Dr Teresa Simon-Yarza

Prof. P Angelika Mielcarek

Patrick Couvreur
email: patrick.couvreur@u-psud.fr

Christian Serre
email: christian.serre@ens.fr

Couvreur

Nanoparticles of Metal-Organic Frameworks: On the Road to In Vivo Efficacy in Biomedicine

Keywords: drug delivery, in vivo evaluation, metal-organic frameworks, nanomedicine, surface modification

Simona, Catherine, Assia, for creating a pleasant atmosphere at work including the scientific discussions and social activities after work. Special thanks to Shan for being a great officemate, creating a friendly and calm atmosphere, and for all the help (especially with Origin). Thank you all for the time we spent together both at work and outside when eating

First of all, I would like to sincerely thank the French government for supporting my PhD thesis with financial grant and my two supervisors Dr. Christian Serre and Prof. Patrick Couvreur for creating this excellent chance for me for self-improvement, for being able to work with experts and creating a wonderful environment in both teams. I appreciate your guidance during my thesis, as well as your help in editing this manuscript. I am very grateful for all the time we spent together on scientific discussions, analyzing results, and for your patience and support during more difficult times.

I wish to express my appreciation to all the members of the thesis Committee, Professor Laure Catala, Professor Karine Andrieux, and Dr. Cédric Boissiere for the acceptance to participate in the evaluation of this thesis.

I am grateful to all of those with whom I had the pleasure to work during this thesis. Special thanks go to i) Dr. Farid Nouar for his help for the scale up of MIL-100(Fe) and collecting SEM images, ii) Dr. Mathilde Lepoitevin for collecting TEM and SEM images, iii) Dr.

Catherine Cailleau for the animal experiments training and iv) Dr. Ghozlene Mekhloufi and Assia Hessani for controlling the safety related to the experiments with radioactive elements. I would also like to thank Prof. Ruxandra Gref and Dr. Xue Li for the collaboration dealing with coating β-cyclodextrin polyphosphate and β-cyclodextrin polyphosphate with adamantine-PEG and for providing materials for experiments. I wish to express my gratitude to all CeleScreen team members: Dr. Luc Lenglet, Prof. Philippe Manivet, Dr. Lisa Martino, Dr. Song-Hua Lee, and Dr. Lucile Fétiveau for the collaboration using nanoMOFs for toxicity tests with worms. I am very grateful to Prof. Nathalie Steunou and Dr. Saad Sene for the collaboration for the coreshell MIL-100(Fe)@γFe 2 O 3 project. I would finally like to thank all, past and present, students and permanent researchers from IMAP: Antoine, Farid, Vanessa, Mathilde, Georges, Nathalie, Bernard, Sujing, Saad, Florian, Oleksii, Iurii, Shan, Heng, Joanna, Lin, Stellina, Chrisa, Nello, Sara, Megane, Raquel, Anna, Amy, Andrea, ChenChen, BingBing, Vincent, Abhinav, and from Institut Galien: Anna, Vincent, Cancer is one of the biggest issues in 21 st century. Scientists still need to improve current available cancer therapies to overcome side effects due to high dosage of anticancer drugs during chemotherapy. One way is to use porous nanoparticles (NPs) as drug nanocarriers. NPs need to fulfill several important criteria such as : low-toxicity, biodegradability, high drug entrapment and controlled release and a particle size usually below 150 nm. In addition, NPs need a long circulation in the blood and reduced recognition by macrophages in order to benefit from the enhanced permeability and retention (EPR) effect leading to a strong accumulation in the tumor.

However nanoparticles in blood are easily covered with proteins through the so-called "opsonization" process, which corresponds to the formation of a "protein corona" at the outer surface of the nanoparticles. This leads to an easy macrophage recognition and removal of the nanoparticles from the bloodstream by the reticulo-endothelial system (RES) organs such as liver, spleen and bone marrow. To overcome this unwanted effect in view of cancer therapy, there are few strategies to obtain a prolonged circulation of NPs in the blood, to render them "stealth". A key parameter is the particle size that shall be ideally smaller than 150 nm; in addition, a highly negative or positive surface charge is desired to avoid aggregation into the vascular compartment after intravenous administration. This is, however, not sufficient and a surface modification of the NPs with molecules known to improve shielding effect towards proteins from the blood is also needed. One prominent class of biomolecules known to render NPs stealth are polymers such as Polyethyleneglycol (PEG). Among the various classes of nanocarriers, Metal-Organic Frameworks nanoparticles (nanoMOFs), that possess a porous crystalline hybrid network, have gained an increasing interest in the past decade for biomedical applications. A particularly attractive MOF is the mesoporous iron (III) trimesate denoted MIL-100(Fe) (MIL stands for Materials from Institut Lavoisier). NPs from this MOF can load high amounts of challenging drugs with, in most cases, a controlled release in physiological conditions. However, MIL-100(Fe) NPs tend to aggregate in physiological media to accumulate in lungs, liver and spleen after intravenous administration in mice. This passive lung targeting of NPs was previously considered successfully for experimental lung metastasis treatment. However, when aiming to treat other types of cancer, such biodistribution may be considered as an important drawback. This thesis, that addresses several and multidisciplinary aspects in nanochemistry and pharmaceutical science, including in vivo animal experiments, was focused mainly on the way to control the in vivo fate of MOFs NPs in order to reduce lung, liver and spleen accumulation after intravenous administration, as well as to prolong blood circulation. To achieve this, we first selected various surface modifications methods, already reported by some of us, that had proven previously to have an improved behavior in terms of colloidal stability and/or in vitro reduction of macrophage recognition. First, we synthetized nanoparticles and coated them with various biomolecules. We first checked the stability of coated NPs in physiological media that mimic in vivo conditions (rat serum), using mainly Dynamic Light Scattering (DLS) analysis. One of the biggest parts of this work was then to evaluate the impact of the nanoMOFs surface modification over the in vivo biodistribution and pharmacokinetics after intravenous administration in mice.

To improve the results, another part of the thesis was focused on the optimization of NPs synthesis to obtain smaller sizes, likely more suitable for intravenous administration due to their potential ability, once coated, to escape macrophage recognition. The last part of the thesis was focused on the encapsulation within NPs of the anticancer drug Methotrexate (MTX) with appreciation of the drug release kinetic. These results were further used for two different potential applications: encapsulation in core-shell magnetic nanoMOFs-iron oxide NPs and the toxicity evaluation of molecules in C. elegans worms.

In details, this PhD manuscript is divided into 3 chapters:

Chapter 1

In this first chapter we describe the state of the art of nanocarriers with a particular emphasis on Metal-Organic Frameworks NPs. First, we have analyzed the different available strategies to achieve "stealth" particles and their corresponding advantages and disadvantages. The next subchapter is based mainly on a review article that we published during this thesis entitled "Nanoparticles of Metal-Organic Frameworks: on the road to in vivo efficacy in biomedicine", reported in Advanced Materials in 2018. To give an updated view of the literature, I have completed it with more recent results. This part was focused on the use of Metal-Organic Frameworks nanoparticles in biomedicine, describing the state of the art in the synthesis of nanoMOFs, the various surface modifications strategies, the stability in physiological media, toxicity, biodistribution and pharmacokinetics, as well as first achievements in theranostics.

