Dr Mario Menendez-Miranda

Ms. Ioanna Vikramjeet Singh

Prof Philippe Roger

Prof Nadine Aubry-Barroca

Prof Caroline Aymes

Dr Antonio Vargas-Berenguel

Dr Juan M Casas

Dr Luc

Dex Dextran

Dr Ruxandra Gref
email: ruxandra.gref@universite-paris-saclay.fr.

Cy5 cyanine 5 DCM Dichloromethane DCTB trans-2-[3-(4-tert-butylphenyl)-2-methyl-2-propenylidene]malononitrile DDS drug delivery systems

Keywords: Porous drug-loaded nanoparticles with engineered coatings core-shell nanoparticles, polymer/oligomer synthesis, "stealth" drug carrier, anticancer nanoparticules coeur-couronne, synthèse de polymère/oligomère

L'utilisation de nanomatériaux hybrides organiques-inorganiques (nanoMOFs) pour le diagnostic et le traitement de maladies graves, telles que le cancer et les infections, suscite un intérêt croissant. Cependant, immédiatement après leur administration, les nanoMOFs interagissent fortement avec leur environnement physiologique, notamment avec les protéines plasmatiques. Par conséquent, ils sont rapidement éliminés de la circulation sanguine et ne peuvent pas atteindre leur cible, par exemple, la tumeur à traiter. Cette élimination prématurée limite considérablement l'utilisation des nanoMOFs pour des applications pratiques.

Pour relever les défis liés à la séquestration des nanoMOFs par le système immunitaire, nous avons développé plusieures stratégies pour recouvrir les nanoMOFs de revêtements "furtifs". Pour ce faire, des chaînes de poly(éthylène glycol) (PEG) ont été greffées de manière covalente à des copolymères qui ont été adsorbés sur les nanoMOFs.

Cependant, par rapport aux nanoparticules denses non-poreuses, la modification de surface des nanoMOFs hautement poreux pose des verrous technologiques et est beaucoup plus délicate. En effet, il a été démontré que les chaînes de PEG peuvent pénétrer dans les pores des nanoMOFs, enduisant une libération prématurée des molécules actives encapsulées. Par conséquent, dans cette étude, nous avons développé différentes stratégies pour modifier la surface des particules de nanoMOFs avec des revêtements multifonctionnels sans perturber leur capacité d'encapsulation de médicaments. En même temps, nous nous sommes attachés à mettre au point des méthodes de recouvrement et d'encapsulation par chimie verte, sans emploi de solvants organiques.

La première stratégie a consisté à synthétiser des copolymères de type « peigne » constitués de dextrane (DEX), de PEG et d'alendronate (ALN). Le DEX joue le rôle de macromolécule volumineuse, empêchant les chaînes de PEG de pénétrer à l'intérieur des nanoMOFs, tandis que les chainons d'ALN sont des groupements complexant le fer en surface des nanoMOFs, pour ancrer fermement les copolymères de type « peigne » de manière rapide et stable aux nanoMOFs. Les copolymères DEX-ALN-PEG ont été obtenus par chimie click en faisant varier les nombre d'unités PEG et ALN qui ont été greffées. Ils ont été caractérisés par un ensemble de techniques physicochimiques, puis adsorbés spontanément sur les nanoMOFs par simple incubation dans l'eau, sans solvant ni agent de couplage.

Comme alternative, d'autres matériaux de recouvrement cette fois biodégradables ont été obtenus en réticulant la γ-cyclodextrine (CD) greffée avec des chaines de PEG avec l'acide citrique. En effet, l'acide citrique peut également se coordonner efficacement aux sites de fer disponibles situés au niveau de la surface des nanoMOFs. En outre, les cavités des CDs offrent des avantages supplémentaires, tels que l'incorporation de molécules actives à caractère hydrophobe qui forment des complexes d'inclusion.

De manière avantageuse, les copolymères CD-citrate peuvent également être fonctionnalisés avec des fluorophores tels que la cyanine-5, permettant ainsi de d'appréhender le mécanisme d'enrobage des nanoMOFs ainsi que leur stabilité dans de milieux variés. Une molécule anticancéreuse d'intérêt, la doxorubicine (DOX), a été incorporée avec succès dans les nanoMOFs, atteignant des charges allant jusqu'à 64±2 wt%. D'après les études de résonance magnétique nucléaire à l'état solide (ssNMR), la DOX se trouvait à la fois dans les coeurs et dans les coquilles des nanoMOFs. Fait remarquable, la DOX incorporée n'a pas été libérée à un pH physiologique de 7,4, alors qu'elle l'a été immédiatement dans un milieu mimant le milieu intracellulaire lysosomial.

Les études par microscopie confocale ont apporté des résultats intéressants et prometteurs, montrant l'influence de la charge en DOX sur son mécanisme de libération. A forte charge en DOX, lorsque les nanoMOFs ont été mis en contact avec des cellules cancéreuses (lignée cellulaire Hela), la DOX n'a pas été libérée dans le milieu extracellulaire. A des charges optimales en DOX, une libération a été possible avec une diffusion de la DOX au niveau du noyau cellulaire. En parallèle les nanoMOFs coeur-couronne ont été dégradés au sein des cellules et un signal fluorescent correspondant au recouvrement a été détecté dans le cytosol. Il est intéressant de noter que les nanoMOFs pénètrent à l'intérieur des cellules avec leur revêtement, puis se dégradent dans le milieu intracellulaire, libérant à la fois leur revêtement et la substance active, ce qui a permis d'éradiquer les cellules cancéreuses.

Les matériaux mis au point dans ce travail sont très versatiles, leur composition peut être modulée de part les paramètres mis en jeu lors de leur synthèse. Ils peuvent être greffés de manière simple et rapide sur les nanoMOFs dans l'eau, par incubation. Notre stratégie repose sur l'établissement de multiples interactions coopératives entre les copolymères et la surface des nanoMOFs, assurant une bonne stabilité même dans de milieux biologiques complexes ? Cette étude ouvre la voie à l'ingénierie de surface de particules hautement poreuses pour des applications biomédicales. frameworks (nanoMOFs) in diagnosis and treatment of severe diseases such as cancer and infections is of increasing interest. Nonetheless, immediately after their administration, nanoMOFs strongly interact with their physiological environment, including plasma proteins and cells. As a result, they are rapidly cleared from the blood circulation and cannot reach their targets. This premature elimination significantly limits the nanoMOF use in clinical practice. To address the challenges related to the nanoMOF sequestration by the immune system, we developed two strategies to cover the nanoMOFs with "stealth" coatings. To do so, poly(ethylene glycol) (PEG) chains were covalently grafted to polymers which were adsorbed onto the nanoMOFs. Compared to dense non porous nanoparticles, the surface modification of porous nanoMOFs is much more challenging. Indeed, it was disclosed that PEG chains can penetrate into the pores, displacing the loaded drug. Therefore, in this study, we developed different strategies to modify the surface of nanoMOF particles with multi-functional coatings without disturbing their drug loading capacity. The first strategy was to synthesize comb-like co-polymers made of of dextran (DEX), PEG and alendronate (ALN). DEX acts as bulky macromolecule preventing the PEG chains to penetrate inside the nanoMOFs, whereas ALN moieties are the iron complexing groups to firmly anchor the coatings to the nanoMOFs. The DEX-ALN-PEG copolymers were obtained by click chemistry. They were spontaneously coated onto the nanoMOFs by incubation in water. The resulting stable shells successfully decreased the nanoMOFs' uptake by macrophages in vitro. In an alternative approach, biodegradable PEGbased coatings were engineered by cross-linking PEG-grafted γ-cyclodextrin (CD) with citric acid, which could also efficiently coordinate with nanoMOF surface iron sites. In addition, the CD cavities offer additional benefits such as drug incorporation inside the shell. The copolymers could be functionalized also with fluorophores such as Cyanines-5, allowing deciphering the coating mechanism. Doxorubicin (DOX) was successfully incorporated into the core-shell nanoMOFs reaching loadings up to 64±2 wt%. According to solid state nuclear magnetic resonance (ssNMR) investigations, DOX located both in the cores and in the shells. Remarkably, the incorporated DOX didn't release out at physiological pH 7.4, whereas it released out immediately in artificial lysosome fluid (ALF). Confocal microscopy investigations were in good agreement with these findings: when nanoMOFs were put in contact with cancer cells (Hela cell line), DOX was not released in the extracellular medium, but only inside the cells. Noteworthy, the engineered nanoMOFs penetrate inside the cells together with their coatings, and then degrade intracellularly, releasing both their coatings and the active dug cargo, which in turn eradicated the cancer cells.

This study pave the way towards engineered the surface of highly porous particles for biomedical applications. L'utilisation de nanomatériaux hybrides organiques-inorganiques (nanoMOFs) pour le diagnostic et le traitement de maladies graves, telles que le cancer et les infections, suscite un intérêt croissant. Cependant, immédiatement après leur administration, les nanoMOFs interagissent fortement avec leur environnement physiologique, notamment avec les protéines plasmatiques. Par conséquent, ils sont rapidement éliminés de la circulation sanguine et ne peuvent atteindre leurs cibles. Cette élimination prématurée limite considérablement l'utilisation des nanoMOFs en pratique clinique. Pour relever les défis liés à la séquestration des nanoMOFs par le système immunitaire, nous avons développé deux stratégies pour recouvrir les nanoMOFs de revêtements "furtifs". Pour ce faire, des chaînes de poly(éthylène glycol) (PEG) ont été greffées de manière covalente à des polymères qui ont été adsorbés sur les nanoMOFs. Par rapport aux nanoparticules denses nonporeuses, la modification de surface des nanoMOFs poreux est beaucoup plus difficile. En effet, il a été révélé que les chaînes de PEG peuvent pénétrer dans les pores, chassant le médicament encapsulé. Par conséquent, dans cette étude, nous avons développé différentes stratégies pour modifier la surface des particules de nanoMOFs avec des revêtements multifonctionnels sans perturber leur capacité d'encapsulation de médicaments. La première stratégie a consisté à synthétiser des copolymères de type « peigne » constitués de dextran (DEX), de PEG et d'alendronate (ALN). Le DEX agit comme une macromolécule volumineuse empêchant les chaînes de PEG de pénétrer à l'intérieur des nanoMOFs, tandis que les fragments d'ALN sont des groupes complexant le fer pour ancrer fermement les revêtements aux nanoMOFs. Les copolymères DEX-ALN-PEG ont été obtenus par chimie click. Ils ont été spontanément adsorbés sur les nanoMOFs par incubation dans l'eau.

Comme alternative, d'autres matériaux de recouvrement cette fois biodégradables ont été obtenus en réticulant la γ-cyclodextrine (CD) greffée au PEG avec de l'acide citrique, qui peut également se coordonner efficacement avec les sites de fer disponibles de la surface des nanoMOFs. En outre, les cavités CD offrent des avantages supplémentaires tels que l'incorporation de médicaments à l'intérieur de la couronne. Les copolymères pourraient également être fonctionnalisés avec des fluorophores tels que les cyanines-5, permettant ainsi de comprendre le mécanisme d'enrobage. La doxorubicine (DOX) a été incorporée avec succès dans les nanoMOFs, atteignant des charges allant jusqu'à 64.3±2.3 wt%. D'après les études de résonance magnétique nucléaire à l'état solide (ssNMR), la DOX se trouvait à la fois dans les coeurs et dans les coquilles. Fait remarquable, la DOX incorporée n'a pas été libérée à un pH physiologique de 7,4, alors qu'elle l'a été immédiatement dans un milieu mimant le lysosome artificiel (ALF). Les études par microscopie confocale ont confirmé ces résultats : lorsque les nanoMOFs ont été mis en contact avec des cellules cancéreuses (lignée cellulaire Hela), la DOX n'a pas été libérée dans le milieu extracellulaire, mais uniquement à l'intérieur des cellules. Il est intéressant de noter que les nanoMOFs pénètrent à l'intérieur des cellules avec leur revêtement, puis se dégradent dans le milieu intracellulaire, libérant à la fois leur revêtement et la substance active, ce qui a permis d'éradiquer les cellules cancéreuses. Cette étude ouvre la voie à l'ingénierie de surface de particules hautement poreuses pour des applications biomédicales.

Acknowledgement

First of all, I am very grateful for the support of my thesis supervisor Dr. Ruxandra Gref for creating the great research environment where I have performed my six months' internship and four years' PhD studies. I am super proud to be one of her PhD student. The joy and enthusiasm she possessed for her great work inspire me during all these years. Moreover, I am quite impressed by her intelligence, immense knowledge, patience, and kindness to me and my family. Not only do I appreciate the time she has taken to guide me in all of my research and the writing of this thesis, I am really thankful to all her contributions of ideas, and connections with the collaborators to make my PhD study productive. Furthermore, I will never forget that five years ago, it is all her great help both mentally and financially that makes the turning point in my life. I left the company in China where I have worked for five years and came back to university (in France) to expand my knowledge. This is a "key" for me to enter into a promising future with more opportunities to pursue my interest, fulfill my dream in science and realize the value of life. I am very happy to have my life established in France with my wife (Xue Li) who is also working with Dr. Ruxandra Gref and my son who is more than 3 years old. This is the life I could not image 5 years ago. Words cannot express all my gratitude, but she is the real angel that God sent to me in my life. I will never forget in all my life. I would like to thank Prof. Jiwen Zhang in Shanghai Institute of Materia Medica, Chinese Academy of Sciences. He is my wife's master supervisor, who not only kindly recommended my wife as a Ph.D candidate, but also supported me to apply for my founding from China Scholarship Council (CSC). Without his valuable support it would be more difficult to start this research.

I appreciate

General Introduction

Among the increasing number of nano-drug delivery systems developed so far (e.g. liposomes, lipids, polymers, metallic nanoparticles, dendrimers, silicon, carbon nanotubes, etc), nanoscale hybrid Metal-Organic Frameworks (nanoMOFs) are presently drawing a growing attention due to their unprecedented porosity, versatile composition and topology, enabling high loadings of a large number of drugs together with their controlled release [1][2][3][4][5] . However, the biological applications of nanoMOFs are still in their infancy: no systems reached the clinic stage and the in vivo investigations of their pharmacological efficacy are still rare. The main reasons are the stability issues and the in vivo fate of nanoMOFs. Indeed, as-synthesized nanoMOFs without surface modification easily aggregate during storage. Moreover, once administered in vivo by the iv route, they can aggregate and are readily recognized and uptook by the immune system.

Uncoated nanoMOFs are prone to be removed from the blood circulation quickly before reaching their targets. Therefore, it is crucial to functionalize the nanoMOFs surface in order to achieve stealth (long-circulating) and targeting properties for biomedical applications.

Since more than 30 years, among all the coating materials, such as silica, lipid, and polymers, it was revealed that surface modification with hydrophilic polymers such as poly (ethylene glycol) (PEG) in a "brush" configuration could remarkably prolong the blood circulation times of NPs by reducing their recognition by the reticuloendothelial system 6,7 . However, so far, only a few cases aiming to functionalize the surface of nanoMOFs with PEG shells have been reported [8][9][10][11][12] Indeed, it is a big challenge to functionalize the surface of nanoMOFs with PEG chains, because they are: 1) biodegradable, requiring green coating methods which will not degrade them or fill their pores with side-products or reactants; and 2) highly porous, so the coating materials should not penetrate into the pores, leading to undesirable drug release. Consequently, the real challenge for the application of nanoMOFs in the field of biomedicine remains not only on achieving high payloads and controlled releases, proper particle size and stability but also on functionalizing the surface of the nanoMOFs, which plays an important role on the in vivo fate of the nanocarriers.

The ''ideal'' nanoMOFs coating should reach the following goals: 1) only anchor onto the surface of nanoMOFs, but not enter into the porous structure; 2) not disturb the loaded drugs;

3) improve the colloidal stability of nanoMOFs during storage and in biological media; 4) have stealth effect, i.e. avoid opsonization to enable a longer circulation time and 5) target to the disease lesions. These topics are presented in Chapter 1, together with a detailed review of the literature [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] .

To achieve these challenging objectives, two main strategies (Schema 1) have been developed

to functionalize the surface of nanoMOFs.

Schema 1: The synthesis of DEX-ALN-PEG (A) and CD-citrate-PEG oligomers (B)

The first strategy was synthesis of co-polymers by linking multifunctional chains in the backbone. In this study, biocompatible dextran was chosen as the backbone, and was grafted with alendronate (ALN) which can spontaneously coordinate to the nanoMOFs surface and PEG chains that can reduce macrophage uptake.

