was sonicated for extra 15-30 minutes at 50

o C. A clear orange solution was obtained from nanoMIL-100(Al) digestion, while digested nanoMIL-100(Fe) contained clear transparent liquid and brown precipitate of iron hydroxide, which was further heated at 80 o C overnight.

Table 3 -

 3 10. Parameters of ssNMR experiments

	NMR	MAS	Pulse		Experiment
	nucleus	(kHz)	sequence	Main parameters	time	Information acquired
				Pulse length 3.35 μs,
				pulse power 100 W
				(MAS 10 kHz) and 50		Presence and integrity
	H	10, 20	Hahn Echo	W (MAS 20 kHz)	< 1 min	of linker
			Cross-	Contact time 2 ms,
			polarization	proton decoupling		Presence and integrity
	C	10	(CP)	SPINAL-64 at 80 W	> 1 h	of linker/drug/coating
				Pulse length 2.8 μs,
			Single pulse,	pulse power 40 W		Coordination number
		12.5,	MQMAS with	(MAS 12.5 kHz) and 50		and coordination sphere
	Al	20	z-filter	W (MAS 20 kHz)	0.5-4 h	of aluminum
				Contact time 3 ms,
			Single pulse,	proton decoupling		Nature of incorporated
	P	12.5	CP, hpdec	SPINAL-64 at 30 W	> 30 min	phosphate species
			CP-HETCOR	Contact time 3 ms
				R421 recoupling, 1.2		Spatial proximity (a few
				ms recoupling time,		Å) between protons and
	H-27 Al	20	D-HMQC	recycle delay 0.5 s	> 8 h	MIL-100 aluminum
				Contact time 3 ms, 1 H		Spatial proximity (a few
				decoupling SPINAL-64		Å) between proton and
	H-31 P	12.5	CP-HETCOR	at 30 W	> 1 h	phosphorus species
				R421 recoupling, 1.7
				ms recoupling time,		Spatial proximity (a few
				0.25 s recycle delay, 1 H		Å) between phosphorus
				decoupling SPINAL-64		species and MIL-100
	P-27 Al	14	D-HMQC	at 50 W	1 day	aluminum

Table 4 -

 4 1. Experimental parameters of in situ NMR measurement

	NMR	MAS	Pulse		Measurement	Information
	nucleus	(kHz)	sequence	Main parameters	time	acquired
						presence	of
				pulse length 3.35 µs,		coordinated
				pulse power 100 W,		trimesate and
			single	delay 4.4 s, acquisition		release of free
	1 H	10	pulse zg	time 1.6 s	≈ 3 minutes	trimesate
				pulse length 3.5 µs,		coordination
			single	pulse power 40 W,		sphere of Al in
	27 Al	10	pulse zg	acquisition time 0.01 s ≈ 27 minutes	nanoMIL-100

Table A2 .

 A2 13 C NMR chemical shift of the synthesized PLA-ITA polymer conjugate

	Chemical shift (ppm)	Assignment
	170.1-169.7 (m (a) )	COO
	132.8	C in H2C=C
	130.9	CH2 in H2C=C
	69.6-69.1 (m)	CH in OCH(CH3)CO2
	61.8	CH2 in OCOCH2CH2CH2OCO
	36.9	CH2 in (CH2=C)CH2CO2H
	27.9	CH2 in OCOCH2CH2CH2OCO
	16.76	CH3 in OCH(CH3)CO2

Table A4

 A4

. Mean hydrodynamic diameter (z_average) and polydispersity index (PdI) of the investigated NPs formulations for PLA-ITA Method [PLA-

ITA] (mg.mL -1 ) Organic and aqueous phase z_average (nm) PdI Observation

	Single emulsion	5		301 ± 12	0.13 ± 0.02
		10		310 ± 7	0.12 ± 0.09
		15	1.5 mL	331 ± 12	0.05 ± 0.03
			dichloromethane,			some precipitate
			4 mL PVA 4 wt%			after solvent
		20	aqueous solution	323 ± 14	0.07 ± 0.01	evaporation
			1.5 mL acetone, 4
			mL PVA 4 wt%
	Nanoprecipitation 5, 10, 15, 20	aqueous solution aggregate
	Nanoprecipitation		0.75 mL acetone,
	small volume	5	2 mL DI water	343 ± 36	0.49 ± 0.07

D. L. Mai VUONG -October 2022

D. L. Mai VUONG -October 2022

Chapter 1 Introduction -Literature review Chapter 1 Introduction -Literature review Chapter 3 Studying degradation mechanism of metal-organic framework nanocarriers

List of publications and author contribution

A part of this thesis's chapter 2 has been published in the academic journal International Journal of Molecular Sciences in the article: Morphology and size of 0.5 mg.mL -1 nanoMOF suspension were observed by transmission electron microscopy (TEM) under a 120 kV JEOL JEM-1400 microscope and scanning electron microscopy (SEM) under a 15 kV JEOL JSM-7001F microscope. For TEM, 3 µL of sample suspension was deposited on copper grids and observed without further staining. For SEM, 40-50 µL of sample was deposited on silicon wafers and metalized with 5 nm-thick gold film. Energy dispersive X-Ray spectroscopy (EDS) coupled with SEM, using a detector from Oxford Instruments, was used to analyze the elemental composition of MOF.