Chapter 2

In chapter 2, we have selected MIL-100(Fe) NPs due to their several key advantages such as: (a) high drug entrapment for a large variety of drug molecules, (b) in most cases a controlled drug release, (c) biodegradability and (d) low toxicity. We then coated the surface of these NPs through different surface modification routes to overcome important limitations of bare NPs such as: (a) low colloidal stability and aggregation in physiological media, (b) fast macrophage recognition, and (c) accumulation in the lungs, spleen and liver after intravenous injections (mice). The main goal was to determine which coating could prevent NPs from aggregation in physiological media and allow a prolonged blood circulation to avoid accumulation in lungs, spleen and liver. The selected known surface modification methods, that have previously shown good in vitro results, are the following: (a) polyethylene glycol though the radical polymerization so-called "Graftfast", (b) covalent grafting of polyethylene glycol through the NHS route, (c) covalent grafting through the use of -cyclodextrin-phosphates or (d) -cyclodextrin-phosphate adamantane polyethylene glycol molecules. For the in vivo experiments Gemcitabine monophosphate was encapsulated into NPs. Both trimesic acid and Gemcitabine monophosphate were radiolabeled to allow their quantitative detection after administration in mice. In vivo pharmacokinetics and biodistribution were then studied, after intravenous injections of different formulations via the tail vein of mice. In addition, to overcome some limitations observed during this in vivo study, part of chapter 2 was focused on the synthesis of smaller MIL-100(Fe) NPs, to ensure better chances of success for this application.

Chapter 3

In the final part, we focused on the encapsulation of the anticancer drug Methotrexate (MTX), into MIL-100(Fe) NPs for two types of potential applications. We first optimized the MTX loading and studied its release profile in different media. The application of MTX encapsulation in these NPs was established using core-shell magnetic nanoMOFs-USPIO (Ultra Small Particles of Iron Oxide), developed in collaboration with prof. Nathalie Steunou (Université de Versailles St Quentin). This study is now being pursued by a new PhD candidate, Heng Zhao. A second potential application was focused on the toxicity evaluation of MTX once encapsulated in these nanoMOFs, based on C. elegans worms, which was performed in collaboration with the CeleScreen start up.

CHAPTER 1 Metal-Organic Framework nanoparticles in biomedicine: Literature review Introduction

This chapter is focused on the state of the art of stealth nanoparticles. We will provide selected examples of strategies proposed to date and analyze which are key parameters for a successful stealth effect. In the second sub-chapter we will focus specifically on Metal-Organic Frameworks (MOFs) used in biomedicine. Part of this chapter is based on a mini-review article published at the beginning of the thesis in Advanced Materials entitled "Nanoparticles of Metal-Organic Frameworks: on the road to in vivo efficacy in biomedicine". This review was written together with Dr Teresa Simon-Yarza, Prof. Patrick Couvreur and Dr Christian Serre. We then devoted a particular attention to analyze the existing surface modification routes of nanoMOFs and to understand the relation between the physico-chemical parameters of the nanoparticles (NPs), their stability in physiological media and therefore their in vivo fate. Finally, we have summarized the current situation of nanoMOFs in biomedicine in terms of toxicity, biodistribution and pharmacokinetics, and finally their application to theranostics.

1.

Stealth properties

Introduction

Cancer is one of the biggest problems of the 21 st century, with a high death rate in Europe. In 2013-2015, a death rate of 261.7 per 100 000 inhabitants occurred in Europe (~1.33 total million deaths in 2015). Such a high rate was observed among woman mainly for breast cancer (95 400 deaths in 2015; 7.2% of all deaths from cancer) while among men it mainly concerned prostate cancer (75 400 deaths in 2015; 5.7% of all deaths from cancer). 1 This highlights the need to improve current available cancer therapies.

Treatment of the most solid tumors involves chemotherapy, surgery and radiotherapy. However, chemotherapy suffers from several drawbacks. The biggest issue is the use of large doses of drugs, which are efficient to kill cancer cells but also are highly toxic for healthy cells. This results in many side effects for the patients such as fever, nausea, vomiting, neuropathy, trouble with breathing, diarrhea, skin rash, hair loss or fertility problems. In addition, repeated chemotherapy can cause multidrug resistance (MDR) in the tumors 2 , which leads to the need of an increase of the drug dosage, and results in worse side effects. To overcome these drawbacks, many scientists started to use nanoparticles (NPs) as drug nanocarriers. This can strongly reduce the dosage of anticancer drugs, and when combined with good targeting abilities, can highly increase therapeutic efficiency together with reduction of the sides effects. 2 Nanoparticles with mean sizes around 100-200 nm are of interest for an efficient tumor targeting, because of the unique phenomenon observed in tumors, the so-called enhanced permeability and retention (EPR) which leads to an accumulation of the nanoparticles in tumors. Specific anatomical and physiological features of tumors, such as leaky tumor vasculature with pore sizes from 100 to 780 nm as well as their poor lymphatic drainage, usually prevent nanoparticles (macrophages, NPs) from their easy removal from the tumor and enable their accumulation in the tumors. 2,3 In 1986, the first evidence of NPs exhibiting an EPR effect was reported. 4 Matsumura and coworkers reported that nanoparticles, so-called smancs, made from a combination of styrene and maleic acid copolymer coupled to the antitumor protein neocarzinostatin, were accumulated after intravenous (i.v.) injection in ddY Sarcoma tumor-bearing mice more effectively by the tumor than the healthy tissues. As a result, the half-life time of smancs was prolonged 10 times compared to free drug. This study was the starting point for many researchers to further develops new-engineered NPs with EPR effect as drug nanocarriers.

As it was first established 2,4 , NPs need a long time circulation in blood to effectively observe the EPR effect. However, when injected in biological fluids, nanoparticles can easily adsorb at their external surface biomolecules such as complement proteins C3b C4b, and the immunoglobulins IgG and IgA, causing the formation of a "protein corona" or "opsonins", in a process called "opsonization". Protein corona on a surface of NPs initiates the removal of the NPs by the phagocytic cells in the blood (e.g. monocytes, neutrophils) and tissues (e.g. Kupffer cells, dendric cells, macrophages). NPs are then eliminated by the macrophages and transferred to the liver, spleen and the bone marrow (i.e. the so-called reticulo-endothelial system, RES). This process drastically shortens the life time of the nanocarriers into the blood circulation and thus strongly impacts the biodistribution with most of the particles being phagocyted and then stored mainly into the liver and the spleen. 4,5 To develop smart nanoparticles based systems with targeting and delivery abilities for many applications, including drug delivery, cancer treatment, gene therapy, imaging and theranostics, "stealth particles" need therefore to be designed. A coating of the external surface of the nanoparticles with highly hydrophilic and flexible polymers is considered as a typical strategy to prevent protein adsorption and allow therapeutic drug nanocarriers to exhibit a prolonged blood circulation for optimal targeting abilities (Figure 1). 6 The long circulation of NPs in blood depends on their size, shape, surface charge, hydrophobicity, and receptor interaction.

The particle size is considered as the most important parameter to design efficient stealth NPs with EPR effect. NPs designed for tumor targeting should have sizes smaller than the leaky gaps of tumor vasculature. However, when NPs are too small (smaller than 10 nm) these can be cleared from bloodstream through kidneys or from the tumor by extravasation. On the other side, NPs bigger than 200 nm can more easily be removed from the bloodstream due to an easier macrophages recognition and thus lead to their fast accumulation in liver, spleen or bone marrow.