Dextran-alendronate-PEG copolymers (DEX-ALN-PEG) were successfully synthesized by click chemistry. The amounts of PEG and ALN moieties in this copolymers were controllable by tuning the molar ratio between them. DEX-ALN-PEG copolymers could spontaneously anchor onto the outer surface of MIL-100(Fe) by green method, without any surfactant or organic solvent, just socking nanoMOFs with DEX-ALN-PEG in pure water. The optimized coatings were able to greatly reduce the nanoMOFs adsorption by HSA, which is the most abundant protein in human blood plasma. What is more, the coating could improve the nanoMOFs's colloidal stability and decrease three times the in vitro macrophages uptake (Chapter 2). However, the synthesis of DEX-ALN-PEG was very complicated, involved multistep reactions with low total yield; DEX-ALN-PEG is not biodegradable.

The second strategy was to synthesize biodegradable γ-CD oligomers using citric acid as cross linker, which also plays a key role in anchoring the oligomers onto the external surface of the nanoMOFs due to the coordination between citrate and available surface-located iron sites from the nanoMOFs. More interestingly, γ-CD-citrate-PEG oligomers were successfully synthesized with the same method.

In order to obtain a biodegradable coating materials, a water soluble biocompatible cyclodextrin (CD)-based materials (γ-CD-citrate oligomers) were successfully synthesized by a facile and environmental friendly method. Stable core-shell NPs relied on iron trimesate nanoMOFs cores with γ-CD-citrate oligomers coatings, successfully prepared through a fast and "green" method.

Furthermore, fluorescence lifetime microscopy was used to decode the coating mechanism due to the presence of the fluorophore which was also linked with the shell. Remarkably, doxorubicin (DOX) encapsulations reached 48 ± 10 wt% or 65 ± 8 wt% respectively, after 1 day or 2 days incubation with the drug solution (Chapter 3). Interestingly, CD coating could promote the loading process and accelerate DOX release. In order to investigate the nanoMOFs component localization at a molecular scale and study their physicochemical properties, [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] Clabeled CD-CO oligomers were synthesized using the reported method, and further coated on the external surface of MIL-100 (Al). A set of high-resolution solid state nuclear magnetic resonance (ssNMR) experiments were employed to achieve a deep understanding of the interactions involved in the MIL-100(Al) core-shell structure. CD-citrate-PEG oligomers were successfully synthesized and further engineered on the outer surface of MIL-100 (Fe), enabled to reduce 2 times the in vitro uptake by macrophages. Remarkably, the DOX release from the DOX loaded nanoMOFs was pH responsive.

Finally, Chapter 4 is a comparative discussion of the coatings considering their synthesis routes, biodegradability, stabilities, interactions with cells, influences on drug loading and release.

Then, a general conclusion and the perspectives of the work are presented.

13. Qiu, J., Li, X., Gref, R. & Vargas-berenguel, A. Carbohydrates in metal organic frameworks : Supramolecular assembly and surface modification for biomedical applications. Metal-Organic Frameworks for Biomedical Applications (Elsevier Inc., 2020). doi:10.1016/B978-0-12-816984-1.00022-6

Chapter 1.1: Introduction

Over the years, nanotechnology is introducing more and more in our daily routine. 1,2 . This revolutionary technology has been applied in multiple fields through an integrated approach.

An increasing number of applications and products containing nanomaterials or at least with nano-based claims have become available. This also happens in pharmaceutical research. The use of nanotechnology in the development of new medicines is now part of our research and in the European Union (EU) it has been recognized as a Key Enabling Technology, capable of providing new and innovative medical solution to address unmet medical needs [3][4][5][6] .

The application of nanotechnology for medical purposes has been termed nanomedicine and is defined as the use of nanomaterials for diagnosis, monitoring, control, prevention and treatment of diseases. Core shell nanoparticles (NPs) with engineered surface are designed to target specifically to diseased cells, which allows for the direct treatment of those unhealthy cells.

Therefore, they are expected to improve efficacy, decrease side effects, and overall improve human health by acting as drug carrier 3,7,8 . Various nanostructures, including liposomes, SLN (solid lipid nanoparticles), polymers, dendrimers, silicon, carbon materials, metallic nanoparticles, and more recently metal organic frameworks (MOFs) have been used as carriers in drug delivery systems [9][10][11][12][START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] .

The first case of reported NP as drug carrier was liposome. Liposomes are mainly composed of phospholipids that form bilayers with an aqueous phase inside. They were considered as typical nanomedicine, frequently evaluated in basic research and clinical medicine. In the recent literature, liposomes have been revealed to enhance the solubility of drugs and ameliorate their pharmacokinetics. For example, they allowed circumventing drawbacks of chemotherapeutic reagents, such as rapid metabolism, they decreased side effects and increase of anticancer activity. There are plenty of drug examples based on liposomal formulations, for instance, anticancer drugs, neurotransmitters (serotonin), antibiotics, anti-inflammatory, and antirheumatic drugs 7,[START_REF] Bulbake | Liposomal Formulations in Clinical Use : An Updated Review development[END_REF] . But the main disadvantages of liposome lies in sometimes low drug loading capacity and lack of stability.

The Polymeric NPs (PNPs) are achieved from synthetic polymers [START_REF] Begines | Polymeric Nanoparticles for Drug Delivery : Recent Developments and Future Prospects[END_REF] , such as polyacrylamide and polyacrylate, or natural polymers, for instance, albumin, DNA, chitosan. PNPs are recognized as biodegradable materials according to the in vivo behavior. Recently, polylactide (PLA) and poly (lactide-co-glycolide) (PLGA) approved by FDA, are most common utilized for the preparation of polymeric biodegradable nano-DDS.

Solid lipid NPs (SLN) are particles stemmed from solid lipids, for example, complex glyceride mixtures, highly purified triglycerides, or waxes stabilized by different kinds of surfactants 7,[START_REF] Montoto | Solid Lipid Nanoparticles for Drug Delivery : Pharmacological and Biopharmaceutical Aspects[END_REF] .

Good colloidal stability, protection of incorporated drugs from degradation, controlled drug release, and good tolerability are the major advantaged properties, which makes that SLN can be widely used for the application in the biomedical field. However, in the recent literature, some disadvantages have been found, such as low loading capacity due to the low solubility of drug in the lipid and the structure and polymorphic state of the lipid matrix, drug bursting out after crystallization.

Dendrimers are unique polymers with well-defined size and structure. They are highly branched macromolecules with controlled near monodisperse three-dimensional architecture emanating from a central core. Dendritic architecture is one of the most common structures detected among all biological systems. Several examples of nanometric molecules with dendritic structure include: glycogen, amylopectin, and proteoglycans. Poly (amido amide) (PAMAM) is an example of synthetic dendrimer material which is widely utilized in biomedical applications [START_REF] Buwalda | Ultrafast in situ forming poly(ethylene glycol)-poly(amido amine) hydrogels with tunable drug release properties via controllable degradation rates[END_REF] .

Not only the synthesis and structure of PAMAM dendrimers, but also the distribution of drugs or genes inside these compounds have been intensively studied.

Carbon nanomaterials employed in DDS can be divided into nanotubes (CNTs) and nanohorns (CNH). CNTs possess particular structure composed of rolling of single or multilayers of graphite with a huge surface area and an outstanding thermal and electronic conductivity [START_REF] Mohajeri | Biomedical applications of carbon nanomaterials : Drug and gene delivery potentials[END_REF][START_REF] Zhao | Mesoporous carbon nanomaterials in drug delivery and biomedical application[END_REF] .

Biocompatibility of CNTs may be ameliorated by chemical modification of their surface with macromolecules, such as employing covalent anchoring of PAMAM dendrimers, or amphiphilic diblock copolymers. Nanohorns which is a single-wall nanotubes-demonstrate similar characters to nanotubes. They can be easily prepared with very cheap materials and can be produced in good quality [START_REF] Zhao | Mesoporous carbon nanomaterials in drug delivery and biomedical application[END_REF] .

Silica materials involved in controlled drug delivery systems are composed of xerogels and mesoporous silica NPs (MSNs). They revealed several advantages as drug carrier systems, containing biocompatibility, an ease in terms of functionalization, and highly porous framework 10 . Silica materials are extensively employed as drug carriers for biological purposes, and are considered of main interest among inorganic NPs.

Metallic nanocarriers are a material of choice, because of their special characteristics, such as controllable size, large surface area, facile synthesis, interesting optical properties, etc 3,8 .

Moreover, metallic NPs have intrinsic properties useful in nanomedicine. For example, iron oxides NPs are usually synthesized as alkaline co-precipitations of Fe 2+ and Fe 3+ salts in H2O in the presence of a hydrophilic polymer such as dextran (DEX) or poly(ethylene glycol) (PEG).

This yields an iron core of 4-5 nm in diameter, which can be used as contrast agents for magnetic resonance imaging. Gold NPs (GNPs) have an unique photodynamic property through absorption of near-infrared lights and emission of light and heat, which has been employed for cancer photothermal therapy. What's more, GNPs are conjugated with plenty of therapeutic agents and targeting moieties, and act as efficient drug carriers.

NanoMOFs

Though there are already so many NPs which can be used as nano DDS with their own advantages, they are also confronted with some limitations, such as, depending on the system, instability, rapid biodegradation, side effects incapacity to load drug combinations, and toxicities. To overcome these limitations, the development of new engineered carriers is essential. It is indeed important to ensure a prolonged and controlled drug release, reducing the administered doses and frequencies to decrease the side effects and improve patient compliance to the treatment.

In this context, a new class of porous hybrid materials, named as Metal-Organic Frameworks (MOFs), has been recently discovered and since has intensively used as a drug carrier to overcome the above mentioned challenges. MOFs are organic-inorganic hybrid porous networks based on organic polydentate ligands and inorganic units. Since their discovery in 1989 [START_REF] Robson | Infinite Polymeric Frameworks Consisting of Three Dimensionally Linked Rod-like Segments[END_REF] , there has been increasingly more attention paid on MOFs because of their fruitful applications in gas storage, separation, catalysis, sensing, energy, optics, and recently, in drug delivery system [9][10][11][12][START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF][START_REF] Dhakshinamoorthy | Catalysis and photocatalysis by metal organic frameworks[END_REF][START_REF] Yang | Recent advances in nanosized metal organic frameworks for drug delivery and tumor therapy[END_REF][START_REF] Ahmed | Nano-MOFs as targeted drug delivery agents to combat antibioticresistant bacterial infections[END_REF] . One of the crucial advantages of the MOFs is their easily tunable composition. The composition and structure of these porous hybrid materials can easily be adjusted through the almost unlimited possible combination of metal inorganic subunits and/or constitutive organic ligands, contributing to thousands of MOFs with special functions [START_REF] Horcajada | Porous metal-organic-framework nanoscale carriers as a potential platform for drug delivery and imaging[END_REF] . The first case of reported use of MOFs as DDS [START_REF] Horcajada | Metal-organic frameworks as efficient materials for drug delivery[END_REF] and as contrast agents [START_REF] Rieter | Nanoscale metal-organic frameworks as potential multimodal contrast enhancing agents[END_REF] Nanosized MOFs (nanoMOFs), due to their small particle size, can indeed ameliorate the drug delivery properties to treat many diseases 9,[START_REF] Ahmed | Nano-MOFs as targeted drug delivery agents to combat antibioticresistant bacterial infections[END_REF] . Furthermore, MOFs possess lots of key characteristics that indeed confer them outstanding attributes as drug carriers for the controlled release of drugs [START_REF] Li | Drug-Loaded Lipid-Coated Hybrid Organic-Inorganic " Stealth " Nanoparticles for Cancer Therapy[END_REF] . Indeed, engineered nanoMOFs possess the following features: 1) high drug adsorption capacities (payload and efficiency), 2) controlled drug release getting rid of "burst" effect (rapid release of the selected drug in very short time), 3) ability to target abnormal cells and organs in a selective way, 4) low toxicity by gradually degrading and elimination of byproducts without accumulation in the body, 5) a facile modification of the outer surface of the NPs for an enhanced in vivo stability and/or targeting, and 6) possibility of incorporating biofriendly contrast agents [START_REF] He | Metal-organic frameworks for advanced drug delivery[END_REF][START_REF] Qiu | Carbohydrates in metal organic frameworks : Supramolecular assembly and surface modification for biomedical applications[END_REF] .

In the past few years, various research works have revealed the great potential of MOFs in the biomedical field as DDS, biosensors, biocatalysts, contrast agents, decontamination agents and photothermal therapy (PTT), and so on 9,12,[START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] . Several of them were published until recently based primarily on their bioactive composition, treatment of severe diseases 11 . Noticeable, nanoMOFs approaching a preclinical in vivo investigation are those relying on Fe carboxylates [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF][START_REF] Baati | In depth analysis of the in vivo toxicity of nanoparticles of porous iron(III) metal-organic frameworks[END_REF] .

Surface Modification of nanoMOFs

Up to now, there is a huge gap between the large number of studies reporting the potential of MOFs to be used for biomedical applications and contrasts with the few number of in vivo pharmacological efficacy investigations studied to date 11 . There are several reasons which can explain this situation.

Firstly, colloidal stability issues and/or lack of organ or tissue targeting properties. It is well known that intravenous (i.v.) injection of NPs needs dispersion in an isotonic solution (for instance 0.9% saline or 5% glucose) for isotonicity. Various nanoMOFs, especially once dried, cannot be re-dispensed yet in these media and have the tendency to aggregate rapidly, making administration impossible due to risk of embolization. Colloidal stability of nanoMOFs is still a main problem that becomes a big challenge for in vivo administration 10 . Selective modification of the outer surface of nanoMOFs is thus crucial in order to overcome such challenges.

Secondly, short blood circulation times (opsonization after injection of uncoated nanoMOFs).

Very interestingly, surface modification of NPs with hydrophilic polymers can greatly decrease the opsonization. Surface functionalization strategies of NPs and liposomes have been investigated during recent decades. The first case of PEGylated liposome was reported by Woodle and colleagues in 1992 [START_REF] Woodle | Versatility in lipid compositions showing prolonged circulation with sterically stabilized liposomes[END_REF] . PEGylation of polymeric NPs was first reported in 1994 by Dr. Gref and co-workers [START_REF] Gref | Biodegradable long-circulating polymeric nanospheres[END_REF][START_REF] Gref | The controlled intravenous delivery of drugs using PEG-coated sterically stabilized nanospheres[END_REF] , since then, an increasing number of studies focusing on PEGylation of other types of NPs have been reported [START_REF] Giménez-Marqués | GraftFast surface engineering to improve MOF nanoparticles furtiveness[END_REF][START_REF] Edward Semple | Synthesis and facile endgroup quantification of functionalized PEG azides[END_REF][START_REF] Zhang | Effects of PEGylated paclitaxel nanocrystals on breast cancer and its lung metastasis[END_REF] . NPs modified with PEG extends the circulation time in vivo, enabling their distribution in other organs. However, up to now, Doxil represents the only FDA-approved PEGylated liposomes for the treatment of cancer using doxorubicin (DOX) [START_REF] Barenholz | Doxil® -The first FDA-approved nano-drug: lessons learned[END_REF] .

Generally, surface functionalization of NPs should meet some standard to be fit for drug delivery, such as a) bio-friendly synthesis method, without using any toxic additive, b) maintain drug release ability, c) not interfere with the loaded drugs, d) not penetrate into the porous architecture, e) improve nanoMOFs physical stability, and f) excellent stability under physiological conditions.

NanoMOFs can be functionalized with silica, lipid layer, carbohydrate, polymer, etc. In the following section, three types of coating materials containing silica coating, lipid layer encapsulation, and carbohydrate coating are considered.

Surface modification of nanoMOFs with thin silica shells can improve the stability of particles, elevate biocompatibility and water dispersibility, and enable functionalization with numerous silyl-derived compounds. In general, the outer surface of NanoMOFs are first modified with a polymer, such as polyvinylpyrrolidone (PVP) which can boost the silica coating process.

Furthermore, tetraethylorthosilicate (TEOS) was used to treat the PVP-coated particles in basic ethanol to provide silica coatings. The silica shell thickness can be manipulated by tuning the reagent concentrations or/and reaction time. Release profile experiments revealed that a silica shell can remarkably slow down the rate of particle decomposition. Moreover, the outer surface of these core-shell architecture can be further functionalized through anchoring a silyl-derived molecule to the silica shell with surface silanol groups to provide important performance, such as optical contrast agents, sensory agents, and targeting ligands 10 . However, surface coatings were sometimes uneven and/or they were unstable. Moreover, no study so far dealt with PEG surface coatings.