The crystallinity of MOF powder was checked by powder X-ray diffraction (PXRD) at λ(Cu Kα) of 1.5406 Å, using a Siemens D5000 diffractometer. The porosity of MOF was evaluated from the nitrogen gas adsorption measured by a Micromeritics Tristar II Plus or porosimeter. Before the measurement, the sample was degassed under vacuum at 100 o C overnight. Infrared spectroscopy (IR) spectra of MOF powder were recorded by a ThermoScientific Nicolet 6700 IR spectrometer with an attenuated total reflectance accessory.

The chemical composition of MOF was analyzed by thermogravimetric analysis (TGA), elemental analysis, and inductively coupled plasma optical emission spectroscopy (ICP-OES). TGA curve was recorded by a Mettler-Toledo thermal analyzer. The sample was heated from 30 to 900 o C in O2 (gas flow rate 20 mL.min -1 ). The content of water, trimesate ligand, and metal in MIL-100 were deduced from the TGA curve and the generally accepted chemical formula M3O(H2O)2OH(C6H3(CO2)3)2 (for M = Fe, Al) 6,8 . First, the wt% of M2O3 (M = Al or Fe) and trimesate in MIL-100 samples were estimated from the values in the TGA curve as marked in Figure 3-80. The metal content in MOF was deduced from the wt% of M2O3. Then, the wt% of coordinated trimesate was calculated from the wt% of metal based on the molar ratio M/trimesate = 1.5. The difference between the wt% trimesate read from the TGA curve and the wt% of coordinated trimesate belongs to the extra-framework trimesic acid trapped in the framework. Next, the molecular weight (MW) of MIL-100 was estimated from the wt% of metal as follows:

Degradation mechanisms of nanoparticle-based drug delivery systems: the case of polymer conjugate and metal-organic framework

Experimental methods

All reagents and solvents were used without any purification.

NanoMIL-100 synthesis and characterization

NanoMIL-100(Al) and nanoMIl-100(Fe) were synthesized and characterized as reported in chapter 3. Briefly, the two nanoMOFs were synthesized by microwave synthesis technique 13,14 . The obtained products were characterized by DLS, XRD, and BET to check their NP size, MOF porosity, and MOF crystallinity.

Degradation of nanoMIL-100

A small-volume reaction between nanoMIL-100 and phosphate buffer was carried out to adapt for the NMR insert needed for in situ NMR measurement. Basically, in a roundbottom 2 mL Eppendorf tube, 5 mg of MOF was mixed with 25 μL phosphate buffer (PB) in deuterated water (D2O)

The phosphate buffer in deuterated water (PB-D2O) was prepared following a published recipe 15 . The stock solution of KH2PO4 1 mol.L -1 (stock A) was prepared by dissolving 1.36 g of the anhydrous salt in 10 mL D2O. For the stock solution of Na2HPO4 1 mol.L -1 (stock B), portion by portion of the anhydrous salt was gradually added to D2O under vigorous stirring and heating at 35-45 o C in a water bath. Afterward, to get PB-D2O 1 mol.L -1 , one gradually pipettes 2.8 mL stock A to 7.7 mL stock B. The stock solutions and PB were stored at room temperature.

Prior to the degradation reaction, 5 mg wetted MOF pellet was centrifuged (11 000 RCF for 15 minutes) from its stock suspension in absolute ethanol (EtOH abs). Then, the obtained pellet was washed twice in D2O as follows: the wet pellet was carefully redispersed in 100 µL D2O first (handshaking and gentle vortex), then 1 mL D2O was added, and the suspension was remixed. D2O was removed by centrifugation (10 000 RCF for 10-15 minutes) to get a clear and non-blurry supernatant. After the last centrifugation step, the supernatant should be discarded as much as possible. Finally, 25 µL of PB-D2O 400 mmol.L -1 was added to the Eppendorf tube. The suspension in PB was dispersed rapidly by vortex. 25 µL of this suspension was injected into the NMR insert using a pipette.

The insert was sealed by a screw cap and then placed inside the NMR rotor for subsequent NMR measurements.

Drug loading of nanoMIL-100 and drug release

The drug loading procedure reported in chapter 3 was adapted to prepare a small volume of samples. Briefly, 5 mg.mL -1 MOF suspension was mixed with 1.5 mg.mL -1 drug solution, corresponding to 30 wt% input drug loading. The volume of the drug solution was determined from the needed weight of MOF. Before mixing with the drug solution, the removal of EtOH abs and washing with D2O were carried out in the same way as described in the degradation experiment. The amount of D2O for washing could be increased twice if a