For optimal ligand-receptor endocytosis to the tumor cells, NPs should therefore ideally be around 30-50 nm. 7 The nature of the polymer coating (density, size, shape) is also crucial for macrophage recognition. Usually non-spherical NPs exhibit a reduced phagocytosis with an increased blood circulation times when compared to the spherical NPs. 8,9 Nanorods have been shown to exhibit the highest cellular uptake, followed by nanospheres, nanocylinders and nanocubes. 8,9 Surface charge impacts strongly as well, the stability and fate of NPs in physiological milieu.

NPs with high positively or negatively charged surfaces (≥±25 mV) exhibit more repulsive forces. Strongly repelled NPs have fewer tendencies to aggregate and there are easily redispersed once injected. Moreover, macrophages are known to have a negative surface charge. In this case the use of more negatively charged NPs should be beneficial to promote a stronger repulsion between NPs and macrophages and slow down NPs clearance from the blood. 9

Strategies for stealth effect-surface modifications of NPs

As described before, a properly designed control of the particle size, surface charge and coating of the nanoparticles are essential for an efficient controlled DDS (drug delivery systems) in order to get a stealth effect. In this part, we will give examples of different surface modifications strategies to improve the stealth effect of nanoparticles. Schematic representations of a variety of designs of topical nanocarriers are shown in Figure 2. One of the most frequently coating is polyethylene glycol (PEG), therefore we will focus first on examples of different nanoparticles coated with PEG through different strategies and then describe alternative coatings to PEG in a view of a stealth effect.

. Schematic variation of design of nanoparticles for drug delivery platform. Design of nanoparticles has an important role to achieve stealth properties and targeting abilities. NPs can be designed as a function of their size, shape, stiffness or surface properties. 7

Examples of PEGylated NPs

The majority of the surface stabilization method of NPs focused on nonionic hydrophilic polymers and/or surfactants, especially polyethylene glycol (PEG). 2 In 1977, Abuchowski and colleagues, reported the first successful attempt of prolonged circulation time of NPs in blood after covalent conjugates of PEG (2 and 5 kDa) to bovine liver catalase (studied on mice). 11 Since then, PEG has been widely used in many NPs systems, such as liposomes 2,3,12 , polymeric micelles [13][14][15][16] , polymeric NPs 2,17 and proteins. 9 . Chain length, shape and density 10,14,18 of the PEG molecules on the surface of NPs are important parameters for a successful prevention of protein adsorption and further have impact on macrophage recognition of these NPs. 10,14,18 This relation will be described in detail later in this paragraph, followed by examples of experimental or computational showcases.

Liposomes

Liposomes consist of one or more concentric spheres of lipid bilayers, which are separated by aqueous compartments. For long-time blood circulation, liposomes can be coated with PEG, however it is difficult to prepare stable PEGylated liposomes, with well-defined dense coating, without disturbing integrity of the lipid membrane. This problem comes from the hydrophilic character of PEG. Stealth liposomal formulations often contains 5 mol% of coated PEG.

However, it is not easy to obtain full protection against macrophage uptake. Among the PEGylated NPs, PEGylated liposomes are among the most well-known drug delivery platforms.

To date, Doxil® represents the only FDA approved PEGylated liposomes for the delivery of the anticancer drug doxorubicin (Dox) 12 . Doxil® (knows internationally as Caelyx®) is used for the treatment of patients with metastatic breast cancer, ovarian cancer, AIDS related Kaposi sarcoma or multiple myeloma. 2,3,12 Doxil® has a very long drug life-time of 72 h and circulation half-time of 36 h, therefore the bioavailibity of doxorubicin after one week is enhanced around 90-fold, compared to the free drug. 12

Polymeric Micelles

Polymeric micelles are formed by self-assembly of block copolymers, which consist of hydrophilic and hydrophobic monomer units. Thus, when mixed with a hydrophobic drug, they form 30-50 nm micelles, which are efficient drug delivery systems. These micelles are usually formed at low concentrations. PEGylation of polymeric micelles is done through the use of PEG as the hydrophilic segment. Poly(lactic acid)-b-poly(ethylene glycol) or poly(caprolactone)-bpoly(ethylene glycol) are among the most studied polymeric micelles, because of their core-shell structure, which has been demonstrated to be resistant to plasma protein adsorption. 13,14 Currently, various polymeric micelle formulations are in clinical trials for various cancers treatments such as for instance NK105 based on PEG-P(aspartate) Paclitaxel block copolymer for stomach cancer, NK012 based on PEG-PGlu(SN-38) block copolymer for breast cancer, NC-6300 based on PEG-P(aspartate) Epirubicin block copolymer for breast cancer, stomach cancer and lymphoma or NC-6004 based on PEG-PGlu(cisplatin) block copolymer to treat solid tumors. 2,15,[19][20][21][22] Product consisting on mPEG-PDLLA (poly(ethylene glycol)-block-poly(D,Llactide)) NPs for the delivery of paclitaxel, Genexol-PM®, is already approved by Korean FDA for the treatment of breast cancer and non-small cell lung cancer (NSCLC) and is currently under clinical trials in the US. 16

Polymeric NPs

Polymeric NPs typically consist of hydrophobic and biodegradable polymers. There are two ways for obtaining PEGylated polymeric NPs. One possibility is to use block copolymers of PEG and a hydrophobic polymer. The second strategy consists of grafting PEG on the external surface of pre-formed polymeric NPs. PEG bearing amine terminal groups can be further covalently conjugated to reactive functional groups (e.g., carboxylic groups) on their outer surface. PEG can also be attached to the surface of the NP via avidin-biotin interactions. 2 Drugs are usually located inside the polymer matrix and are released by diffusion and/or through a degradation of the matrix. 2 Brigger and colleagues have reported PEG-coated poly(hexadecylcyanoacrylate) (PEG-PHDCA)

nanospheres of approximately 150 nm with stealth properties as well as tumor brain targeting effect. In vivo biodistribution and pharmacokinetics experiments were performed on Male CD 344 Fischer rats bearing a 9L gliosarcoma cancer model in comparison with a healthy animal control group. The resulting PEGylated nanospheres showed a longer-circulation time in both, healthy and tumor-bearing animals and with high nanoparticle accumulation in tumor. 23 Following these promising results, authors further loaded these engineered nanoparticles with the anticancer drug doxorubicin (DOX). Nevertheless, Doxorubicin-loaded nanospheres demonstrated a lower accumulation in 9L gliosarcoma compared to the unloaded particles that were mostly accumulated in lungs and spleen. Additionally, size and surface charge experiment, mimicking physiological conditions (rat serum), indicated that while unloaded nanospheres remain small after contact with serum, Dox-loaded nanospheres increased drastically its size in the same conditions. 17 These results demonstrated how tricky it is to design nanocarriers with good stealth properties. In this case, drug encapsulation led to a change in the particle size and surface charge, resulting into a negligible benefit in terms of long-circulation effect after PEGylation.