Lipid have recently been used as coating materials of choice to modify the surfaces of nanoMOFs. Lipid bilayers not only slow the pharmacokinetic but also improve biocompatibility [START_REF] Sicklinger | their uptake by cancer cells †MOF nanoparticles coated by lipid bilayers and their uptake by cancer cells[END_REF] . The flexible nature of lipid bilayer is an advantage over the rigid silica coating. Wuttke and his colleagues reported the lipid DOPC (1,2-dioleoyl-snglycero-3phosphocholine) coated nanoMOFs as a potential and new type of nanocarriers [START_REF] Sicklinger | their uptake by cancer cells †MOF nanoparticles coated by lipid bilayers and their uptake by cancer cells[END_REF] . They revealed that this novel drug delivery system can efficiently load dye molecules inside the porous structure of the MOF while the lipid bilayer avoids their burst release. Effective uptake of the lipid coated nanoMOFs by cancer cells enables these nanocarriers promising for drug delivery and diagnostic. Based on this study, we further functionalized here nanoMOFs with PEGylated lipid, namely DSPE-PEG2000 [START_REF] Li | Drug-Loaded Lipid-Coated Hybrid Organic-Inorganic " Stealth " Nanoparticles for Cancer Therapy[END_REF] .

We found that the outer surface of nanoMOFs functionalization with lipids provides a better control over their degradation in biological media and drug release. Moreover, we found that the coatings were able not only to decrease macrophage uptake in vitro but also to eradicate cancer cells by carrying the drug inside.

Carbohydrates, one of the four major classes of biomolecules have been well-recognized as edible materials. The excellent properties of hydrophilicity and versatility play an important role for the carbohydrates to be used as surface modification materials for preformed drug nanocarriers. They present the advantages of: (i) low toxicity; (ii) could mediate the cell internalization; and (iii) can be easily further engineered with other functional moieties.

Recently, carbohydrate-based materials, for instance, CDs 39 , heparin [START_REF] Bellido | Heparin-Engineered Mesoporous Iron Metal-Organic Framework Nanoparticles: Toward Stealth Drug Nanocarriers[END_REF] , hyaluronic acid [START_REF] Giménez-Marqués | GraftFast surface engineering to improve MOF nanoparticles furtiveness[END_REF] , and chitosan [START_REF] Hidalgo | Chitosan-coated mesoporous MIL-100(Fe) nanoparticles as improved bio-compatible oral nanocarriers[END_REF] , which possess bigger molecular size than the pore size of nanoMOFs, have been frequently investigated as coating materials.

The first investigation of nanoMOF surface modification with carbohydrate coating material was reported by the group of Dr. Gref [START_REF] Noiray | A "green" strategy to construct non-covalent, stable and bioactive coatings on porous MOF nanoparticles[END_REF] , where phosphorylated CD (CD-P) derivatives were employed to functionalize iron trimesate nanoMOFs MIL-100(Fe) in aqueous media, adopting a facile, one-step, and completely "green" (without any organic solvent or surfactant) method.

Besides, CD-P were too bulky to penetrate into the porous of the nanoMOFs, leaving CD-P anchored only on the outer surface of nanoMOFs. In addition, the CD-P functionalization enables the possibility of PEGylation by formation of guest-host inclusion with adamantane-PEG chains [START_REF] Noiray | A "green" strategy to construct non-covalent, stable and bioactive coatings on porous MOF nanoparticles[END_REF] . Giménez-Marqués and her co-workers employed GraftFast surface engineering procedure to modify nanoMOF, using PEG and hyaluronic acid (HA) acrylates [START_REF] Giménez-Marqués | GraftFast surface engineering to improve MOF nanoparticles furtiveness[END_REF] . HA-PEG was covalently grafted by a single step GraftFast reaction on the surface of nanoMOFs. The resulting PEGylated nanoMOFs demonstrated improved colloidal stabilities in different kinds of biological media as compared with uncoated nanoMOFs, while preserving their porosity which enabled the adsorption of bioactive molecules. The PEG coating could decrease macrophage uptake in vitro. However, the in situ synthesized HA-PEG is not biodegradable and the amount of PEG chains could not be well controlled by GraftFast reaction.

More recently, the water-soluble copolymer poly(3-hydroxyalkanoate) grafted with PEG chains were synthesized 42 to functionalize nanoMOFs. The resulting PEGylated nanoMOFs displayed improved colloidal stabilities in aqueous and physiological fluids and no cytotoxicity (up to 100 µg/mL) on murine macrophages J774.A1 cells. However, the stealth property still needs to be further investigated.

In this context, we focused in this thesis on the design and synthesis of novel biocompatible carbohydrate polymers with: 1) controlled PEG contents and 2) biodegradability. To reach the first goal, we grafted PEG chains to dextran backbones at different ratios [START_REF] Cutrone | Comb-like dextran copolymers: A versatile strategy to coat highly porous MOF nanoparticles with a PEG shell[END_REF] . As the second strategy, we applied polymerized PEGylated CDs, creating biodegradable ester bonds [START_REF] Cutrone | Design of engineered cyclodextrin derivatives for spontaneous coating of highly porous metal-organic framework nanoparticles in aqueous media[END_REF] . After the design and successful synthesis, we prepared core-shell NPs as drug delivery system using the synthesized coating materials as shell and the porous nanoMOFs as the core [START_REF] Cutrone | Comb-like dextran copolymers: A versatile strategy to coat highly porous MOF nanoparticles with a PEG shell[END_REF][START_REF] Cutrone | Design of engineered cyclodextrin derivatives for spontaneous coating of highly porous metal-organic framework nanoparticles in aqueous media[END_REF][START_REF] Qiu | Self-assembled multifunctional coreshell highly porous metalorganic framework nanoparticles[END_REF][START_REF] Li | Chapter 3.3 Porous nanoparticles with engineered shells release their drug cargo in cancer cells In preparation Aldrich, Saint-Quentin-Fallavier, France, 95%) and absolute ethanol (EtOH, Carlo Erba, Valde-Reuil, France, 99%) were used for the synthesis of MIL-100 (Fe). Water was purified by a Millipore MilliQ system. Doxorubicin hydrochloride (DOX, Sigma-Aldrich, 98%), and Deferoxamine mesylate salt (DFOM, Sigma-Aldrich, 92.5%), were used for drug encapsulation and drug release studies. latrunculin A (LA, Sigma-Aldrich, ≥85%) and cytochalasin D (CyD, Sigma-Aldrich, 98%) were used for inhibitors. Cell culture. Murine macrophage cell line J774.1 (ATCC® TIB-67™) was grown in Dulbecco's Modified Eagle's Medium (DMEM, ThermoFisher Scientific, Villebon-sur-Yvette, France) supplemented with 10% v/v decomplemented fetal bovine serum (FBS)[END_REF] . A set of complementary characterizing experiments were performed to characterize the coated nanoMOFs. Drug of interest was loaded into the coated nanoMOFs, in order to study the drug loading and release. The optimized formulations were finally evaluated in vitro.

General Objectives and Author Contributions

The first case of nanoMOFs was reported by the scientists in 1989. Since their discovery, nanoMOFs have attracted a large number of interest because of their excellent properties.

However, it is absolutely necessary to functionalize the outer surface of nanoMOFs with coating materials such as PEG chain to prevent them from the recognition and removal by the innate immune system, an indispensable condition for biomedical applications. Until recently, this was still a big challenge because of the highly porous composition and degradability of nanoMOFs. In this context, we achieved the surface modification of the external surface of MIL-100(Fe) nanoMOFs by a green and environmental friendly method, utilizing a new family of comb-like copolymers made of dextran (DEX) grafted with both PEG and alendronate (ALN) moieties. DEX-ALN-PEG copolymers were synthesized with click chemistry, controlling both the amount of PEG and ALN moieties by adjusting the molar ratio between them. A stable coating was formed while DEX-ALN-PEG copolymers spontaneously anchored onto the outer surface of MIL-100(Fe).

Complementary methods were used to prove the efficacy of DEX-ALN-PEG copolymers on the performance of stealth effect.

1) The interaction of coated or uncoated nanoMOFs with HSA, the most abundant protein in human blood plasma, was investigated in order to gain further understanding on protein adsorption. There results clearly revealed that the optimized coatings are able to decrease five times the adsorption of this modal protein.

2) Optical microscopy was used to observe the cells after incubation with the nanoMOFs.

After 2 h incubation, lots of nanoMOFs were took up by macrophage. On the contrary, with the same incubation time, the uptake of coated nanoMOFs was dramatically reduced.

3) Inductively coupled plasma mass spectrometry (ICPMS) further quantified the amount of internalized uncoated and coated nanoMOFs. The optimized coating reduced three times the in vitro macrophages uptake, affording new perspectives for the applications in biomedical field.

All the authors have contributed in this work. I performed the experiments on the nanoMOFs synthesis, characterization, surface modification, colloidal stability and cell interaction, and also contributed to the analysis of the data and writing of the main manuscript.

a These authors contributed equally.

Supplementary Material

Table of Contents

I. 1 H NMR spectra for compounds 3-8

S2-S7

II. [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C NMR spectra for compounds 3-5

S8-S10

III. [START_REF] Woodle | Versatility in lipid compositions showing prolonged circulation with sterically stabilized liposomes[END_REF] P NMR spectra for compounds 3 and 6-8

S11-S14

IV.

Compositional model for the structural analysis of dextran derivatives 5-8

S15-S20

V. In order to analyze the structure of dextran derivatives 5-8 from compositional data obtained by elemental (for C, H and N) and HR-ICP-MS (for P, Na, Cu and Li) analyses, the following mathematical model is proposed. Although dextran presents a structure consisting of glucose moieties linked through α(1→6) glycosidic bonds with some branches mainly from α(1→3) and α(1→4) linkages, in order to simplify the model we will consider a linear, nonbranched, main chain containing k glucose residues:

SEC chromatograms S21

Dextran T-40 (DT40) with 40000 Da of average molecular weight was used. For that reason, the average number of glucose units for each DT40 chain (degree of polymerization) is calculated as: When DT40 is propargylated by using CDI and propargylamine, a number of propargyl carbamate group are formed according to this reaction: Thus, each propargyl group introduced in the polymer modifies the molecular formula of DT40 by adding C4H3ON.

In addition, the final product has to be considered hydrated by an undefined number of H2O molecules (m), which obviously are reflected in the compositional analysis of the product. As a result, the molecular formula for DEX-PC

N atoms n

In this analysis we assume that all nitrogen atoms present in the molecule are part of propargyl carbamate groups, and thus no imidazole carbamate groups remain on the structure. This hypothesis is based firstly on the large excess of propargylamine that was used during the reaction, and secondly on the subsequent dialysis process that should break all unreacted imidazole carbamates to the original alcohols. DEX-PC 5 is subsequently reacted with alendronate azide derivative 3 and poly(ethylene glycol) azide 4 in the presence of LiCl and CuSO4/sodium ascorbate catalyst system in order to promote CuAAC onto the propargyl carbamate groups. Those dextran OH groups that support a propargyl carbamate group which reacted with alendronate azide 3 will be then transformed into: 1 where na will be the number of sodium atoms, cu the number of copper atoms, and li the number of lithium atoms per alendronate branch. We will consider x as the number of alendronate groups per DEX molecule.

In a similar fashion, those dextran OH groups that support a propargyl carbamate group which reacted with poly(ethylene glycol) azide 4 will be then transformed into: 1 where p is the number of CH2CH2O moieties (degree of polymerization) of poly(ethylene glycol). Considering that we used poly(ethylene glycol) of an averaged molecular weight of 2000 Da, p can be calculated as follow:

We will consider z as the number of poly(ethylene glycol) groups per DEX molecule.

In summary, functionalized dextran with alendronate and poly(ethylene glycol) branches will contain x alendronate groups, z PEG groups, n non-reacted propargyl groups 1 and m hydration molecules of H2O. In addition, alendronate groups might contain na atoms of Na, cu atoms of copper and li atoms of Li. Thus, molecular formula of such derivative will be:

Li atoms li•x

On the basis of this analysis, weight percentage of each element can be defined as follows:

%𝐶 = 100𝑐 1 (6𝑘 + 4𝑛 + 14𝑥 + (2𝑝 + 5)𝑧) 𝑑 %𝐻 = 100ℎ 1 (10𝑘 + 2 + 3𝑛 + 2𝑚 + (25 -𝑛𝑎 -𝑐𝑢 -𝑙𝑖)𝑥 + (4𝑝 + 6)𝑧) 𝑑 %𝑂 = 100𝑜 1 (5𝑘 + 1 + 𝑛 + 𝑚 + 9𝑥 + (𝑝 + 1)𝑧) 𝑑 %𝑁 = 100𝑛 1 (𝑛 + 5𝑥 + 4𝑧) 𝑑 %𝑃 = 100𝑝 1 (2𝑥) 𝑑 %𝑁𝑎 = 100𝑛𝑎 1 (𝑛𝑎 • 𝑥) 𝑑 %𝐶𝑢 = 100𝑐𝑢 1 (𝑐𝑢 • 𝑥) 𝑑 %𝐿𝑖 = 100𝑙𝑖 1 (𝑙𝑖 • 𝑥) 𝑑
where: From elemental and HR-ICP-MS analyses, the following data were measured: Chapter 3 multifunctional core-shell drug loaded nanoparticles for biomedical applications Chapter 3.1 Self-assembled multifunctional core-shell highly porous metal-organic framework nanoparticles

d = c1[6k + 4n + 14x + (2p + 5)z] + h1[10k + 2 + 3n + 2m + (25 -na -cu -li)x + (4p + 6)z] + o1[5k + 1 + n + m + 9x + (p + 1)z] + n1(n + 5x + 4z) + p1(2x) + na1(na•x) + cu1(cu•x) + li1(li•x) 𝑘 = 40000-

General Objectives and Author Contributions

Due to the special properties originating from the combination of core and shell materials, Coreshell nanoparticles (NPs) have been gaining increasing interest in the field of nanomedicine. It is well known that nanoMOFs are able to entrap a large number of drugs with high payloads.

However, as previously described, it is absolutely necessary to functionalize the nanoMOFs with engineered coatings to improve the colloidal stability, enable a controllable drug release and extend the in vivo circulation time. A novel water soluble biodegradable cyclodextrin (CD)based shell grafted or not with fluorophore were successfully synthesized in order to fulfill such challenging objective. Compared with monomeric CDs, the oligomeric CD coatings crosslinking with citric acid (γ-CD-citrate oligomers) provide better possibilities to graft functional moieties such as fluorescent molecules, spontaneously and effectively anchored on the surface of porous nanoMOFs.

We succeeded to prepare the stable core-shell NPs based on iron trimesate nanoMOFs (~200 nm) with γ-CD-citrate oligomer through a fast and "green" (without any organic solvent and surfactant) method achieved by simply mixing the two components in water. The amounts of γ-CD-citrate oligomers coated on the nanoMOFs reached up to 53 ± 8 wt%, the yield was as high as 86% in the optimized condition. In addition, fluorescence lifetime microscopy can decode the coating mechanism because of the presence of the fluorophore grafted to the shell.

DOX loadings reached 48 ± 10 wt% or 65 ± 8 wt% after 1 day or 2 days incubation with the drug solution, respectively. This study paves the way toward the design of multifunctional coreshell nanoMOFs of interest for applications in the biomedical field.

General Objectives and Author Contributions

Aluminium-based Metal-organic frameworks (Aluminium-based MOFs) were among the first synthesized and studied MOFs based on their stability and highly porous. They have been drawing an increasing interest because they can be synthesized with "green" hydrothermal method. As previously described, more recently, nanoMOFs engineered with versatile coatings are attracting numerous attention in the field of nanomedicine.

In this context, CD-based polymer (CD-CO) coated nanoMIL-100(Al) particles loaded with DOX molecules were successfully prepared and characterized. A high amount of drug was encapsulated in the nanoMOFss due to the high affinity of DOX with both the CD-CO coatings and the MOF cores. More interestingly, it was discovered that the CD-CO coatings favored the release of DOX from the cores, particularly at early stage (< 6h) and for low DOX payloads (especially TDL of 10%). Moreover, the CD-based coating accelerated the DOX loading process.

A series of high-resolution solid state nuclear magnetic resonance (ssNMR) experiments ( 1H-1H and 13C-27Al) were employed to obtain a deep understanding on the multiple interactions involved in the MIL-100(Al) core-shell system. To do so, 13C-labeled CD-CO oligomers were synthesized and further coated on the outer surface of the nanoMOFs. The ssNMR results clearly confirmed the high affinity of DOX with both the CD-CO shells and the nanoMOFs. The ssNMR methodology associated to the selective isotope labeling strategy demonstrated to be effective for the characterization of this system, relying on the diamagnetic character of the nanoMOFs. This study provided the grounds of a methodology to evaluate the nanoMOFs component localization at a molecular scale and to study the physicochemical properties of the nanoMOFs which play a main role on their biomedical applications.

In this study I contributed mainly with the synthesis of all the coating materials and nanoMOFs, surface modification of the nanoMOFs with coating materials, performance of DOX loading and release, and so on.