Proteins

Sharma et al. have successfully obtained stealth nanoparticles. Authors focused on 5-Fluorouracil (5-FU), a common anticancer drug for colorectal cancer treatment. This drug was conjugated to recombinant human serum albumin (rHSA), which was further chemically coupled to PEG monoamine. The size of PEGylated nanoparticles was approximately 65 nm, while surface charge was highly negative (~-26 mV). After NPs dispersion in mixture of serum and PBS (20% v/v serum to PBS), a strong decrease in drug release capacity was observed after PEGylation (approximately 30% less) in agreement with a slower diffusion of the drug due to the PEG corona. The coating did not impact the cytotoxicity, demonstrated using HT-29 cells. In vivo pharmacokinetics study was then performed for three groups (5-FU, 5-FU-rHSA-NPs and 5-FU-rHSA-PEG-NPs), which were intravenously injected through the tail vein of the Swiss albino male mice. A strong improvement of the half-life (t 1/2 ) of 5-FU-rHSA-PEG-NPs was observed in comparison with other groups. 9

PEG interactions with proteins

Since grafting PEG is one of the most popular strategy for the surface modifications of NPs, many studies have devoted a particular effort to analyze and understand the relationship between the particle size, PEG length and PEG surface density and the shielding ability for protein adsorption. 10,14,18 The "stealth" dependence of corona-core nanoparticles such as PEG-coated poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA), and poly(ε-caprolactone) (PCL) nanoparticles, was analyzed by Gref et al.. The PEG chain length and surface density were tuned and the impact evaluated by phagocytic uptake with polymorphonuclear (PMN) cells and competitive plasma protein adsorption. In case of PEG-PLA45K, PEG of 5000 g/mol with PEG content between 2 and 5 wt% was reported to be optimal for reduction of opsonization. Higher Mw or PEG density had no influence in terms of protein adsorption. Phagocytosis and zeta potential data were in agreement with results from protein adsorption. Importantly, above a certain Mw, the polymer brush started to fold, which leads to a better protection of nanoparticle surface and a minimization of protein adsorption. 14 Another group, Lin et al. also analyzed the effect of PEG density and PEG chain length on NPs shielding ability against a range of molecules such as proteins. According to coarse-grained molecular dynamics (CG MD) and dynamic light scattering (DLS) measurement NPs with coating densities of >40%, a minimal aggregation in water was deduced. Moreover, authors have calculated shielding abilities against several biologically relevant molecules: water-0.14 nm, butane-0.23 nm (used as the model of small molecules with four carbon chain), a short alpha helix peptide-0.6 nm (representative of small protein molecules), and albumin-3.5 nm (example of large protein). Water molecules were shown to penetrate freely through the PEG corona even at the highest PEG coating density and chain length. Shielding abilities against butane or alpha helices increased rapidly as the density and chain length of PEG increased. In case of shielding against albumin or similar molecules, accessible hydrophobic content drop to 0% when PEG density reaches 40%. It can be concluded that accessibility of proteins through the PEG corona could be minimized if high-density PEG coating is considered. 10 Walkey and colleagues also focused on the relationship between size, surface charge of NPs and their surface modification on serum protein adsorption and macrophage uptake. Authors used gold nanoparticles of different sizes: 15, 30, 60 and 90 nm, with grafted thiolated, methoxy-terminated PEG with densities from 0 to 1.25 PEG/nm 2 with fixed Mw of 5000 g/mol. They demonstrated that for a fixed nanoparticle size, increasing the density of PEG leads to a reduction of the protein adsorption from human serum, almost to a minimum value (Figure 3). A highest reduction in macrophage uptake, performed with J774A.1 murine macrophage-like cells, was reported for nanoparticles grafted with a density of 0.64 PEG/nm 2 . Further increase of PEG density did not lead to a further decrease of macrophage uptake. In addition, it was shown that, in order to reduce macrophage uptake, it is not required to fully eliminate protein adsorption. 18 ) have less adsorption of the proteins, but shielding effect is not enough, thus NPs are still recognized by macrophages. NPs with both higher densities (above 0.64 PEG/nm 2 ) adsorbed much less proteins, resulting in significantly reduction of macrophage uptake. 18 Therefore, to obtain stealth particles, these crucial parameters (size, charge, coating parameters: chain length and density) need to be verified and precisely selected for optimal results.

Particularly, for PEG coating, NPs coated with dense PEG coating have better shielding effect against proteins from the blood. 14 Remarkably, coating density above 40% led to significant reduction in adsorption of the proteins. 18 Limited adsorption of proteins, however, does not induce macrophage uptake. 10,14,18 Therefore, in a nut-shell, the use of a PEG coating for NPs exhibits multiple advantages such as:

(a) an easy coating process, (b) a positive stealth effect, (c) a lack of toxicity, (d) a good stability of the coating, (e) a better stability in physiological media, or (f) shielding against proteins in the blood, if density is high enough. However, PEG coating often suffers also from several drawbacks or limitations such as: (a) no systematic prevention of adsorption of blood proteins, (b) difficulty to determine for each nanocarrier the optimal coating parameters to reach an improved stealth effect or (c) a lower cellular internalization efficiency and/or a modification of binding interactions with the target disease proteins, (d) activation of the complement at the surface of PEGylated nanoparticles, and (e) absence of biodegradation and urinary excretion.

Alternative coatings

PEG is still one of the most extensively studied polymers for the surface modification of nanoparticles in biomedicine. There are to date other methods to coat the surface of nanoparticles to ensure the stealth effects such as the use of other polymers, proteins, lipid bilayers or more complex molecules.

Other polymers

To date, many polymers have been proposed as an alternative to PEG coating. 2,24 Poly(2-ethyl-2oxazoline) with hydrophilic segment polyoxazolines (POZ) can be made with high quality, with different architectures and can be modified with different functional groups. POZ was conjugated with poly(caprolactone), poly(1,3-trimethylene carbonate), or poly(aspartic acid) for polymeric micelles formation. [25][26][27][28] POZylated therapeutics have better bioavailability compared to PEGylated formulations and better stability in physiological media. 29,30 For instance, poly(2methyl-2-oxazoline) show no cytotoxicity at concentration up to 20 g/l in cell culture 31 and up to 2 g/kg dosage after intravenous injection in rats. 29 Moreover, POZ used in liposomes formation showed satisfactory stealth effects. 32 Polyoxazolines have the same beneficial properties as PEG to achieve stealthness of particles, without disadvantageous effect in animal model. 29 Moreover, compared to PEG, it has some special characteristics beneficial for drug delivery. They are nonionic, stable and highly soluble in water and organic solvent. POZ is rapidly excreted by the kidneys with no significant accumulation in the body and does not activate immune response after administration. 29 Another suitable group of polymers is based on betaines (polybetaines), which are zwitterionic molecules that possess an ability to bind water via electrostatic interaction. 33 Polybetaines showed a decreased protein adsorption, formation of biofilms and bacterial adhesion on various surfaces. 3435 Poly(carboxybetaine) has been used to coat silica 36 , gold 37 , iron oxide 38 , hydrogel or PLGA 39 NPs, to provide better stability and shielding effect towards proteins from blood. 2,[36][37][38][39] Surface hydration is one of the key reasons for PEG resistance to nonspecific protein adsorption. While PEG achieves hydration via hydrogen bonding, polybetaines bind water even stronger via electrostatic interactions. Therefore, polybetains showed excellent low fouling properties against complex media (e.g. undiluted blood serum). 3840 Some poly(amino acids) such as poly(hydroxyethyl L-glutamine) and poly(hydroxyethyl-L-asparagine) have also been proposed as an potential stealth polymeric coating. Liposomes coated with poly(hydroxyethyl-L-asparagine) showed prolonged blood circulations. 2,41 In contrary to PEG, poly(amino acids) are rapidly degraded by proteases and thus accumulating in the body is limited, as well as toxicity is reduced. 41 Another well-developed groups of polymers considered for stealth effect are polysaccharides (e.g. derivatives of chitosan [42][43][44] , dextran 45,46 , heparin 47,48 or hyaluronic acid 49 ). 2 Polysaccharides have many advantages, such as biodegradability, low immunogenicity and low toxicity. 42 They also provide multiple functional groups, which can be beneficial for conjugation of drugs or cell-penetrating ligands. 24 Nanoparticles coated with polysaccharides have prolonged half-life circulation time of encapsulated drugs with enhanced accumulation in neoplastic tissues. For instance, chitosan at lower pH become positively charge and can be easier bind to negatively charged surfaces of tumor tissues. Thus, targeting abilities are much improved. 43

Proteins

As already described, proteins in the blood adsorb onto the surface of NPs, which typically leads to a destabilization and aggregation of NPs, leading to their macrophage uptake and removal from the blood. However, if NPs are coated by selected proteins in a well controlled manner, it is possible to design NPs with good shielding abilities against proteins from the blood. 50 For instance, Oh and coworkers coated mesoporous silica nanoparticles (MSN) with a recombinant fusion protein GST-HER2-Afb, in which HER2-binding affibody (Afb) combines with glutathione-S-transferase (GST) (Figure 4). The protein corona shield (PCS) nanoparticles were described to reduce serum protein adsorption, while retaining its targeting ability. PCS

Biomimetic Route

Red blood cell (RBC) membrane has been reported as an alternative stealth coating to NPs. 52,53 Che-Ming and colleagues used erythrocyte membrane to coat polymeric NPs (PLGA NPs). 53 RBC coated NPs had a size around 80 nm. RBC-membrane-camouflaged NPs exhibit longer circulation half-life after i.v. injection in mice, compared to NPs covered with PEG-lipid. Piao and coworkers, coated gold nanocages (AuNCs) with RBC membranes. 52 Authors compared RBC-AuNCs with PVP-AuNCs (NPs were coated with poly(vinylpyrrolidone)). Size of RBC coated NPs was also around 80 nm. The circulation half-time of the RBC-AuNCs was significantly improved compared to PVP-AuNCs. In addition, RBC-AuNCs, combined with NIR irradiation, showed satisfactory results with PTT cancer treatment. However, in both studies 52,53 , NPs coated with RBC did still accumulate in the liver and the spleen. 52,53 Nevertheless, these coating methods may have one significant drawback if one considers clinical translation. This approach may face some challenges related to immunogenicity or disease transmission. In contrary to animal studies, human erythrocytes contain numerous surface antigens that can be classified to many different blood groups. RBC membrane used for stabilization will need to cross-matched to patients' blood, similarly as in the case of blood transfusion. There are many other approaches using endogenous indirect carriers for increasing the stealth properties and improving the tumor translocation. Sobot and coworkers proposed squalene derivative of gemcitabine (SQGem) nanoparticles for the best targeting to cancer cells with high lipoprotein (LP) receptors expression. SQGem exhibited spontaneous strong interaction with plasma LPs.

This concept was proven in vitro in human blood and in vivo in rats. 54 Intravenous administration of liposomal or polymeric nanopharmaceutics incites autonomic, muco-cutaneous and cardiopulmonary reactions in some human patients. Therefore, Wibroe and colleagues, proposed the system, where carboxylated polystyrene spheres (500 nm) were associated with human or pig erythrocytes. These nanoparticles attached to erythrocytes had poor complement activation, and thus they are able to decrease particle-mediated cardiopulmonary distress by avoiding early interactions with macrophages. 55 Huang and coworkers, to successful delivery of drug SN-38 into lymphoma tumors, surface functionalized T cells with SN-38-loaded nanocapsules. T cells had retained homing receptors mirroring lymphoma cells. This concept was proven successfully on mice model. 56 To summarize, the majority of reported surface modification routes of nanoparticles are important aspect for improving the stability of nanoparticles in physiological media and to prolong blood circulation after i.v. injection. There are many parameters to be considered when designing stealth nanoparticles in nanomedicine leading to series of different strategies. Some works have demonstrated the feasibility of getting stealth nanoparticles in vivo, however sometimes losing this effect upon drug encapsulation due to a change of morphology and/or surface charge of the nanoparticles. Therefore, it is very difficult to predict the behavior of nanoparticles in a complicated system such as the human body. Most of the efforts are also focused on in vitro tests (cells) or animals (mice, rats), which cannot be easily translated to humans. Only a few nanocarriers could reach clinical trials on humans and even less were approved by FDA (e.g. Doxil®, Genexol-PM®). Some clinical studies demonstrated that the EPR effect is rather variable from one patient to another, as well as depending on the tumor type, growth and extracellular matrix. The large tumor-to-body weight ratio in mice compared to human patients may also significantly alter the pharmacokinetics and biodistribution of drug nanocarriers. Most of the rodent tumors grow much faster than tumors in human. Thus tumor environment in human is much more complex. For instance, due to fast growth of tumor in mice, blood vessels in mouse tumor are much more leaky. In contrary, in humans, not all tumor vessels are leaky, which causes heterogeneous distribution of pore sizes and in result heterogeneous extravasation and delivery of drug nanocarriers. 57 Human patients generally receive chemotherapy treatments at longer time intervals (2 to 4 weeks), compared to mice (~3 days).

Long intervals between treatments in human patients allow patients to recover from the toxic, side effects of the treatment. Nevertheless, during this time cancer is also able to recover and gain resistance to the treatment. Much more aggressive schedule in mice reduce the chance of tumor recovery and evolution. Moreover, most studies in mice model, due to short lifespan of the mouse, do not focus in long-term effects of the anticancer treatment. 58 Thus, the pre-clinical models are not always very relevant. [57][58][59][60][61] Lee and coworkers treated human patients with HER2positive metastatic breast cancer with 64 Cu-MM-302 NPs ( 64 Cu-labelled HER2-targeted PEGylated liposomal doxorubicin). Variable uptake of NPs occurred both across lesions within a patient and across patients. For instance, in patients with multiple lesions, not all of lesions had the same level of uptake. 62 Harrington and colleagues used 111 In-DTPA-labelled PEGylated liposomes (IDLPL) 17 patients with locally advanced cancer (breast cancer, SCCHN, lung cancer, high-grade glicoma, and cervix cancer). Positive tumor images were observed in 15 of 17 studies, although levels of tumor liposome uptake varied between and within tumor types. 63 Ramanathan and coworkers studied imagine properties of furumoxytol (FMX) iron nanoparticles and nanoliposomal irinotecan (naI-IRI) in tumors, identified by quantitative MRI. All patients had metastasis disease with two lesions. Heterogeneous uptake was observed across patients and across individual lesions within patient. Moreover, significant accumulation in liver and spleen was observed. 64 Subbiah and coworkers used NC-6004, a cisplatin nanoparticles (nanomicelles), and gemcitabine for treatment of advanced solid tumors in human. Tumor shrinkage was observed in 11 of 22 patients, while partial responses in 3 of 22 patients. 65 In the next part, we will focus on Metal-Organic Frameworks nanoparticles (nanoMOFs), in which our group is specialized. Their unique features, such as (a) the porous structure of NPs, able to encapsulate variety of drugs at high content, (b) their easily tunable structure, (c) their controlled particle size distribution, (c) their controlled surface charge or (d) their easily surface modification with different coatings, make nanoMOFs promising candidates for biomedical applications.

Metal-Organic Framework nanoparticles (nanoMOFs)

Introduction

The following sub-chapter, is based on a review article entitled "Nanoparticles of Metal-Organic

Frameworks: on the road to in vivo efficacy in biomedicine", published in Advanced Materials in 2018 during the course of this PhD. To give an updated view of the literature, more recent results have been added. The full article is available in the Annex.