The ssNMR studies were carried on in collaboration with Dr. Charlotte Martineau-Corcos.

General Objectives and Author Contributions

Due to their huge surface areas and high pore volumes, MIL-100 (Fe) was demonstrated to incorporate a large number of drugs, for instance antibiotics and anticancer drugs, which are able to effectively penetrate within the porous MOF structures. Without surface modification, the uncoated nanoMOFs can be easily recognized and taken up by macrophages and removed from the blood circulation. Finally they will never reach the target sites. Therefore, it is crucial to engineer the external surface of nanoMOFs with 'stealth' properties for biomedical applications. Moreover, there are many biological barriers before achieving an efficient drug delivery. These challenges are related to the capacity to target special cells and tissues, and deliver the cargo to particular intracellular zones. Remarkably, in this study, after incorporating DOX in nanoMOFs, it is shown that DOX didn't release at pH 7.4 in both phosphate buffer saline (PBS) and cell culture medium with high payload, but released very fast and totally in artificial lysosomal fluid (ALF) at pH 4.5.

Briefly, this study included four parts:

1) Biodegradable γ-CD-citrate-PEG oligomers were successfully prepared by grafting CD-PEG monomers with citric acid by a simple one-pot reaction. The external surface of the nanoMOFs was modified with γ-CD-citrate-PEG oligomers by a green method (without any organic solvent or surfactant).

2) The anticancer drug DOX was successfully impregnated into the coated nanoMOFs with high payload up to 50wt%, the drug encapsulation was not affected by the surface modification.

Moreover, DOX released very fast in artificial lysosomal fluid (ALF), compared with a slow release in PBS or cell culture medium.

3) The anticancer effect of the nanoMOFs loaded with anticancer drug and oligomers-coated was studied by MTT test and visualized by confocal investigations.

4) The PEG "brush" allowed to efficiently escape macrophage uptake, as showed by quantitative determination of the amount of internalized nanoMOFs utilizing inductively coupled plasma mass spectrometry (ICP-MS).

Introduction

Metal-organic frameworks (MOFs), first reported in 1989 (Hoskins and Robson, 1989) Since long, it was revealed that surface modification with hydrophilic polymers such as poly (ethylene glycol) (PEG) in a "brush" configuration could significantly prolong the NPs' blood circulation times by reducing their recognition by the reticuloendothelial system (Gref et al., 1995, 1994). Numerous PEGylated drugs are currently used in the treatment of cancer, hepatitis and hemophilia ("FDA approved PEGylated drugs 2021," 2021). However, there are still scarce examples of PEGylated NPs approved by the Food and Drug Administration (FDA): the first PEGylated liposome Doxil® was approved in 1995 (Barenholz, 2012), and the micellar formulation Genexol-PM made using PEG-based block copolymers was approved in 2007 in Korea and marketed in Europe (Pillai, 2014).

However, engineering the surface of porous MOFs with PEG shells remains very challenging as compared to liposomes or dense NPs made of biodegradable polymers. Indeed, the PEG chains can readily penetrate into the nanoMOF pores, releasing the encapsulated drugs in an uncontrolled manner (Agostoni et al., 2015). Recently, in order to avoid PEG penetration into the highly porous MOFs, acryl PEGs were polymerized onto the surface by a GraftFast method (Giménez-Marqués et al., 2018). However, this polymerization method does not allow a precise control of the molecular weight of the resulting PEG-based copolymers.

As an alternative, we previously developed a PEG-based shell by click chemistry (Cutrone et al., 2019b). A block copolymer was synthesized by coupling onto a dextran (DEX) backbone two types of moieties: 1) PEG chains to reduce macrophage uptake and 2) alendronate (ALN) to enable a strong coordinative interaction with the nanoMOFs' surface. Both PEG and ALE contents were well controlled by click chemistry, however, the synthesis method involved many steps and the resulting copolymers were non-biodegradable.

Herein, biodegradable shells based on PEG and γ-cyclodextrin (CD) were successfully synthesized. First, CD-PEG monomers were prepared, where each CD was linked to a PEG 2000 g/mole polymer. They were crosslinked with citric acid, leading to the formation of CD-PEG-CO oligomers. Moreover, the citrate moieties played a major role in anchoring the oligomers onto the external surface of the nanoMOFs, due to strong coordination with the available iron sites from the nanoMOFs. The bulky structure of CD-PEG-CO prevented the PEG chains from penetrating inside the 3D nanoMOFs networks, allowing for the formation of PEG shells to evade the recognition by macrophages. What's more, the surface modification didn't influence the drug encapsulation. The drug of interest was doxorubicin (DOX), an anticancer molecule widely employed to treat many types of carcinoma, soft tissue sarcomas, and hematological malignancies (Aryal et after 60 h, while only 37% and 30% DOX were released at pH 6.8 and 7.4 after 60 h, respectively. However, the porous nanoMOFs were not coated in these studies.

A more systematical study was carried on here, incorporating DOX at different payloads in MIL-100(Fe) nanoMOFs coated or not with CD-PEG-CO and studying its release in various media. The optimized DOX formulations didn't release their content at pH 7.4 (PBS and cell culture medium), but rapidly released DOX in artificial lysosomal fluid (ALF) at pH 4.5. The anticancer efficacy of the drug loaded and surface-modified nanoMOFs with was evaluated.

The ability of the engineered nanoMOFs to evade macrophage uptake and to get internalized in cancer cells was also investigated. Cell internalization was visualized by confocal microscopy and quantified by inductively coupled plasma mass spectrometry (ICP-MS). The core-shell nanoMOFs demonstrated a high versatility in terms of functionalities and DOX controlled release.

Experimental

Materials and reagents

Chemical materials. Potassium chloride (KCl, Sigma-Aldrich, ≥99%), sodium phosphate monobasic dihydrate (NaH2PO4. 2H2O, Sigma-Aldrich, ≥99%), γ-cyclodextrin (γ-CD, Cyclolab, Budapest, Hungary, ≥90%), dimethylsulfoxide (DMSO, ACS), phosphate-buffered saline (PBS, Gibco) and RPMI (Roswell Park Memorial Institute) 1640 medium ( Gibco) were purchased from commercial sources and used without further purification otherwise indicated. sodium chloride (NaCl, Sigma-Aldrich, ≥99.5%), sodium hydroxide (NaOH, Emsure®, ≥99.0%), citric acid (Sigma-Aldrich, 99.0%), calcium chloride (CaCl2, Sigma-Aldrich, ≥99.0%), disodium phosphate (Na2HPO4, Sigma-Aldrich, ≥99.0%), sodium sulfate (Na2SO4, abcr Gmbh, Germany, 99.0%), Magnesium chloride (MgCl2, Sigma-Aldrich, ≥98.0%), glycerin (Sigma-Aldrich, ≥99.0%), citrate sodium dihydrate (Sigma-Aldrich, ≥99.0%), sodium tartrate dehydrate (abcr Gmbh, Germany, ≥99.0%), sodium pyruvate (Fluka, 99.0%) were used for the preparation of artificial lysosomal fluid (Table S1). Iron (III) chloride hexahydrate (FeCl3. 6H2O, Alfa Aesar, Schiltigheim, France, 98%), 1,3,5-benzenetricarboxylic acid (BTC, Sigma-dried mixture underwent a treatment at 140 °C for 15 ~ 30 min under reduced pressure (10-15 mmHg) in a system connected to a suction pump. A yellowish product was obtained in the bottom of the flask. Then, 8 mL deionized water was added and the crude was sonicated for 3 min, followed by filtration in order to separate the insoluble fraction from the soluble part. The soluble fraction was recovered and dialyzed for 48 h with a cellulose membrane (20 kDa or 50 kDa, Spectrum Laboratories, Rancho Dominguez, US). The solution was finally dried using a freeze dryer to yield a slightly yellow powder. In addition, CD-CO and CD-CO-Cy5 were synthesized as previously reported (Qiu et al., 2020).

The compositions of synthesized CD-PEG-CO and γ-CD-citrate oligomers (CD-CO) were evaluated by proton nuclear magnetic resonance ( 1 H-NMR) spectroscopy. 1 H-NMR spectra were recorded in D2O using a Bruker Avance 360 MHz. The citric acid/ CD-PEG molar ratio in the products was estimated by integrating the peaks assigned to 4 protons of the citric acid methylene group divided by 3 methyl protons of CD-PEG units. This allowed calculating the relative monomer:citrate molar ratios in the oligomers, and further have an estimation of the CD-PEG wt% contents, by taking into account the respective molecular weights of CD-PEG (3374 g/mol) and citrate (192 g/mol) moieties.

Size-exclusion chromatography (SEC), coupled on-line with multiangle light scattering (MALS) and refractive index (RI) detectors (SEC/MALS/RI, Wyatt Technology, Santa Barbara, USA), was used to determine the average molecular weight values of CD-PEG CO, using a dn/dc value of 0.14 mL/g(Puskás et al., 2013), two OHPAK SB 804/OHPAK SB806 HQ columns, and 0.1 M LiNO3 as eluent. Data were determined thanks to the Astra® 6.1.1 software from Wyatt Technology using Zimm order 1 between 44.8° and 140°.

Synthesis and characterization of MIL-100 (Fe) nanoMOFs

Iron trimesate nanoMOFs MIL-100 (Fe) were synthesized by using a previously described hydrothermal method assisted by microwave irradiation (Agostoni et al., 2013). Concisely, 20 mL of an aqueous mixture containing 0.84 g BTC and 2.43 g iron (III) chloride hexahydrate was heated at 130 °C for a short time (6.0 min) under stirring by microwave irradiation (Mars-5, CEM, USA). The nanoMOFs were harvested by centrifugation at 10,000 g for 10 min and purified by washing with absolute ethanol 6 times, to remove unreacted BTC and discard aggregates. The crystallinity was investigated by X-ray powder diffraction (XRPD) using a homemade setup based on a copper rotating anode generator (RU-200BEH, Rigaku Ltd., Tokyo, Japan), as previously described by our group (Li et al., 2021; Qiu et al., 2020). The samples were filled in round glass capillaries (1.5 mm in diameter, WJM-Glas Müller GmbH, Berlin, Germany). The X-ray beam was filtered and focused by a confocal system (CMF-12-38Cu6 from Osmic Inc., Cleveland, OH, USA). The accessible scattering vector range was 0.035 ~ 0.5 Å -1 .

The hydrodynamic diameters and size distributions were determined by dynamic light scattering (DLS, Malvern Nano-ZS, Zetasizer Nano series, France). The nanoMOFs BET (Brunauer-Emmett-Teller) surface area was evaluated by nitrogen sorption experiments at -196 °C using an ASAP 2020 (Micromeritics, USA) after degassing at 100 °C for 16 h under secondary vacuum. NanoMOFs were kept in EtOH at room temperature and re-suspended in aqueous media before use.

Surface modification of nanoMOFs with CD-PEG-CO

NanoMOFs were centrifuged at 6,000 g for 10 min to remove the storage solvent (ethanol) and then re-dispersed in water by vortex. For the coating procedure, they were incubated at room temperature with different initial amounts of CD-PEG-CO (at the weight ratio CD-PEG-CO: nanoMOFs of 1:10 ~ 1:1). After 1 day incubation, the non-attached CD-PEG-CO fraction was removed by centrifugation (10,000 g, 10 min). The same method was employed to coat the nanoMOFs with CD-CO.

Characterization of CD-PEG-CO coated nanoMOFs

Morphologies of nanoMOFs, before and after surface modification and/or drug loading were investigated using a JEOL 1400 transmission electron microscope (TEM, 120 kV, Japan). The crystallinity of the particles was characterized by X-ray powder diffraction (XRPD), as previously mentioned.

STEM associated with elemental chemical energy-dispersive X-ray spectroscopy (EDX) cartography was used to evaluate the elemental (C, O, N, S, Cl and Fe) distribution in the particles. The MIL-100 (Fe) nanoMOFs coated or not with CD-PEG-CO and loaded or not with DOX were placed on a copper grid covered with a pure carbon membrane. The STEM observations were made on a Titan Themis 200 microscope (FEI/Thermo Fischer Scientific) equipped with a geometric aberration corrector on the probe. This microscope was also equipped with the "Super-X" systems for EDX analysis with a detection angle of 0.9 sr. The observations were made at 200 kV with a probe current of about 70 pA and a half-angle of convergence of 24 mrad. High-angle annular dark-field (HAADF)-STEM images were acquired with a camera length of 110 mm (inner/outer collection angles were respectively 69 mrad and 200 mrad).

Colloidal stability investigation. Colloidal stabilities of CD-PEG-CO coated nanoMOFs in

water were evaluated by DLS. Mean hydrodynamic diameters were measured at 0, 1, 2, 6, 10, and 14 days storage at 25 °C. Zeta potential (ZP) of nanoMOFs were determined at 25 °C using a Zetasizer instrument (Malvern Nano-ZS, Zetasizer Nano series, France) in a pH range of 1 to 10. For the measurements, nanoMOFs were diluted using a KCl solution (1 mM) to a final concentration of 100 μg/mL. Thermogravimetric analyses (TGA) analyses were used to quantify the amount of coating materials attached on the surface of nanoMOFs using a TGA 4000 Instruments (Perkin Elmer, 100-240 V/50-60 Hz) using our previous protocol (Qiu et al., 2020). During analysis, samples were submitted to an O2 flow of approximately 20 mL/min. Around 10 mg of nanoMOF samples, coated or not, were heated from 25 to 600 °C at a rate of 3 °C/min. Prior to the TGA analysis, the samples were dried overnight at 60 °C to remove the absorbed water. The associated amounts of CD-PEG-CO (wt%) were calculated with respect to dry MIL-100(Fe) nanoMOFs weight.

DOX encapsulation in CD-PEG-CO or CD-CO coated nanoMOFs

MIL-100(Fe) nanoMOFs suspension (8 mg/mL) was first centrifuged at 10,000 g for 10 min to recover the nanoMOFs, which were further re-suspended in water. For the DOX encapsulation, 0.25 mL of nanoMOFs aqueous suspension (4 mg/mL) was mixed with different volumes of DOX solution (5 mg/mL), with or without 125 μL CD-PEG-CO (4 mg/mL) or CD-CO (4 mg/mL). The final volume was adjusted to 1 mL. Theoretical drug loading (TDL), calculated as the weight ratio (%) between DOX and nanoMOFs, was in the range of 10-100%. The resulting suspensions were gently stirred for 1 to 6 days at room temperature to allow DOX incorporation. Then, 1 mL of nanoMOF suspensions were centrifuged (17,000 g, 15 min) and both supernatants and pellets were recovered to assess the amounts of incorporated DOX.

The supernatants were used to quantify the drug payload, which was calculated as Eq. ( 1):

𝑃𝑎𝑦𝑙𝑜𝑎𝑑 (%) = Encapsulated Drug (mg) NanoMOFs (mg) × 100 (1)
Encapsulation efficiency (EE) was calculated, as according to Equation (2):

𝐸𝐸 (%) = Encapsulated Drug (mg) Initial Drug (mg) × 100 (2)
DOX quantification was performed using fluorescence spectrometry with the excitation (λex) and emission (λem) at 500 and 590 nm, respectively. DOX in the supernatant was diluted with a water/DMSO mixture (1:1 v/v), to dissociate the complexes γ-CD: DOX. Calibration curves of DOX were obtained between an interval of 10 -3 to 10 -4 mg•mL -1 (correlation coefficient > 0.999). Pellets were digested in 1 mL of DFOM (10 mg/mL in PBS), then incubated at 37 o C in a rotator with a rotational speed of 200 rpm for overnight, then diluted with a water/DMSO mixture (1:1 v/v) to determine DOX payloads.

Release of DOX from MIL-100 (Fe)

The DOX release experiments were performed by incubating the nanoMOFs in different media (PBS, RPMI, AFL and water) at 37 °C in a rotator with a rotational speed of 200 rpm. At periodic time intervals, 1 mL of the suspensions (100 μg/mL) were centrifuged (17,000 g, 15 min) and the supernatants were recovered for DOX quantification.

The DOX concentration in the supernatant media was analyzed by fluorescence spectrometry.

The percentage of the released DOX was calculated as Eq. ( 3):

𝑅𝑒𝑙𝑒𝑎𝑠𝑒𝑑 𝐷𝑂𝑋 (%) = DOX amount in the supernantant DOX amount loaded in nanoMOFs × 100

The release study was further confirmed by digesting the pellet in 1mL of DFOM (10 mg/mL in PBS), then incubated at 37 o C in a rotator with a rotational speed of 200 rpm for overnight, after that diluted with water/DMSO mixture (1:1 v/v) to determine the amount of DOX in the pellets.

NanoMOF interaction with macrophages

NanoMOFs quantification inside cells using ICP-MS. The amounts of nanoMOFs internalized inside the cells were determined by iron quantification using ICP-MS. Briefly, macrophage cells (J774A.1) were seeded as previously described. Cells treated or not with inhibitors were then incubated with 50 μg/mL nanoMOFs coated or not with CD-PEG-CO or CD-CO for 4 h. Then, the cells were washed with PBS for three times to eliminate the nanoMOFs in the medium. Four different wells were used to prepare the cells at one condition,

where the cells in one well were counted and the cells in other three wells were finally dried at 60°C for 2 h and digested using aqua regia (15 minutes under sonication) for Fe quantification.

Fe quantification was carried out by ICP-MS (Agilent 8800) as previously reported (Cutrone et al., 2019b, 2019a; Li et al., 2020b). Co was added at a concentration of 10 μg/L as internal standard. Fe and Co isotopes were detected using "on-mass mode" (54Fe + , 56Fe + , 59Co + ). Fe was finally quantified using external calibration prepared from certified 1000 mg/L Fe standard solution (Merck, Germany).

Anticancer efficacy of drug loaded particles evaluated on Hela cells

Confocal visualization. Hela cells were seed at the density of 3.0×10 4 cells/well on sterile glass slides placed at the bottom of each well and let incubate overnight. Then, nanoMOFs empty or loaded with DOX (at TDL of 20%, and 50%) and coated or not with CD-CO-Cy5 and CD-PEG-CO (50 wt%) were added at a concentration of 50 µg/mL in each well. By simply mixing with CD-PEG-CO, the nanoMOFs were efficiently labelled after incubation overnight. After incubation with living cells for 2 h, 6 h, and 24 h, cells were washed three times with PBS to remove free particles. The living cells were visualized using a confocal microscope (LEICA SP5 system), equipped with a CO2 regulated (5% CO2) and a thermostatically controlled chamber (37°C). DOX and Cy5 were excited at 488 nm and 614 nm, respectively. The emissions from 550-650 nm and 650-750 nm were collected for DOX and Cy5, respectively.

The images were processed with the Image J software.

NanoMOFs internalization in Hela cells quantified by ICP-MS.

The same protocol for J774A.1 cells was used for Hela cells, but no inhibitor was used. The internalized nanoMOFs were directly quantified by ICP-MS as described in section 2.2.7.

Cytotoxicity assays.

The cytotoxicity of nanoMOF samples, with and without coatings and loaded or not with DOX was evaluated by MTT. First, Hela cells were seeded in 96 well plates at the density of 1.0×10 4 cells/well and incubated overnight. Then, the Hela cells were exposed to various concentrations (0-100 µg/mL) of nanoMOFs coated or not with CD-CO/CD-PEG-CO and/or loaded with DOX (at TDL of 20%). After 24 h incubation, all the medium in the wells was removed and excess unbounded nanoMOFs were discarded by rinsing two times with PBS. Then, the cells were incubated for another 4 h with an MTT solution (100 µl, 500 µg/mL).

Finally, 100 µl DMSO was introduced in each well to dissolve the formazan crystals followed by absorbance detection using a plate reader (GloMax® Discover Microplate Reader, Promega) at the wavelength of 490 nm.

Results and discussion

Synthesis and characterization of CD-PEG oligomers

Synthesis of CD-PEG oligomers.

Prior to CD-PEG oligomers synthesis, the CD-PEG monomer was successfully synthesized by click chemistry.

The CD-PEG oligomers were synthesized by adjusting a previous method reported for CD-CO (Qiu et al., 2020) (Scheme 1). The polymerization reaction between CD-PEG and citric acid was investigated by optimizing the heating time and the molar ratio between CD-PEG, citrate acid and catalyst. The different synthesis conditions are summarized in Table 1. Initially, the reaction mixture was concentrated by evaporation, and the resulting dried mixture was heated at 140 °C for other 15-30 min under reduced pressure (entries 1-4). The average molar masses (Mn) increased with the heating time from 7500 to 11200 g/mole for 25 min and 40 min, respectively, whereas the polymerization yield was practically unchanged (30-35 %).

However, the polydispersity increased from 1.7 (25 min) to 4.7 (40 min) (entries 1-4,) which could be possibly explained by random crosslinking in between the CD-PEG monomers in the viscous reaction mixture. Therefore, to maintain low polydispersity values, the total heating time was fixed at 25 min, and the molar ratio among CD-PEG, citrate acid and catalyst was systematically investigated (entries 5-11). It is found that when the molar ratio among CD-PEG, citrate acid and catalyst was 1:15:4, the Mw was high (33,600 g/mole), with acceptable polydispersity (Ð=1.8) (entry 7). Furthermore, the oligomers with the lowest molar masses were removed by purification by dialysis using a membrane with a cutoff of 50 kDa, resulting in a series of copolymers with large Mw (entries 12-15; SEC chromatograms can be found in Fig. S1). The best condition was entry 12, corresponding to a Mw of 53400 g/mole and Ð=1.7. The composition of this CD-PEG-CO copolymer and of the corresponding CD-PEG monomer were assessed by 1 H-NMR spectroscopy (Figs. S2-S3). Fig. S3 shows that as expected, the composition of CD-PEG corresponds to one PEG chain grafted to each CD unit, in average.

In the case of CD-PEG-CO, the characteristic peak of citric acid (f) was well observed, indicating its efficient incorporation in the copolymers (Fig. S3). The citric acid/CD-PEG molar ratio was estimated by integrating the peaks assigned to the 4 protons of the citric acid methylene groups (f) and the three methyl (a) protons of the CD units. The molar ratio: citric acid/CD-PEG was found as 5.9, corresponding to a CD-PEG-CO with an average composition (CD-PEG)12 (citrate)72, with high PEG contents (45 wt%).

XRPD patterns (Fig. 1E). Furthermore, high resolution STEM investigations clearly evidenced the crystalline structures with well delimited planes (Fig. 2). It was possible to access to the nanoMOFs' 6-fold symmetry and interplanar spacings (Fig. S4). The obtained values of 2.61 Surface modification of nanoMOFs with CD-PEG-CO was achieved by a "green" method (free of organic solvents and surfactants), consisting in a one-step incubation at room temperature of the nanoMOFs in an aqueous solution of the synthesized oligomers. The procedure was similar to the one described to coat similar MIL-100(Fe) nanoMOFs with CD-CO oligomers (Qiu et al., 2020) or with phosphorylated CD derivatives (Agostoni et al., 2015). After overnight incubation, CD-PEG-CO coated nanoMOFs were harvested by centrifugation and the amount of CD-PEG-CO associated to the nanoMOFs was determined by TGA (Fig. 1F). In the thermograms of the uncoated nanoMOFs, it can be observed that a residual inorganic Fe2O3 material (~35 wt%) remained at 600 °C, in agreement with our previous reports (Qiu et al., 2020). The amount of residual material was diminished in the case of CD-PEG-CO coated nanoMOFs, to 25±2 wt%, 26±2 wt% and 27±2 wt% when the CD-PEG-CO: nanoMOFs weight ratio in the coating procedure was 1:1, 1:2 and 1:3, respectively. This correspond to CD-PEG-CO associated amounts of 32±2, 29±2 and 24±2 wt%, respectively.

In conclusion, the maximal amount of CD-PEG-CO associated to the nanoMOFs in these studies reached to around 32 wt%, which is in the same order of magnitude as in the case of other PEGylated coatings such as those obtained by GraftFast (29 wt%)(Giménez-Marqués et al., 2018), and by using comb-like PEGylated dextran copolymers (32 wt%) (Cutrone et al., 2019b). Noteworthy, the amount is twice higher than for CD-Phosphate-PEG coatings ( 16wt%) (Cutrone et al., 2019a). The strong affinity of numerous carboxyl groups in CD-PEG-CO for the iron sites on the nanoMOFs' surface most probably accounts for this efficient association. There is no significant difference in terms of the CD-PEG-CO associated amount between the samples prepared at weight ratio CD-PEG-CO: nanoMOFs of 1:1 and 1:2, however, the coating yields were 32% and 58%, respectively. Therefore, the coating condition was fixed to a weight ratio CD-PEG-CO: nanoMOFs of 1:2 for the following studies.

Remarkably, the CD-PEG-CO coated nanoMOFs maintained their size, facetted morphology (Fig. 1B) and crystalline structure (Fig. 1E) despite the high amounts of associated coatings.

There was a slight size increase in hydrodynamic diameters, from 255 ± 14 nm to 268 ± 15 nm, before and after CD-PEG-CO coating, respectively (Fig. S5) which could be due to the presence of the shell.

Furthermore, in contrast to the uncoated nanoMOFs, the CD-PEG-CO coated ones were stable upon storage up to 2 weeks (less than 15% size variation, Fig. S5). The presence of the CD-PEG-CO coating was further supported by the significant ZP variations as compared to uncoated particles (Fig. 3B). The first observation was that the ZP of the nanoMOFs, before and after surface modification, was strongly dependent upon the pH (in a range of 1~10) of the suspension media, in agreement with previous reports (Li et al., 2019a). The ZP of assynthesized nanoMOFs was positive (+24±2 mV) at pH 5, but shifted to negative values (-15±3 mV) at the same pH value after coating with CD-PEG-CO. This is possibly due to the presence of citrate-based coatings with a negative charge, as reported in previous studies (Cutrone et al., 2019b).

DOX encapsulation and drug release

After the successful surface functionalization, DOX was encapsulated as drug of interest to evaluate if the core-shell NPs are able to encapsulate and release drug cargoes. First of all, DOX encapsulation was performed by simply incubating the nanoMOFs in aqueous DOX solutions under gentle rotating agitation at room temperature. Fig. 3 A indicates the kinetics of DOX incorporation in 6 days quantified by fluorescence spectroscopy. Briefly, DOX loading kinetics were studied at different DOX initial amounts while the TDL of DOX was 10%, 20%, and 50%.

When TDL was 10%, the DOX payload reached 9±0.2 wt% within 1 day and plateaued in 2 days at 9.6±0.2 wt%, indicating the excellent DOX encapsulation efficiency up to 96%. When TDL was increased to 20%, the DOX payload plateaued in 3 days at 19±1.0 wt%, corresponding to more than 95% of DOX encapsulation efficiency. Remarkably, the DOX payload reached up to 32±1.5 wt% after 1 day impregnation when TDL was 50%. It further mounted to 41±2 wt% after 6 days impregnation. Complementary TGA experiments confirmed these findings (Fig. 1F) showing a 42±2 wt% DOX payload. These results clearly showed that MIL-100(Fe) nanoMOFs act as "nanosponges", effectively adsorbing DOX in their cages. Surprisingly, when the TDL was increased to 100%, the DOX payload reached up to 65±3.0 wt% (Fig. S6), which is similar to the reported DOX loading (62 wt%) in MIL-100 (Al) metal-organic gels (Feng et al., 2018). However, this DOX loaded amount is theoretically larger than the available nanoMOF pore volume (Agostoni et al., 2013), suggesting that DOX most probably be located both inside the pores and on the surface of the nanoMOFs. This hypothesis was supported by ZP investigations, before and after DOX loading (Fig. 3B).

At the DOX payload of 19±1 wt%, there was no significant ZP variation between loaded and unloaded nanoMOFs in the pH range from 7 to 10. On the contrary, ZP significantly shifted to more positive values when DOX payload increased to 41±2 wt%. As an example, ZP values were -16±4 mV and +7±2 mV for empty nanoMOFs and DOX loaded nanoMOFs, respectively, at a pH of 7. This could possibly be attributed to the cationic character of DOX located at the nanoMOFs' surface (Unnier et al., 2007). This supports the hypothesis of DOX location in the nanoMOFs' top layers at highest (41 ± 2 %) drug loadings as compared to 19±1wt%.

The maximal DOX payload in MIL-100 (Fe) nanoMOFs reported in the literature was 33 wt% Noteworthy, whatever the DOX loadings, the nanoMOFs' faceted morphology was preserved, before and after DOX encapsulation, as shown by TEM investigations (Fig. 1C). In addition, at the highest DOX loading (42 wt%), the cages of the nanoMOFs were well observed (Fig. 2B and 2D). PXRD patterns of DOX-loaded nanoMOFs demonstrated that the crystalline structure of the MIL-100 (Fe) nanoMOFs was maintained despite of their unprecedented DOX payloads (Fig. 1E). Advantageously, regardless of their DOX loading, all the nanoMOFs samples were stable for at least two weeks. (Fig. S5).

Interestingly, the surface modification with CD-CO and CD-PEG-CO didn't significantly affect the DOX loading (Fig. 3A). Due to the guest-host inclusion between DOX and γ-CD, CD-CO accelerated the drug loading process, in agreement with our previous investigations (Qiu et al., 2020).

Finally, there were no significant differences in the drug loading values, at the plateau reached after 6 days impregnation. When the DOX TDL was 50%, the measured DOX loadings were 41±2 wt%, 44±3 wt%, and 40±2 wt% for bare nanoMOFs, nanoMOFs coated with CD-CO, and CD-PEG-CO, respectively.

Similarly, when the DOX TDL was 20%, the DOX payloads were 15±2% and 17±3%, for nanoMOFs coated or not with CD-PEG-CO. Moreover, despite the high drug loading and surface modification with CD-PEG-CO, the nanoMOFs maintained their crystalline structures (Fig. 1E) and facetted morphologies (Fig. 2C). Complementary investigations were carried out by EDX to gain further insights on DOX distribution within the nanoMOFs and to estimate the individual nanoMOFs drug loadings. To track the drug, EDX mappings of Cl and N elements specific of DOX and not present in nanoMOFs were performed (Fig. 4). Only Cl mapping gave reliable information, as the N signal was shielded by the large signals arising from C and O.

Table 2 summarizes the estimated drug loadings of randomly chosen nanoMOFs. In average, the DL was 15.7±1.9 and 15.2±3.9 wt% for the uncoated and CD-PEG-CO coated nanoMOFs, respectively. The EXD data are just estimations of the loadings, but the results clearly prove that the nanoMOFs are homogeneous in terms of composition and that the presence of the CD-PEG-CO coatings do not influence the drug loading. DOX release studies were further carried out systematically in different media, i.e. simulated physiological medium (PBS buffer), cell culture medium (RPMI), and simulated intracellular medium (ALF). As shown in Fig. 5, the DOX release is dependent on the drug loading. At low DOX payload (9.6±0.2 wt%), DOX release in both PBS (pH = 7.4) and RPMI (pH = 8.0) approached 12 ± 1.2% within 6 h, and it reached 88 ± 2.3% and 70 ± 2.2% in 1 month (Fig. 5 A and B).

When the DOX payload increased to 19 ± 1.0 wt%, DOX released amounts were reduced. For example, only 9 ± 0.4% and 60 ± 2.0% DOX amounts were released in PBS in 6 h and 1 month, respectively. Similar release profiles were found for DOX release in RPMI (9 ± 1.0% and 38 ± 3.0% for DOX release in 6 h and 1 month, respectively). At high DOX payload of 41 ± 2.1 wt%, DOX didn't release in both PBS and RPMI within 48 h (<1%) and there were only 16 ± 3.2% and 12 ± 1.0% released in 1 month in PBS and RPMI, respectively. These results were confirmed by quantifying the DOX amount in the pellet, For instance, when the DOX payload was 19±1.0 wt%, only 9.2±0.3% and 61±1.8.0% DOX amounts were released in PBS in 6 h and 1 month, respectively.

The incomplete DOX release at high loadings could be explained by the self-association of the drug molecules confined inside the nanoMOFs cages. Similar results were reported in the case of topotecan-loaded MIL-100(Fe) nanoMOFs (Di Nunzio et al., 2014). When concentrated inside the cages, topotecan was prone to aggregate by -stacking and the formed cluster were too large to be released out.

In contrast, the DOX release rate was remarkably increased in ALF (pH = 4.5) for all the studied cases (Fig. 5C, and Fig. S7). For instance, there was around 90±3% and 94±2% DOX release in 6 h and 24 h when loaded at 9.6±1 wt% (Fig. S7). Even at DOX loading of 41±2 wt%, a very high DOX release up to 86±3% was observed in 24 h (Fig. 5C). Interestingly, the "burst" effect was significantly reduced when the DOX-loaded nanoMOFs were coated with CD-CO or CD-PEG-CO (64±1%, 46±1% and 42±3% was released in 1 h from bare nanoMOFs, nanoMOFs coated with CD-CO, and CD-PEG-CO, respectively). However, all the samples reached a similar plateau (80-90% release) in 24 h. Noteworthy, non-released DOX could be extracted from the nanoMOFs using DFOM, after the plateau was reached. Results were in agreement with HPLC data, confirming that the release was not complete. TDL was 50% (dashed line ----), TDL was 20% (straight line), and TDL was 10% (dashed line ). DOX loadings were 41±2 wt%, 44±3 wt%, and 40±2 wt% for bare nanoMOFs, nanoMOFs coated with CD-CO, and CD-PEG-CO (dashed line ----); 19.2±0.9 wt%, 19.7±1 wt%, and 19±0.6 wt% for bare nanoMOFs, nanoMOFs coated with CD-CO, and CD-PEG-CO (straight line); 9.8±0.2 wt%, 9.7±0.2 wt%, and 9.6±0.2 wt% for bare nanoMOFs, nanoMOFs coated with CD-CO, and CD-PEG-CO (dashed line ).