Metal-Organic Frameworks (MOFs) or porous coordination polymers were first reported in the late 80's 66 and this domain started to expand continuously since the end of the 90's [67][68][69] . The structure and composition of these crystalline hybrid solids can easily be tuned through the almost infinite possible combination of metal inorganic sub-units (clusters, chains or layers of transition metals, 3p, lanthanides…) and/or constitutive organic ligands (carboxylates, phosphates, azolates…) leading to thousands of MOFs with unique features [70][71][72][73] . MOFs exhibit therefore highly porous structures that span over a large range of pore sizes (micro-or mesopores) or pore shapes (cages, channels…), and possess either rigid or flexible frameworks.

Examples of various MOFs are showed in Figure 5. Moreover, schematic formation of MOFs is represented with example of MIL-100(Fe) (MIL stands for Materials of Institut Lavoisier) (Figure 5a). Reaction of Fe(III) trimers with trimesic acid led to formation of hybrid supertetrahedra, which further assemble to the highly porous structure of MIL-100(Fe). MIL-100(Fe) consists of big mesoporous cages (29Å), accessible through hexagonal windows (8.6Å)

and small mesoporous cages (25Å), accessible through pentagonal windows (5.6Å). 74 One can also further tune their polar/apolar character through the use of polar or apolar organic functionalities, often carried out through direct synthesis or post-synthesis modification on the ligand or grafted on the metal sites, which strongly impacts the sorption properties of the solids 75 .

As a consequence, a large number of potential applications of MOFs have been proposed to date such as gas adsorption/storage or separation [76][77][78] , catalysis 79,80 , energy [81][82][83][84] , optical properties 85 , sensing [86][87][88] and biomedicine 89,90 among others 73 . Downsizing materials to the nanoscale is also a suitable method to tackle down new applications, compared to the ones of their bulk analogues [91][92][93][94][95][96][97][98][99] , such as in biomedicine, whereas applications of nanoMOFs is a rapidly developing topic of interest. Nanomaterials, due to their smaller particle size, can indeed improve the drug delivery performances for the treatment of several diseases 90,100 . Moreover, the high and regular porosity and the unique combination of well dispersed metal sites and organic groups within the framework of nanoMOFs, combined with the low toxicity of polycarboxylic acids and selected metals (Fe, Zn, Ca…) make these porous solids appealing nanocarriers for the controlled release of drugs 75 . Considering the drug delivery requirements, the best nanovectors are the ones that fulfill the following conditions [101][102][103][104][105][106] : (a)

high drug entrapment (payload and efficiency), (b) controlled drug release without 'burst' effect, (c) ability to target diseased cells and tissues in a highly selective manner, (d) lack of toxicity through a progressive degradation and absence of accumulation in the body, (e) an easy engineering of the outer surface of the nanoparticles (NPs) for an improved in vivo stability and/or biodistribution (the distribution and accumulation in the different organs and tissues) and

(f) possibility of NPs to be detected by imaging techniques.

In the past few years, numerous studies have demonstrated the great potential of nanoMOFs at the pre-clinical level for biomedical applications. Many of them were reported very recently based on their bioactive composition 107 , anticancer application 72,[108][109][110][111][112][113][114][115] , or from a general drug delivery/theranostic perspective 116 . In this section, we aim at providing a global view of the studies that evaluated MOFs biomedical applications at the pre-clinical stage, when in vivo tests are described either for pharmacological applications or for toxicity evaluation (Table 2). Of note, most of the nanoMOFs reaching a pre-clinical in vivo evaluation are those based on Fe carboxylates or Zn azolates, except recently a few metal(IV) polycarboxylates. We first described the current surface engineering approaches that are crucial to understand the in vivo behavior of the nanoMOFs. Finally, after a detailed and comprehensive analysis of the in vivo studies reported with MOFs so far, and considering the general evolution of the drug delivery science, we suggest new directions for future research in the use of nanoMOFs for biomedical applications. 117

Synthesis of nanoMOFs

Different synthesis routes for the synthesis of MOFs nanoparticles have been reported to date based on either microwave irradiation, sonication, emulsion, electrochemistry, thermomechanical approach, pump flow technology or microfluidics. The choice of the preparation route strongly impacts the nanoparticles' parameters, such as particle size and shape, surface charge or crystallinity. Each method exhibits advantages and drawbacks and depending on the type of application, one shall consider one of these routes to prepare nanoMOFs. 5,75,118 When considering bioapplications, especially to prepare NPS for i.v. injections, smaller NPs are preferred, usually with a size below 200 nm. Larger particles can not only lead to embolia through a blockade of the small blood capillaries, but larger particles have a higher tendency for macrophages recognition and elimination 119 . In addition, to avoid any toxicity issues, one shall privileged the use of green solvents while the possibility to scale-up easily on a reproducible manner the nanoMOFs is of importance. NPs synthesized for bioapplications purposes shall also be biodegradable.

In recent years, among different synthesis method for preparation of nanoMOFs, microwave route was among the most popular one. 5,75,118,119 This method requires the use of microwave irradiation in water or an organic polar solvent. This method has specific advantages: (a) short reaction time, (b) good crystallinity of resulting products, (c) homogenous and controllable NPs particle size distribution, possibility to downsize the NPs, (d) a satisfactory reaction yield.

However microwave suffers from a few disadvantages such as (a) specific equipment and (b) difficulty to scale up due to safety issues. 5 The reaction time affects the size and crystallinity of MOF NPs (example MIL-101(Cr)). Short reaction time led to the formation of very small NPs, but with a poor crystallinity, while longer reaction times led to bigger and well crystalline NPs. 120 Another method to obtain nanoMOFs is sonochemical synthesis, which requires the use of acoustic waves. The synthesis time and power can be easily tuned, which affects NPs formation. 121,122 This method exhibits many advantages: (a) homogenous nucleation, (b) fast reaction time, (c) green solvents, (d) monodisperse, controllable size, (e) formation of nano-or microcrystals. 5 Lestari and coworkers, synthesized [Zn 3 (BTC) 2 ] by the sonochemical method.

They demonstrated that longer synthesis times (over 90 minutes) lead to larger particles with however an agglomeration of NPs. 122 One drawback of the sonication route its rather low yield.

The scale-up of this synthesis route is thus questionable.

Recently, Cacho-Bailo and coworkers, reported the use of microfluidics to prepare MOFs. This involves the use of oil-segmented droplets to confine the growth of the nanoMOFs (see Figure 6).

The residence time for the crystallization solutions can be tuned by varying the reactor length.

Longer reaction times in the PTFE tube led to larger, more angular shaped NH 2 -MIL-88B(Fe) particles, as well as a broader particle size distribution. Furthermore, when lower temperatures were considered (55 and 75 vs 95°C), this led to the formation of less polydispersed nanoparticles of a smaller size. 123 However, it is questionable whether this method is scalable. To date, to our opinion, there is no perfect synthesis route (biofriendly route without toxic organic solvent, controlled small particle size with narrow size distribution, or scalable method) of nanoMOFs for biomedical applications and thus there is still a strong synthetic effort by MOF teams to develop better routes to prepare nanoMOFs for this type of application.