After the systematic characterization of DOX loading, surface functionalization, and drug release, we investigated the biological effect of the prepared formulations. Firstly, the stealth effect of CD-PEG-CO modified nanoMOFs was evaluated in vitro on macrophages (J774A.1 cells). The quantity of nanoMOFs internalized in the J774 was determined using ICP-MS as previously described (Cutrone et al., 2019b). This quantification of iron content in cells was performed to study the nanoMOFs cellular uptake. Around 75% uncoated nano-MOFs/3×10 5 cells were taken up by J774A.1 cells after 4 h incubation, thus showing that macrophages avidly took up uncoated nanoMOFs, in agreement with our previous results (Cutrone et al., 2019a, 2019b; Li et al., 2020b). There was about 70% CD-CO coated nanoMOFs internalized in macrophages, meaning that the CD-CO coating didn't have stealth effect. Interestingly, the presence of the CD-PEG-CO coating reduced the nanoMOFs internalization to 50% at the same incubation time. What's more, the internalized amount of CD-PEG-CO coated nanoMOFs was further reduced to 35% and 40% in the presence of the inhibitors latrunculin A and cytochalasin D, respectively (Fig. S8), which block actin polymerization and hence disrupt microfilamentmediated internalization, manifesting the role of phagocytosis. These results are agreement with our previous studies (Li et al., 2019a).

Cytotoxicity assays of nanoMOFs on cervical cancer cells

All the studied nanoMOFs, coated or not, were nontoxic for the Hela cervical cancer cells up to 100 µg/mL (Fig. 6A), with more than 87% cell viability after 24 h incubation, which is in line with previous reports (Baati et ). Noteworthy, the newly synthesized CD-CO and CD-PEG-CO oligomers were not toxic neither, up to 100 µg/mL. In contrast, as expected, the anticancer drug DOX (Fig. 6B, blue histograms) exerted a cytotoxic effect with 67% cell viability at 1 µg/mL, which further diminished to 33% at 5 µg/mL. Indeed, free DOX molecules are efficient in eradicating cancer cells, since they can easily pass through their cell membrane and reach the nuclei. In this context, the engineered nanoMOFs could play a role in vivo, acting as nanocarriers which maintain their DOX cargo until they reach the cancer cells, where they would release DOX, thus reducing DOX toxicity in healthy organs. Keeping these aspects in mind, preliminary in vitro studies were carried on here. The cytotoxicity of DOX-loaded nanoMOFs was dependent on their internalization inside the cancer cells (Fig. 7). According to ICP-MS quantifications, around 15% of the nanoMOFs put in contact with the Hela cells were internalized after 24 h incubation. For example, when DOX loaded nanoMOFs was incubated with Hela cells at DOX concentration of 5 µg/mL, the internalized DOX was around 0.75 µg/mL. Taking into account the amount of internalized nanoMOFs, the DOX loaded nanoMOFs didn't significantly affect the DOX efficacy (60% and 67% for DOX loaded nanoMOFs at 5 µg/mL and free DOX at 1 µg/mL, respectively). These results are in agreement with reported data (Xue et al., 2019), where around 65% cell viability was observed for MCF-7 cells after 24 h incubation with DOX-loaded MIL-100(Fe) nanoMOFs. 2015). It was shown that nanoMOFs acted as "Trojan horses" by internalizing inside the cancer cells, carrying their drug cargo to eradicate cancer cells (Li et al., 2019b). In this study it was shown that interestingly, the presence of a CD-based coating (PEGylated or not) did not reduce the nanoMOF anticancer efficacy on Hela cells.

Localization of nanoMOF in Hela cells

The internalization of nanoMOF in Hela cells was further investigated using confocal microscopy. To do so, fluorescent CD-CO was synthesized as previously described by crosslinking CD-Cy5 with citric acid (Qiu et al., 2020) and was used as coating to track the nanoMOF internalization and intracellular stability of the coating material. The fluorescent dye Cy5 with Ex/Em of 647/666 was chosen in order to avoid overlapping with the signal rising from DOX molecules (Ex/Em=470/593). The CD-CO-Cy5 coating showed a low detachment from the nanoMOFs in cell culture medium (< 25% after 24 h incubation). Thus, taking advantage of the fluorescent DOX molecules and the low drug release in cell culture medium (< 10% in 24 h, Fig. 5), the drug localization was also visualized simultaneously with the coating. As shown in Fig. 7, after 2 h incubation, the nanoMOFs were efficiently internalized inside Hela cells maintaining their coatings and the encapsulated DOX. Indeed, the green (Cy5labelled coating) and red (DOX) signals were colocalized (Fig. 7C). The nanoMOFs preferentially located in the cytoplasm at early stage, and no fluorescent signal was found in the nuclei. These findings suggest that the nanoMOFs' coatings were intact bearing and the incorporated DOX was not released. This is in agreement with the reported confocal images in the case of DOX loaded mesoporous silica NPs modified with poly acrylic acid (Minati et al., 2013), where DOX signal was observed only in cytoplasm after 2 h incubation. However, the coating material was not labelled in this case.

In the case of DOX-loaded MIL-100 nanoMOFs with CD-PEG-CO coating, after 6 h incubation, the situation was totally different and the two fluorescent signals were not colocalized anymore: DOX diffused into the nucleus but the coating material remained in the cytoplasm (Fig. 7F). This suggests that the coating materials most probably detached from nanoMOFs. This is in agreement with previous reports that nanoMOFs didn't penetrate inside the cell nuclei (Li et al., 2020a) and that nanoMOFs were prone to degraded within a few hours inside cells (Li et al., 2019a), releasing the drug cargos. Similar phenomenon was observed in the case of DOX loaded nanoassemblies composed of poly (ethylene glycol)-β-cyclodextrin and a disulfide-containing adamantine-terminated doxorubicin prodrug (Xiong et al., 2018).

DOX started to penetrate inside the cell nuclei after 8 h incubation with HepG2 cells. In the case of DOX-loaded mesoporous silica NPs coated with poly acrylic acid, the authors reported that DOX was mainly located in cell nuclei after 17 h incubation. However, no intermediate incubation times were investigated between 2 and 17 h (Minati et al., 2013).

Presumably, in the case studied here, the DOX molecules released out of the nanoMOFs exerted their cytotoxic effect in the nucleus by the well-known mechanisms consisting in intercalating within DNA base pairs, causing breakage of DNA strands and inhibition of both DNA and RNA synthesis (Johnson-arbor and Dubey, 2020). Indeed, after 24 h interaction with the DOX-loaded nanoMOFs, the Hela cells were practically all eradicated, and the few remaining cells showed shrinked morphologies and reduced nuclei sizes. In addition, the Cy5 also started to quench, exhibiting a much less bright fluorescent intensity. In a nutshell, these studies demonstrates that nanoMOFs localized in the cytoplasm where they degraded, releasing the coating materials from the shell and the DOX molecules from the core. The detached coating shell remained in the cytoplasm, whereas the released DOX diffused inside the nucleus, causing DNA damage. Noteworthy, the shell was released inside the cells, but not in the outer medium.

Conclusion

These studies pave the way to the design of engineered core-shell nanoMOFs that do not release their drug content in biological media, but only in intracellular environment. Further studies will concern nanoMOF surface modifications with CD-PEG-CO coatings bearing targeting moieties.

General Discussion and Perspectives

Recent developments in the field of nanomedicine have demonstrated that the surface modification of nanoparticle (NPs) plays an important role in the drug efficacy, being a key point for their successful biomedical application 1,2 . Among an increasing number of drug nanocarriers developed so far, nanoscale hybrid iron carboxylate Metal-Organic Frameworks (nanoMOFs) have been reported as one of the most promising drug carriers [1][2][3][4][5][6] based on: i) their adjustable porosity and versatile composition, , ii) low in vivo toxicity 7 , iii) unprecedented loading capacities of a variety of drugs together with iv) the ability to achieve controlled release under physiological conditions and v) their biocompatibility and biodegradability.

It was reported that surface modification with hydrophilic PEG chains in a "brush" configuration could dramatically prolong the blood circulation times of NPs by reducing their recognition by the reticuloendothelial system 8,9 . However, up to now, only a few number of cases have been reported on engineering the surface of nanoMOFs with PEG shells 10,11 . Indeed, it is more challenging to functionalize the porous surface of MOFs, as compared for example, to dense NPs composed of biodegradable polymers or to liposomes. Indeed, it has been reported that PEG chains can penetrate into the MOF pores, decrease the drug loading capacity, and lead to uncontrolled drug release 10 . In order to prevent the PEG from penetrating into the highly porous MOFs, the method of GraftFast was employed to form PEG-based shells. This method was based on polymerization of acryl PEGs 11 . However, the molecular weight of the resulting PEG-based copolymer could not be effectively controlled.

Therefore, there is an urgent need to develop an engineered PEG coating which can efficiently decrease reticuloendothelial sequestration of nanoMOFs. Herein we employed two strategies to engineer nanoMOFs surface-modified with PEG chains. The first strategy is to use the bulky macromolecule DEX as a backbone, followed by coupling two types of moieties: i) PEG chains to reduce macrophage uptake and ii) alendronate (ALN) to spontaneously coordinate to the external surface of nanoMOFs (chapter 2). The second strategy is to synthesize oligomers based on the cage molecule "CDs" coupled with PEG chains (chapter 3). The first strategy adopted mainly click chemistry to "click" different functional moieties to the backbone whereas the second strategy was to "click" PEG chains to CDs, followed by crosslinking the CD-PEG units.

These two strategies share common characteristics and at the same time both of them have their own advantages and disadvantages.

First strategy: synthesis of DEX-ALN-PEG

In chapter 2, both PEG and ALN were successfully grafted onto the DEX backbone and the number of grafted units was controlled. Three dextran derivatives with different molar ratios of ALN and PEG branches were synthesized, namely DEX-ALN75-PEG25, DEX-ALN50-PEG50

and DEX-ALN25-PEG75, which were confirmed by 1 H NMR, [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C NMR and 31 P NMR spectroscopies. MIL-100(Fe) nanoMOFs with a mean hydrodynamic diameter around 200 nm were successfully synthesized by a green method, without organic solvents, assisted by microwave and within a short time. The three copolymers with increased PEG contents were used as coating materials for the surface modification of nanoMOFs, following a fast, green method (without any organic solvent or surfactant) consisting of simply soaking nanoMOFs in aqueous solutions of the prepared copolymers. This facile coating protocol, devoid of the use of any organic solvents or surfactant, is beneficial for the applications in biomedical field and for scaling up. In order to gain deep insights on protein adsorption, the interaction of nanoMOFs, coated or not with copolymers, with HSA, the most abundant protein in human blood plasma, was investigated. The optimized DEX-ALN-PEG coatings could reduce five times the adsorption of the model protein HSA. There results clearly reveal that the DEX-ALN-PEG coating materials greatly decrease protein adsorption, which plays an important role on macrophage uptake.

The ability of DEX-ALN-PEG coated nanoMOFs to evade macrophage uptake was assessed on the murine macrophage J774A.1 cell line. After 2 h incubation, the optical microscopy images illustrated that a large amount of nanoMOFs clustered inside the macrophages. On the contrary, with the same incubation time, the uptake of DEX-ALN25-PEG75 coated nanoMOFs was dramatically decreased. To further quantify the macrophage uptake, ICP-MS was employed in this study. These results reveals the efficacy of the DEX-ALN-PEG coatings upon macrophage uptake. After 4 h incubation, the nanoMOFs amounts internalized in the cells were 77, 39, 32 and 24% for control (uncoated nanoMOFs), DEX-ALN75-PEG25, DEX-ALN50-PEG50, and DEX-ALN25-PEG75, respectively. This demonstrates that the higher the PEG contents in the functionalized material, the lower the nanoMOFs uptake. To the best of our knowledge, these are the first quantitative results on PEG-coated nanoMOFs uptake by macrophages based primarily on a direct quantification and visualization of the iron content in the cells.

In spite of all their advantages (easy to click a series of functional moieties, stealth effect, green method for coating procedure, etc), the DEX-ALN-PEG copolymers are not biodegradable, involve multiple synthesis steps, and therefore their synthesis is not easy to scale up. This can hamper their application in the biomedical field.

Second strategy: synthesis of CD-PEG-CO

To improve the biodegradability of the PEG-based coating materials, our second strategy was to synthesize oligomers based on CDs. It is recognized that CDs possess excellent properties to form host-guest inclusion complexes in aqueous solutions with various hydrophobic active molecules, due to their peculiar structure with a hydrophobic interior and a hydrophilic exterior.

Therefore, we employed CDs in the shells to elaborate versatile coating materials. As shown in chapter 3.1, a core-shell nanomaterial was devised using MIL-100 (Fe) nanoMOFs as the core and a CD-based layer as the shell.

Synthesis of CD-CO

First of all, biodegradable CD-CO (with and without fluorescent dye) were prepared adopting citric acid as cross linker, which plays an important role in anchoring the oligomers onto the surface of the nanoparticles due to the coordination between the citric acid and the available iron sites from MIL-100 (Fe). As far as we know, it is the first reported case that fluorescent CD-CO were synthesized to further understand the interaction between the shell and the core.

The synthesized CD-CO with and without fluorescent dye were characterized by a set of complementary techniques, including FTIR, SEC and NMR spectroscopies. The other novelty in this study was the use of fluorescence-lifetime microscopy to characterize the coating processes. It is well recognized that fluorescence lifetime is an inherent property of a fluorophore, no matter of its concentration. However, it could change partially according to its interaction with its environment (solvent, quenchers, NPs, etc). Indeed, the mean fluorescence lifetime (τ) of fluorescent CD-CO was remarkably modified upon interaction with nanoMOFs. Fluorescence-lifetime microscopy provided a straightforward method for an in situ characterization of the interactions between nanoMOFs and their coating materials, playing an important role to gain deep insights at the molecular level.

In the subchapter 3.1, DOX payload in nano-MOFs reached up to 48 ± 10 wt% after 24 h incorporation when the nanoMOFs and DOX were incubated at a weight ratio of 1:1 . DOXloaded nanoMOFs had a tendency to aggregate (mean hydrodynamic diameter raised to > 750 nm in 10 days). On the contrary, the colloid stability of DOX-loaded nanoMOFs was dramatically improved (size variation <10% in 3 weeks) after coating with CD-CO. Prior to the NMR study, DOX impregnation kinetics were studied at different DOX initial amounts (TDL of 10%, 20%, 50%). The maximum DOX payload was around 28 wt% after 6 days incubation, which obviously demonstrates that there is a strong interaction between DOX and the MIL-100 nanoMOFs. The morphology with faceted-type structures and the XRPD crystalline structure of the MIL-100 (Al) nanoMOFs were not affected after DOX encapsulation. Remarkably, with the highest DOX payloads, the BET surface area were greatly decreased (1720 ± 65 m 2 /g, and 200 ± 35 m 2 /g, for nanoMOFs before and after DOX loading (27±2.0 wt%), respectively, indicating that DOX penetrated inside the pores of the nanoMOFs.

Meanwhile, the CD-CO( 13 C) was successfully synthesized to track the localization of CD-CO in nanoMOFs and the interaction between CD-CO and DOX. Similarly to MIL-100 (Fe) nanoMOFs, MIL-100 (Al) nanoMOFs were easily modified with CD-CO ( 13 C) by soaking in water at room temperature. The morphology and crystallinity of the nanoMOFs was preserved after surface modification. BET surface area was only slightly lowered (1720 ± 65 m2/g &1500 ± 130 m2/g for nanoMOFs before and after surface modification), which demonstrates that the coatings mainly remained on the surface of the nanoMOFs.

Interestingly, the 1 H-1 H NMR spectra showed close contact between the DOX and the CD-CO inside the particles. The aromatic protons of the DOX have much higher intensity on the 1H NMR spectrum after transfer from the [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C-citrate than in the normal 1H MAS NMR spectrum.

This indicates that the aromatic parts of the DOX molecules have a preferential interaction with the 13CD-CO polymer.

Interestingly, DOX release was highly dependent on the DOX payloads. At the highest DOX loading, DOX release was much reduced. For example, 60±4%, 48±3%, and 22±2% DOX was released in 4 h for DOX loaded nanoMOFs with TDL of 10%, 20% and 50%, respectively.