NanoMOFs in biomedicine

Surface modifications of nanoMOFs

So far, the great number of studies describing the potential of MOFs to be used for pharmaceutical applications contrasts with the reduced number of in vivo pharmacological efficacy studies performed to date 5 . There are several reasons why most of the physicochemical and in vitro studies dealing with MOFs in the biomedical field have not been systematically evaluated in vivo yet. Some are related to MOFs design development itself, such as colloidal stability issues and/or lack of organ or tissue targeting properties 5 . These challenges are already being overcome by several surface coating strategies. Similarly to what has already been demonstrated for other nanocarriers, it is possible to tune pharmacokinetics and biodistribution by incorporating motifs onto the outer surface of the nanoMOFs, and the aim of this section is to give an overview of the latest advances in this field (the most recent works are summarized in Table 2).

First, it must be noted that for isotonicity, intravenous (i.v.) administration of nanomaterials needs dispersion in an isotonic solution such as 0.9 % saline or 5% glucose. Many nanoMOFs, particularly once dried, cannot be dispersed yet in these media and/or are not stable enough and tend rapidly to form aggregates, making i.v. administration not possible due to risk of embolization. In some cases, nanoMOFs could nevertheless be administered at (low) doses at which particle aggregation does not occur. For instance, Baati et al. did not observe any stability problems for low doses of iron (III) carboxylates MIL-100(Fe), MIL-88A(Fe) and MIL-88B_4CH 3 (Fe) when dispersed in glucose solution 124 . But colloidal stability of nanoMOFs still remains a major issue that makes difficult in vivo administration 125 . Selective functionalization of the external surface of nanoMOFs is thus required in view of their biomedical applications.

Apart from the colloidal stability, one major objective of the surface engineering of nanoMOFs is also to allow blood long circulating properties (i.e. "stealthness") of the NPs (Figure 7). Indeed, surface functionalization of NPs with hydrophilic polymers could reduce NP opsonization and recognition and elimination by macrophages to the liver, spleen and the bone marrow. 5 Surface modification strategies of NPs and liposomes were described in previous subchapter. Among different surface modyfications, polyethylene glycol (PEG) is the most studied coating. The PEGylation of polymeric NPs was first proposed in 1994 by Gref and colleagues 126 and a great number of studies dealing with PEGylation of other types of NPs has been carried out since that time 127,128 . Nevertheless, to date, Doxil® represents the only FDA approved PEGylated liposomes for the delivery of the anticancer drug doxorubicin (Dox) 12 .

Besides, "active" targeting implies the incorporation of molecules on the surface of the nanoMOFs that are specifically designed to recognize particular receptors or antigens expressed onto the membrane of the diseased cells and tissues 5,129,130 .

Figure 7.

Surface modification of nanoMOFs directly affects colloidal stability before and after in vivo administration, drug release and burst effect, in vivo targeting and prolonged circulation in the organism.

In general, surface coating of NPs should fulfill some criteria to be suitable for drug delivery, such as 103,131 (a) biofriendly synthesis method, without any toxic additive, (b) avoid intrusion of the targeting molecule into the nanocarrier, (c) preserve drug release capacity, (d) lack of interference with entrapped drugs, (e) improvement of particle colloidal stability and (f) stability under physiological conditions. Furthermore, in the case of nanoMOFs, the presence of a high porosity makes the surface modification even more challenging, due to possible non-specific intrusion of the molecules into the pores, blocking them and/or decreasing drug release capacities 131 . Numerous strategies have been proposed so far to modify the surface of nanoMOFs using a large variety of molecules, such as nucleic acids 132,133 , polymers 110,[134][135][136] , cyclodextrins (CDs) 131 or lipids 137,138 , resulting into various degrees of improvement of the particles properties.

One can classify these approaches into three categories: covalent 131,133,[146][147][148][149]134,[139][140][141][142][143][144][145] , noncovalent 129,136,[150][151][152][153][154][155] (Table 1) or core-shell 86,130,163,137,[156][157][158][159][160][161][162] , as detailed below.

Table 1. Covalent and non-covalent strategy for surface modifications of nanoMOFs (CD-P: cyclodextrin-phosphates, Py-PGA-PEG-pyrene derived polyethylene glycol).

Covalent route

Recently, Zimpel et al. have reported the covalent attachment of functional polymers, such as amino-PEG and a derived oligoamino amide named Stp10-C and constituted by two terminal groups, a primary amine and a thiol, connected via a repetitive diaminoethane motif with protonsponge characteristics, grafted on the external surface of the mesoporous biocompatible iron carboxylate MIL-100(Fe) NPs (Figure 8) 134 . This resulted in a successful improvement of the colloidal stability of the functionalized NPs, together with satisfactory cellular uptake of MIL-100(Fe) NPs functionalized with Stp-10C by murine neuroblastoma (N2A) cells. Authors showed nevertheless a negative impact of the polymer coating on MIL-100(Fe) MRI (magnetic resonance imaging) properties. Concerning PEG coating, no information was provided about its influence on macrophages uptake and thus it is not yet possible to conclude about the stealth properties of this system 134 .

Interactions Type of interactions Example

Non-covalent -van der Waals Moreover, authors demonstrated the good stability of CD-P coating in aqueous solution, where three washings did not lead to any CD-P leaching, while three washings with phosphate buffer saline (PBS) led to only 7% CD-P detachment. Similarly, it was shown that less than 10% of the total CD-P coating was released from the nanoMOFs after 24 h of incubation in PBS or in cell culture media. Loading and release of the antiretroviral azidothymidine-triphosphate (AZT-TP) was also not affected by MOF coating, in all cases drug cargo was almost entirely delivered after 24 h, under physiological simulated conditions (PBS, 37°C). Importantly, the coating with CD-P did not cause any significant toxic effects on the cell lines (J774, MCF7 and LP-1). In addition, cellular uptake of MIL-100(Fe) NPs functionalized with mannose-bearing CD-P derivative (CD-P-man) by the human retinoblastoma cell line Y79 exhibited more than twice higher penetration inside cells as compared to uncoated NPs 131 . The benefit of such a covalent coating in vivo remains however still to be demonstrated.

Morris and colleagues, took advantage of the click chemistry route to covalently functionalize the microporous zirconium carboxylate UiO-66-N 3 NPs with oligonucleotides 132 , by a strain promoted click reaction between DNA append with dibenzylcyclooctane and azide-functional UiO-66-N 3 . As a result, colloidal stability in NaCl was meaningfully improved and cellular uptake (HeLa-human cervical cancer cells) was significantly higher, whereas cell viability was not altered.

A more sophisticated system was developed by Wang et NPs 139 . This led to a better stability in PBS and prevented from fast degradation, compared to uncoated particles. Surface modified nanoMOFs presented negligible uptake in α v β 3 integrin negative non-cancerous COS7 (African green monkey kidney) cells, while NPs were internalized in α v β 3 integrin expressing HeLa (human adenocarcinoma) cells. In vivo antitumor efficacy in hepatoma H22 tumor bearing mice was then demonstrated through tumor inhibition effect for coated NPs loaded with Dox and for free Dox, while no side effect was observed in coated NPs.

Nevertheless, a main inconvenient lies here in the lack of demonstration of the effectiveness of the coating since the NPs were locally administered by subcutaneous injection, near the tumor.

Schmitt and coworkers, have reported functionalized and hierarchically structured MOFs embedding magnetic core particles (magMOFs) through a layer-by-layer (LbL) synthesis route 143 .

Core/multishell particles were obtained following several steps: (a) MOF growth of Cu(BA-TPDC) (BA-TPDC-Bis(azidomethyl)-terphenyl-dicarboxylic acid) around a magnetic core, (b) click reaction of blue dye, (c) MOF growth of Cu(TPDC) (TPDC-terphenyldicarboxylic acid), (d) MOF growth of Cu(BA-TPDC) and (e) finally, a click reaction of the red dye was performed.