Therefore, not all the drug was released out at high DOX contents in the nanoMOFs. Possibly, DOX self-associates inside the cages causing the incomplete release, as it was previously been shown with another drug, topotecan (ref). What is more, it appears that one additional advantage of CD coatings is to better release DOX out from the cores, in particular, at early phases (< 6h)

and for low DOX payloads (particularly TDL of 10%). In conclusion, CD coating speeds up the loading process and promotes DOX release.

One novelty was the use of ssNMR spectroscopy to characterize the core-shell nano-MOFs.

13

C-labeled CD-citrate oligomers were successfully synthesized and coated on the nanoMOFs outer surface to perform informative 1 H- [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C and [START_REF] Li | Drug-Loaded Lipid-Coated Hybrid Organic-Inorganic " Stealth " Nanoparticles for Cancer Therapy[END_REF] Al- [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C NMR experiments. The ssNMR methodology related to the selective isotope labeling approach proved to be effective for the characterization of this system based on diamagnetic nanoMOF.

In a nutshell, biodegradable CD-CO (with and without fluorescent dye) were successfully synthesized with a facile green method, DOX was well incorporated into the pores of nanoMOFs with high drug payload, the external surface of DOX loaded nanoMOFs was successfully engineered with CD-CO. Interestingly, CD coatings have also potential capability to interact with drugs.

Synthesis of CD-PEG-CO

The CD-PEG-CO was synthesized by a two-step procedure: the first step was to click PEG on CD followed by the second step of crosslinking. Every CD molecule was linked with one PEG chain and then, linked with citric acid. The PEG density in the synthesized CD-PEG-CO was around 9% (grafted glucopyrano. Compared to the DEX-ALN-PEG copolymers, CD-PEG-CO have the advantage of biodegradability due to the degradable ester bonds. However, the PEG density in CD-PEG-CO was slightly lower than that in DEX-ALN-PEG (around 13% of units were grafted).

After the successful synthesis of CD-PEG-CO, the surface modification of nanoMOFs was performed using the same "green" strategy (no organic solvents, no surfactant) as in the case of DEX-ALN-PEG. The maximal amount of associated CD-PEG-CO was up to around 30 wt% measured by TGA analysis, similar to that of DEX-ALN-PEG (32 wt%). Moreover, the drug loading capacity remained unaffected after surface modification. For example, the DOX payloads of the CD-PEG-CO-coated or the bare nanoMOFs were 40 ± 3 wt% and 41 ± 2 wt% respectively, after 6 days impregnation. One other novelty in this study was the use of highresolution STEM to characterize the bare or coated nanoMOFs. The images clearly revealed the crystalline structures before and after drug loading. Remarkably, the incorporated DOX didn't release out at physiological pH 7.4, whereas it released out immediately in artificial lysosomal fluid (ALF). Confocal microscopy investigations were in good agreement with these findings: when nanoMOFs were put in contact with cancer cells (Hela cell line), DOX was not released in the extracellular medium, but only inside the cells. Noteworthy, the engineered nanoMOFs penetrated inside the cells together with their coatings, and then degraded intracellularly, releasing both their coatings and their active dug cargo, which in turn eradicated the cancer cells. However, despite their biodegradable character, the CD-CO and CD-CO-PEG coatings were less effective to reduce the macrophage uptake than the DEX-ALN-PEG ones, presumably because of lower PEG contents or unfavorable PEG orientation to form a protective "brush" at the nanoMOFs' surface.

In a nutshell, "click" chemistry enabled successfully to produce both DEX-ALN-PEG and CD-PEG-CO copolymers. It is a well-known powerful tool, able to conjugate functional groups to DEX backbone or to CDs. Despite their different synthesis procedures, both DEX-ALN-PEG and CD-PEG-CO adsorbed spontaneously and efficiently onto the external surface of nanoMOFs, enabling reaching high coating amounts, up to around 30 wt%. The presence of the coating materials didn't significant influence the drug loading. In the case of DEX-ALN-PEG with highest PEG density, the best "stealth" effect was observed, possibly due to the stronger anchoring interaction between ALN and Fe sites. However, the DEX-ALN-PEG is a bulky linear molecule, prone to induce bridging effects in between the nanoMOFs. This explains why large amounts (weight ratio between nanoMOFs and the coating material of 1:1) of DEX-ALN-PEG copolymer was required for the surface modification. In the case of lower molecular weight CD-PEG-CO, the bridging effect was negligible. Indeed, regardless of the weight ratio between nanoMOFs and CD-PEG-CO, no aggregation was observed after surface modification.

By applying the "camouflaging" technique on nanoMOFs, the macrophage uptake of the nanomaterials was significantly reduced, indicating the success in escaping the recognition by these immune cells. . As perspective, ligands for active targeting could be further conjugated to the PEG chains. The resulting copolymers could be similarly adsorbed onto the nanoMOFs endowing them with active targeting abilities. The potential targeting ligands include small molecules, peptides, antibodies, engineered proteins, or nucleic acid aptamers. More specific examples are the antibody Trastuzumab for HER2 receptor and folic acid.. After conjugating with the active targeting ligand, the cancer cellular uptake is expected to further increase by receptor-mediated endocytosis, contributing to increased drug accumulation in cancer cells.

This mechanism relies on the interaction between tumor ligands conjugated on the surface of nanoMOFs and the overexpressed receptors/antigens on cancer cell surfaces.

It is recognized that targeting drug delivery systems exhibit several advantages over the conventional chemotherapeutic drugs, particularly in regard to decrease of side effects and enhancement of drug efficacy and safety. Nevertheless, one main disadvantage of many active targeting drug delivery system lies in the expensive and sophisticated technology required for their production. Coatings on nanoMOFs by simple incubation with preformed functionalized copolymers could offer simple and reproducible strategies of interest to be explored in the future. -Therefore, the selection of the targeting ligand will be of upmost importance for the successful coating. Besides, it will be interesting to further explore drug incorporation into functionalized core-shell NPs, as well as shell stabilities in biological media. The strategies presented here could be employed to form coating based on other materials, such as hyaluronic acid, or chitosan. Last, the coating strategies could be applied to other types of biodegradable MOFs, such as UIO-66 and ZIF-8.

The new biomaterials engineered here could find many other biomedical applications in preparing drug nanocarriers, implants or gels to deliver drugs in a controlled manner. The versatile family of biomaterials synthesized here open new fascinating fields of investigation and their in vitro/in vivo toxicity and (bio)degradation is worthwhile to be further deeply investigated.

Reference:

investigations were in good agreement with these findings: when nanoMOFs were put in contact with cancer cells (Hela cell line), DOX was not released in the extracellular medium, but only inside the cells. Noteworthy, the engineered nanoMOFs penetrated inside the cells together with their coatings, and then degraded intracellularly, releasing both their coatings and their active dug cargo, which in turn eradicated the cancer cells.

However, despite their biodegradable character, the CD-CO and CD-CO-PEG coatings were less effective to reduce the macrophage uptake than the DEX-ALN-PEG ones, presumably because of lower interactions with the iron trimesate cores.

In a nutshell, these studies reveal the potential of coatings based on carbohydrate copolymers to produce core-shell nanoMOFs for biomedical applications, paving the way for applications particularly in the field of cancer treatment. Further studies will deal with ligand grafting onto the shells for targeting cancer cells.

  Abbreviations ............................................................................................................................. General Introduction .................................................................................................................. Chapter 1.1: Introduction ..................................................................................................... Chapter 1.2: Carbohydrates in metal organic frameworks: Supramolecular assembly and surface modification for biomedical applications ................................................................ Chapter 2 .................................................................................................................................. Comb-like dextran copolymers: A versatile strategy to coat highly porous ............................ MOF nanoparticles with a PEG shell ....................................................................................... Chapter 3 multifunctional core-shell drug loaded nanoparticles for biomedical applications Chapter 3.1 ........................................................................................................................... Self-assembled multifunctional core-shell highly porous metal-organic framework nanoparticles ......................................................................................................................... Chapter 3.2 ......................................................................................................................... Doxorubicin-loaded metal-organic frameworks nanoparticles with engineered cyclodextrin coatings: insights on drug location by solid state NMR spectroscopy ............................... Chapter 3.3 ......................................................................................................................... Porous nanoparticles with engineered shells release their drug cargo in cancer cells ....... General Discussion and Perspectives ..................................................................................... General Conclusion ................................................................................................................ Annex I: Drug-Loaded Lipid-Coated Hybrid Organic-Inorganic "Stealth" Nanoparticles for Cancer Therapy ...................................................................................................................... Annex 2: Highly Porous Hybrid Metal-Organic Nanoparticles Loaded with Gemcitabine Monophosphate: a Multimodal Approach to Improve Chemo-and Radiotherapy ................ Annex 3: Design of Engineered Cyclodextrin Derivatives for Spontaneous Coating of Highly Porous Metal-Organic Framework Nanoparticles in Aqueous Media ...................................

Chapter 1 . 2 :

 12 Carbohydrates in metal organic frameworks:

8 Figure S1. 1 H 3 Figure S2. 1 H 4 Figure S3. 1 H 5 Figure S6. 1 H 8 II. 13 C-NMR spectra for compounds 3- 5 Figure 5 III. 31 P-NMR spectra for compounds 3 and 6- 8 Figure

 8131415181355318 Figure S1. 1 H NMR spectrum (600 MHz, D2O, 25 ºC) for compound 3

  𝑘 =(40000 -18.02) 162.14 = 246.59 where 18.02 Da is the molecular weight of the ending groups OH and H in the polymer and 162.14 Da is the molecular weight of the glucose residue (C6H10O5). Thus, the molecular formula of DT40 is:(𝐶 6 𝐻 10 𝑂 5 )𝑘 + 𝑂𝐻 + 𝐻 = 𝐶 (6𝑘) 𝐻 (10𝑘+2) 𝑂 (5𝑘+1)

  c1 = 12.011 h1 = 1.008 o1 = 15.999 n1 = 14.007 p1 = 30.974 na1 = 22.99 cu1 = 63.546 li1 = 6.94 n = number of non-reacted propargyl carbamate branches m = number of hydration water molecules x = number of alendronate branches z = number of poly(ethylene glycol) branches na = number of Na cations per alendronate branch cu = number of Cu cations per alendronate branch li = number of Li cations per alendronate branch and where in addition %C + %H + %O + %N + %P + %Na + %Cu + %Li = 100 Such equation system was implemented within Wolfram Mathematica v11.0 software as follows (where carbono = %C, hidrogeno = %H, oxigeno = %O, nitrogeno = %N, fosforo = %P, cobre = %Cu, sodio = %Na and litio = %Li; values indicated for these variables in the following program correspond to those obtained for DEX-ALN50-PEG50 7 derivative):

V

  Figure S14 SEC chromatograms of DEX-PC 5 (A), DEX-ALN50-PEG50 7 (B), and DEX (C)

Figure S1

 S1 Figure S1 FTIR spectra of citric acid (purple), γ-CD (orange), Cy5-γ-CD-citrate oligomers

Figure

  Figure S9 MALDI-TOF-MS spectrum (DCTB as matrix) for derivative 2.

  , are one of the latest classes of ordered crystalline hybrid solids. They attract growing interest because of their versatile architectures and compositions. They are tuned based on the nearly unlimited possible coordination of almost any metal and a variety of organic ligands (carboxylates, phosphonates, sulfonates or imidazolates, etc.), leading to thousands of MOFs with various pore sizes and shapes (He et al., 2021; Horcajada et al., 2012; Simon-Yarza et al., 2018). Iron(III) polycarboxylates-based nanoscale MOFs (nanoMOFs), especially MIL-100 (Fe) (MIL stands for Materials from Institut Lavoisier) are among the most studied for biomedical applications as they were proven to be biodegradable and non-toxic (Agostoni et al., 2013; Baati et al., 2013; Quijia et al., 2021); (Simon-Yarza et al., 2016, 2017). They were shown to incorporate high payloads of hydrophilic, amphiphilic and hydrophobic drugs including anticancer, antiviral and antibiotics. (Agostoni et al., 2013; Horcajada et al., 2012; Li et al., 2020b, 2019a; Rodriguez-Ruiz et al., 2015) ; (Anand et al., 2014; Chalati et al., 2011; Di Nunzio et al., 2014; Rezaei et al., 2018; Singco et al., 2016). However, as for other type of nanoparticles (NPs), the in vivo fate of nanoMOFs in the living body (biodistribution, pharmacokinetics and targeting abilities)strongly depends upon their surface physicochemical properties. Bare nanoMOFs can be easily taken up by macrophages and eliminated from the blood circulation, preventing them from reaching the tumor sites(Baati et al., 2013; Simon-yarza et al., 2016). Therefore, it is of outmost importance to functionalize the surface of the nanoMOFs with shells able to confer them "stealth" properties for biomedical applications.

  al., 2009; Jong Oh Kim, Alexander V. Kabanov, 2009; Wang et al., 2017). As known, when administered as a free molecule, DOX lacks tumortargeting ability resulting in poor biodistribution and low therapeutic efficacy as well as severe side effects (Lee et al., 2010; Wang et al., 2017). Different drug delivery systems have been used to encapsulate DOX, including FDA-approved liposomal DOX formulations (Pegylated Doxil®, Lipodox® and non-pegylated Myocet®), and pre-clinical delivery platforms such as polymeric NPs, metal NPs, micelles, emulsions, gels, DNA intercalated NPs, targeted NPs etc (Zhao et al., 2018). More recently, nanoMOFs emerged as a new type of NPs for DOX encapsulation due to their high drug loading capacity (Anand et al., 2014; Bhattacharjee et al., 2020; Di Nunzio et al., 2014; Giménez-Marqués et al., 2018; Qiu et al., 2020; Zhu et al., 2016). For example, DOX loading in nanoMOF gels made of MIL-100(Al) (MIL stands for Material of Institute Lavoisier) reached up to 62 wt%. Moreover, a pH-triggered DOX release was reported, with less than 10% DOX released out in 100 h in phosphate buffer saline (PBS) at pH 7.4 and more than 90% DOX release in 50 h in PBS at pH 5.5. The biocompatible iron trimesate MIL-100 (Fe) nanoMOFs were further investigated as drug carriers for DOX and maximal reported payloads reached 65 wt% (Qiu et al., 2020). A pH responsive release was reported by Xue et al (Xue et al., 2019), with 66% release at pH 5.5

and 1 .

 1 52 nm correspond well to the 220 (2.55 nm) and 422 (1.48 nm) crystalline planes of the cubic structure of the MIL100 Fe (a: 0.7234 nm) (Agostoni et al., 2013; Horcajada et al., 2007).

Figure 1

 1 Figure 1 TEM images of nanoMOFs, before (A) and after (B) surface modification with 29

Figure 2

 2 Figure 2 STEM images of nanoMOFs: empty nanoMOFs (A); MOFs loaded with 18 wt% DOX

Figure 3

 3 Figure 3 DOX loading and surface charge of DOX loaded nanoMOFs, coated or not with CD-

(

  Xue et al., 2019) but in these studies, only a one day impregnation time was investigated. The DOX loading in other type of NPs was much lower. For instance, DOX payload in the commercial Doxil ® (DOX-loaded liposomes)(Barenholz, 2012) and poly-lactic-co-glycolic acid (PLGA) (Pieper and Langer, 2017) was less than 15 wt%. Altogether, these data highlight the affinity of DOX for MIL-100(Fe) nanoMOFs.

Figure 4

 4 Figure 4 STEM-EDX images of DOX-loaded nanoMOFs (18 wt%) coated by CD-PEG-CO (29 wt%). Mapping of the constitutive C, Fe, O, Cl and N elements. Si signal arises from the support.

  These drug release profiles clearly show that the DOX release behavior can be controlled by the amount of entrapped drug molecules and by the composition of the release media. Such a behavior is of interest in the design of DOX nanocarriers for cancer treatment(Aryal et al., 2009; Jong Oh Kim, Alexander V. Kabanov, 2009;Lee et al., 2010; Wang et al., 2017). Indeed, if administered by intravenous injection, the DOX-loaded nanoMOFs developed here could preserve their DOX payloads in physiological media at pH 7.4, contributing to significantly decrease the DOX side effects in normal tissues, especially heart. In contrast, DOX would be released out of the nanoMOFs once they get internalized in the targeted cells. Indeed, it was shown here that, once the DOX loaded nanoMOFs were internalized in the cells most probably by endocytosis, DOX was rapidly released out and efficiently killed the cancer cells. Moreover, to the best of our knowledge, it is the first time to report that DOX at high loading can be released out (>85%) from nanoMOFs in 24 h.

Figure 5

 5 Figure 5 DOX release in PBS (A); in RPMI (B); in artificial lysosomal fluid (C) : DOX in bare

Figure 6 A

 6 Figure 6 A: A: Cell viability of HeLa cells incubated for 24 h with CD-PEG-CO (light blue),

Figure 7

 7 Figure 7 Confocal images of Hela cells incubated with DOX loaded nanoMOFs coated with

A

  series of CD-PEG-CO water-soluble oligomers were successfully synthesized, reaching molecular weights up to 53 kD and low polydispersities of 1.7. They spontaneously anchored onto the surface of MIL-100(Fe) nanoMOFs in aqueous media forming stable coatings. The surface modification didn't significantly affect the nanoMOFs' DOX loading capacity. DOX loadings reached 64±3%, and the presence of the shell didn't interfere with the nanoMOF incorporation capacity. Remarkably, DOX didn't release out (<1%) at pH 7.4 from the formulation of TDL of 50% DOX loaded nanoMOFs coated or not with CD -PEG-CO oligomer in two days, whereas it rapidly released out in ALF at pH 4.5. Confocal investigations showed that the DOX loaded nanoMOFs penetrated inside cancer cell together with their PEG-based shells, followed by DOX release and diffusion inside the nucleus to eradicate the cancer cells.

Figure S2 1

 1 Figure S2 1 H-NMR spectrum of CD-PEG monomer: The PEG-2000/CD molar ratio in the compound was estimated by integrating the peaks assigned to the 3 protons of the PEG-2000 methyl group (a) divided by 8 anomeric protons (d) of the CD units. The ratio PEG/CD=(3/3)/(8/8)=1,0; meaning that one CD was linked with one PEG. For the part of PEG-2000, the calculated proton amount of the CH2CH2 is 172, which is very close to the theoretical value (44×4=176).

Figure S3 1 H

 1 Figure S31 H-NMR spectrum of CD-PEG-CO. The citric acid/CD-PEG molar ratio in the copolymer was estimated by integrating the peaks assigned to the 4 protons of the citric acid methylene groups (f) and the three methyl (a) protons of the CD units. The ratio gave the molar ratio: citric acid/CD-PEG of 6.0; considering that the Mw was around 53400 g/mole, this means

Figure

  Figure S4 X-Ray diffraction pattern of empty MIL-100 (Fe) nanoMOFs highlighting their ordered structure.

Figure S5

 S5 Figure S5Colloid stability of DOX-loaded nanoMOFs before and after surface modification. Uncoated nanoMOFs (blue), DOX in bare nanoMOFs (purple), and nanoMOFs coated with CD-PEG-CO (green) or CD-CO (red) at DOX TDL of 50% (dashed line ----), TDL of 20% (straight line), and TDL of 10% (dashed line).

Figure S6 Figure S7

 S6S7 Figure S6 DOX payload in nanoMOFs coated with CD-CO at TDL of 100% using different loading methods (A: one pot loading; B: loading first then coating; C: coating first and then loading; D: bare nanoMOFs without coating).

Figure S8

 S8 Figure S8 NanoMOFs internalization in cells quantified by ICP-MS: nanoMOFs coated with 29 wt% CD-PEG-CO (green), nanoMOFs coated with 42 wt% CD-CO (red), and uncoated nanoMOFs (Fe) (blue).

  It was found that after overnight impregnation, the amounts of DEX-ALN75-PEG25, DEX-ALN50-PEG50, and DEX-ALN25-PEG75 associated onto the surface of nanoMOFs reached 29 ± 2 wt%, 27 ± 2 wt% and 32 ± 3 wt%, respectively. These amounts were determined by ICP-MS by direct quantification of the P content in the coated nanoMOFs. The nanoMOFs coated or not with DEX-ALN-PEG were characterized by a series of complementary techniques. XRPD studies demonstrated that the coating materials based on DEX-ALN-PEG polymers did not disturb the crystalline structure of nanoMOFs, despite the high amounts of DEX-ALN-PEG associated to the NPs. Besides, TEM experiments indicated that the facetted morphology of the nanoMOFs was not affected after surface modification. Uncoated nanoMOFs aggregated during storage in water, whereas on the contrary, the colloid stability of nanoMOFs was greatly improved after coating with all the DEX-ALN-PEG samples (mass ratio DEX-ALN-PEG:nanoMOFs 1:1), illustrating that their coating effectively prevent nanoMOFs from aggregation.

  Similar to the DEX-ALN-PEG coated nanoMOFs, the CD-CO coated nanoMOFs was performed by easily impregnating the nanoMOFs in aqueous solutions of the synthesized CD-CO. The facetted morphology of nanoMOFs visualized by TEM and the XRPD crystalline structure were preserved after surface functionalization10 . The amount of γ-CD-citrate oligomers anchoring on the surface of nanoMOFs was determined both by fluorescence spectroscopy and by TGA, an indirect tool to measure the fluorescent dye content in the supernatant after impregnation and a direct quantification of the amount of coating material on the surface of NPs, respectively. The two set of values obtained from these two different techniques were in good agreement with each other. There was 53 ± 9 wt% of CD-CO associated to the nanoMOFs, which was almost twice the amount of the attached DEX-ALN25-PEG75 (32 ± 3 wt%). This effective association could be probably related to the strong coordination of the abundant carboxyl groups for the iron sites located at the external surface of the nanoMOFs. Indeed, citric acid, possessing the same number of carboxyl groups as trimesic acid (the organic linker of nanoMOFs), is well recognized to have a strong affinity to Fe sites.In addition to the conventional techniques, such as TEM, XRD, TGA and DLS, XPS was employed to assess the presence of CD-CO within the nanoMOFs' top layers (~10 nm depth) and to quantify the atomic surface composition. Both "fingerprint" of CD-CO (main contribution at 286.2 eV) and carbon C1s ''fingerprint'' of the carbon skeleton of nanoMOFs (284.8 and 289 eV: C-C or C-OOH, respectively) were well observed for CD-CO coated nanoMOFs, confirming that the surface of nanoMOFs was well shielded with CD-CO.

  ssNMR spectroscopy has been used to characterize the DOX localization in CD-CO coated nanoMOFs,. It emerged as an indispensable analytical technique to obtain a comprehensive atomic-scale characterization of complex porous materials, such as MOFs. ssNMR was a powerful tool to investigate the supramolecular structure of highly porous core-shell drug nanocarriers, namely, CD-phosphate (CD-P) coated MIL-100 nanoMOFs. The first case of reported investigation was carried on paramagnetic MIL-100(Fe) nanoparticle employing 1 H solid-state magic-angle spinning (MAS) NMR analysis. The resulting 1 H NMR spectra showed that cooperative interactions took place between the shell and the core, based on phosphate-iron coordination, improving the colloidal stability of the core-shell nanoMOFs. However, numerous spinning side bands were detected due to the presence of the Fe paramagnetic center, restricting the volume of information that could be drawn from the NMR spectra. In the subchapter 3.2, we synthesized MIL-100(Al) nanoMOFs with hydrodynamic diameter of around 260 nm for NMR characterization.

  5 derivative, containing n propargyl carbamate groups and m hydration water molecules is:

		For k	For the chain	For n propargyl	For m molecules
		glucose	ending groups	carbamate	of H2O
		residues		groups
	C atoms	6k		4n
	H atoms	10k	2	3n	2m
	O atoms	5k	1	n	m

Table S1

 S1 Atomic composition (%) evaluated by XPS for γ-CD-citrate oligomers, uncoated nanoMOFs and nanoMOFs coated with γ-CD-citrate oligomers (50 wt%)

			Chapter 3.2
	Doxorubicin-
		γ-CD-citrate oligomers (%) Uncoated nanoMOF (%) Coated nanoMOF (%)
	Fe2p	0	7	2
	C1s	60	58	58
	O1s	39	34	40

loaded metal-organic frameworks nanoparticles with engineered cyclodextrin coatings: insights on drug location by solid state NMR spectroscopy

Table 1 .

 1 Optimization of the CD-PEG-CO synthesis conditions. Influence of the molar ratio of the reactants and the catalyst on the average weight and number molar masses Mw and Mn, and on the polydispersity Ð and polymerization yield.

		CD-
	Series	PEG:citric acid: catalyst	Total heating time (min)	M n (g/mol)	M w (g/mol)	Ð	Yield (%)
		(molar ratio)
	1	1:5:2	25	7500	13000	1.7	35
	2	1:5:2	30	7600	36800	4.8	30
	3	1:5:2	35	8400	29200	3.5	33
	4	1:5:2	40	11200	52600	4.7	35
	5	1:10:2	25	10000	14300	1.4	35
	6	1:10:4	25	12900	22100	1.7	33
	7	1:15:4	25	19000	33600	1.8	34
	8	1:20:4	25	14700	23000	1.6	30
	9	1:30:4	25	9600	37500	3.9	32
	10	1:10:10	25	13800	24300	1.7	35
	11	1:15:15	25	25300	160000	6.3	32
	12	1:15:4	25	30700	53400	1.7	23
	13	1:20:4	25	24200	35600	1.5	22
	14	1:10:4	25	21300	35100	1.6	21
	15	1:10:10	25	21600	50300	2.3	20
	20 kDa and a 50 kDa cut-off dialysis membrane were used for the dialysis of entries 1-11 and
	12-15, respectively.

Table 2 Experimental

 2

	series	Fe/Cl (wt%)	DOX payload (%)
	1.1	10.6	21.0
	1.2	11.0	19.5
	1.3	11.1	19.1
	1.4	11.5	16.6
	1.5	11.5	16.6
	1.6	11.7	16.2
	1.7	11.7	16.2
	1.8	11.8	16.0
	1.9	14.5	13.6
	Average	11.7±1.2	15.7±1.9
	2.1	10.4	21.3

Fe/Cl weight ratios of a series of DOX-loaded nanoMOFs; 1.1-1.9 are DOX loaded nanoMOFs (18 wt%), 2.1-2.9 are DOX loaded (18 wt%) and CD-PEG-CO (29 wt%) coated nanoMOFs

Table S1

 S1 Composition of the artificial lysosomal fluid. Weight and concentrations of products used to prepare a total volume of 250 mL.Figure S1Elution profiles and molar mass distributions of CD-PEG-CO after purification by dialysis. Samples are named according to Table1: entry 10 (brown), entry 12 (green), entry 13 (dark blue), entry 14 (pink), entry 15 (light blue). Signals are from differential refractometric detector (dRI) : solid line, and light scattering (LS) detector: dotted line.

			Molar Mass vs. volume
	Q-0903-2a	Q-0903-1a	Q-0903-3a	Q-0903-4a	Q-0903-5a
	Compound
	Sodium chloride Sodium hydroxide Molar Mass (g/mol)		Concentration (mg/mL) 3.2 6.0		Weight (mg) 800 1500
	Citrate acid
				21			5250
	Calcium chloride 1000.0			0.10			25
	10.0	12.0	14.0	16.0	18.0	20.0	22.0	24.0
	Disodium			volume (mL)
	phosphate			0.18			24
	Sodium sulfate
			0.039			10
	Magnesium
	chloride			0.11			13
	Glycerin
			0.059			15
	Citrate sodium
	dihydrate		0.077			19
	Sodium tartrate
	dehydrate		0.090			22.5
	Sodium pyruvate
			0.086			21.5

Defining alendronate and poly(ethylene glycol) branches starting from original OH dextran groups instead of propargyl carbamate groups allows to redefine n as free, non-reacted propargyl carbamate groups instead of the total number of propargyl carbamate groups, which simplifies the analysis.

Acknowledgements

We are grateful for help from Ludivine Houel Renault, responsible of the platform CPBM/CLUPS/LUMAT FR2764 for confocal microscopy. We acknowledge Dr. Doru

Abbreviations Chapter 2

Comb-like dextran copolymers: A versatile strategy to coat highly porous MOF nanoparticles with a PEG shell

Supplementary information

Comb-like dextran copolymers: a versatile strategy to coat highly porous MOF nanoparticles with a PEG shell Giovanna Cutrone ,1,a , Jingwen Qiu 2,a , Mario Menendez-Miranda 2 , Juan M. Casas-Solvas 1 , Ahmet Aykaç 1 , Xue Li 2 , Daniel Foulkes 2 , Borja Moreira-Alvarez 3 , Jorge R. Encinar 3 , Catherine Ladavière 4 , Didier Desmaële 5 , Antonio Vargas-Berenguel 1,* and Ruxandra Gref 2,*

Abstract

Highly porous nanoscale metal-organic frameworks (nanoMOFs) attract growing interest as drug nanocarriers. Engineering "stealth" nanoMOFs with poly(ethylene glycol) (PEG) coatings is challenging as the PEG chains readily penetrate inside the nanoMOFs pores. Here we address the goal of coating nanoMOFs with PEG shells by synthesizing biodegradable cyclodextrin (CD)-based oligomers with a bulky structure (larger than the nanoMOFs pores) to avoid their penetration inside the open porosity. The PEG chains were grafted by click chemistry onto the CDs which were further crosslinked by citric acid. In addition, the free citrate units allowed a spontaneous anchoring of the coatings onto the nanoMOFs by complex formation with the iron sites in the top layers. Stable coatings representing around 31 wt% were achieved. The anticancer drug doxorubicin (DOX) was successfully entrapped in the core-shell nanoMOFs with loadings up to 41 wt%. High resolution TEM (HRTEM) showed that the organized crystalline structures were preserved. Remarkably, at the highest loadings, DOX was not released out at pH 7.4, but at pH 4.5 of artificial lysosomal fluid. Confocal microscopy investigations showed that the DOX-loaded nanoMOFs penetrated inside Hela cancer cell together with their PEG shells. Then, they released the drug which further diffused inside the nucleus to eradicate the cancer cells. The optimized copolymer (entry 12) was further used in all the coating studies. Besides, in order to track the localization of nanoMOFs in cells using confocal microscopy, a fluorescent CD-CO-Cy5 copolymer was synthesized to allow stable fluorescent coating onto the nanoMOFs, as previously described (Qiu et al., 2020). Made by crosslinking CD-Cy5 monomers with citric acid, it had an average molar mass of 31050 g/mol, and a polydispersity Ð=1.6.

Surface modification of MIL-100(Fe) nanoMOFs with CD-PEG-CO and their physicochemical characterizations

The porous nanoMOFs were successfully synthesized with a mean hydrodynamic diameter of 255 ± 14 nm and a BET surface area of 1750 ± 45 m 2 g -1 . They exhibited a facetted morphology (Fig. 1A) as previously reported (Agostoni et al., 2013)

General Conclusion

In this study, two different types of carbohydrate polymers were used to functionalize highly porous iron trimesate nanoMOFs MIL-100 (Fe) in view of possible biomedical application.

1. DEX-ALN-PEG copolymers were successfully synthesized by click chemistry with high yields in each step. Both PEG and ALN grafting densities were controlled, and the PEG:ALN molar ratios ranged from 75:25 to 25:75. The colloidal stability of the nanoMOFs was greatly improved after their external surface modification. The copolymers were firmly anchored in aqueous media by a "green" (organic solvent-free) method. More importantly, the DEX-ALN-PEG coatings were able to decrease 3 times the in vitro uptake by macrophages.

2. Biodegradable CD-CO grafted or not with Cy5 fluorophore was successfully synthesized and effectively adsorbed onto the surface of porous nanoMOFs, achieving up to 53 ± 8 wt% of coating material associated with the nanoMOFs. Compared with the bare nanoMOFs, these core-shell nanoMOFs were very stable upon storage. taking advantage of the presence of a fluorophore conjugated to CD-CO, fluorescence lifetime microscopy was employed to decipher the coating mechanism. Taking advantage of the high versatility of CDs' chemical modifications and capacity to form "host-guest" inclusion with a wide range of drug molecules, we successfully encapsulated DOX in the core-shell nanoMOFs.

3. In order to characterize the drug localization in the core-shell nanoMOFs, highresolution ssNMR spectroscopy was employed. [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C-labeled CD-CO was synthesized to perform informative 1 H- [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C and [START_REF] Li | Drug-Loaded Lipid-Coated Hybrid Organic-Inorganic " Stealth " Nanoparticles for Cancer Therapy[END_REF] Al- [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] C NMR experiments. CD- [START_REF] Yang | Metal -Organic Frameworks for Biomedical Applications[END_REF] CO coated MIL-100(Al) nanoMOFs loaded with DOX were successfully synthesized. A set of NMR data clearly confirmed the high association of DOX with both the nanoMOF and the CD-CO moieties. CD coatings were found to be promising in accelerating the loading process and in promoting DOX release. The ssNMR methodology related to the selective isotope labeling strategy was proven to be effective for the characterization of this system based on diamagnetic nanoMOF.

Supplementary Material

Drug-loaded lipid-coated hybrid organic-inorganic "stealth" nanoparticles for cancer therapy