In a next step, the SURMOF final product was converted into dye-loaded SURGELs (the surface grafted gels) capsules around the magnetic core. The advantage of SURGELs results from the possibility to control the release of the dye molecule, depending on the pH of the environment.

The drug loading of the MOFs inner layers was carried out followed by a coating with a protective polymer layer. The resulting magMOF particles could be converted to magGEL by EDTA treatment. Different kinetics of release were reported as a function of the pH with a maximum of release reached at pH 11 after 2 hours. This product was considered for the oral administration of drugs 143 with however neither in vitro nor in vivo assays to assess the benefit of this strategy. In the same study, the growth of SURMOF films on gold coated mica substrate was described followed by their conversion into SURGELs via click chemistry with arginine-glycineaspartic acid (RGD) to favor cell adhesion via specific interaction with the integrin receptors of the cell membrane. GRD is well known to be able to induce adhesion of osteoblasts onto implanted surfaces and to improve bone formation. The proof of concept has been performed in vitro using osteoblast-like CAL72 cell line. The cells were seeded onto the functionalized SURGEL substrates. Finally, no toxicity was observed until 24 hours and more adherent cells were observed onto functionalized SURGELs. Additionally, microfluidic shear force assay was performed which confirmed that CAL72 cells interacted stronger at the surface of functionalized SURGELs, as compared to the non functionalized counterpart 144 .

Recently, Guo and coworkers proposed a novel application of MOF-based particles for infection treatment, "pathogen-like" particles. Two iron-based MOFs (MIL-100(Fe) and MIL-88A(Fe))

were coated with sulfo-Cys5-NH 2 (fluorescent labeling) and mannosamine via amine-carboxylic acid coupling reaction, to ensure a maximal uptake by alveolar macrophages (AMs), lungs cells involved in infection defense system. Non-functionalized MIL-100(Fe) had higher and faster cellular uptake than MIL-88A(Fe), however after coating with mannose, MIL-88A(Fe) uptake increased significantly, while mannosylation had no effect at internalization of MIL-100(Fe). In Furthermore, a short blood circulation half-life was reported, suggesting that surface PEG density was not efficient to provide sufficient stealth properties to the nanoMOFs.

In a first attempt to prepare MOFs for oral administration, MIL-100(Fe) NPs were coated by the bioadhesive polysaccharide chitosan (CS) 150 . The surface modified NPs exhibited an improved chemical and colloidal stability under oral simulated conditions. In vitro, permeability through a model of intestinal barrier and cytotoxicity were improved, in comparison with the non-coated nanoMOFs. The viability and integrity of the intestinal barrier was investigated using an in vitro model of polarized Caco-2 monolayer cells. Systemic and mucosal immune response were also studied, through complement activation tests and by cytokine profile, resulting in the absence of any complement activation for both coated and uncoated NPs, while cytokine production decreased from one to two orders of magnitude for the CS coated NPs 150,151 .

Another non-covalent approach was reported by Bellido and coworkers, through heparin coating of MIL-100(Fe) NPs 152 . The colloidal stability of these NPs was improved in water and PBS, while reduced cell recognition was also observed in vitro for the heparin coated nanoMOFs using a macrophage cell line (J774.A1) at short time of incubation (up to 4 h), without any complement activation and reactive oxygen species production. Furthermore, surface modification preserved NPs encapsulation capacities, as demonstrated with the active caffeine, while a decrease in the release kinetics compared to uncoated NPs was observed. However, the eventual benefit of these nanoMOFs in terms of pharmacokinetics and biodistribution still needs further demonstration.

Another example of non-covalent approach to decorate MOF NPs was developed recently by Cai and coworkers that successfully developed surface engineered MIL-100(Fe) with hyaluronic acid (HA) and indocyanine green (ICG) with improved colloidal stability and good cellular uptake by cancer cells using the CD44-positive MCF-4 cell line 136 .

Qu and coworkers described the ability of non-covalent functionalized iron carboxylate MIL-101(Fe) with unmethylated cytosine-phosphate-guanine oligonucleotydes (CpG ODNs) (MIL-101(Fe)-CpG nanoconjugates) to enhance the immune response. In addition, T 2 -magnetic response imaging ability was tested in vitro and in vivo 155 . CpG ODNs were adsorbed onto the MOFs by π-π interactions between the CpG ODNs and the terephthalic acid organic ligands.

Regarding cellular experiments, the obtained nanoconjugates did not show any toxicity on RAW264.7 cells. After cell internalization, the nanoconjugates interacted with TLR9 and triggered the secretion of cytokines. Furthermore, MIL-101(Fe)-CpG nanoconjugates, displayed a higher immune response comparatively to CpG ODNs alone, both in vitro and in vivo.

Moreover, T 2 -weighted MR images of the tumor-bearing mice, before and after subcutaneous injection of MIL-101(Fe)-CpG nanoconjugates, showed a strong signal at the site of injection, whereas the tumor position was well visible. Due to T 2 -MRI ability of the nanoconjugates, it was proposed to use these nanoconjugates to track the labeled immune cells and to monitor in vivo the CpG-ODN-based drugs or vaccines 155 .

Recently, Nejadshafiee and coworkers reported magnetic bio-MOF NPs (Fe 3 O 4 @Bio-MOF), composed by zinc as metal ion and curcumin (CUR) as organic linker; they loaded the NPs with 5-Fluoroacil (5-FU) (Figure 12), as chemotherapeutic agent, and coated them with chitosan (CS) conjugated folic acid (FA) through electrostatic interactions. This model could simultaneously deliver CUR and 5-FU to the tumor. Release profiles of 5-FU and CUR were obtained for pH 5.5 and 7.4. 5-FU. At pH 5.5, a first burst release occurred until 10 h, followed by a more progressive and constant release at longer time; at pH 7.4 5-FU was released to a lower extent (~30% less), which is a very positive result as the drug might be released faster and in higher quantitative in cancer cells, which are more acidic. Therefore, cytotoxicity could be focused mainly on tumor site. CUR was released only in small amount (16%), only at pH 5.5.

"Core-shell" strategies

Besides covalent and non-covalent surface modification, another strategy related to the so-called "core-shell" approach, consists in the covering of the nanoMOFs core by a dense coating, identified as the shell of the system.

Rieter and coworkers were the first to consider silica coating to form core-shell systems made of nanoMOFs to avoid a too fast degradation of the NPs in body fluids 86 . Further functionalization of silica-coated NPs was also performed using diopicolinic acid. Authors described a general method to obtain variable thickness of silica shells on the lanthanide dicarboxylate Ln(BDC) 137 . They demonstrated that MIL-100(Fe) and the chromium dicarboxylate MIL-101(Cr) NPs, coated with a lipid bilayer consisting of (1,2-dioleoylsnglycero-3-phosphocholine; DOPC), exhibited a better colloidal stability than the uncoated nanomaterial. Noteworthy, the integrity of the lipid bilayer was confirmed and led to a higher uptake by cancer cells (T24 bladder carcinoma cells), without cytotoxic effect of empty coated

NPs. The same team reported very recently the coating of MOF NPs with exosomes via lipid fusion (Figure 9) 157 . This method ensured the preparation of an exosome delivery system with unprecedented loading efficiency. The resulting exosome-coated NPs showed no burst leakage with an efficient release of their cargo into the cells. 117 130 Wang and colleagues also reported nanoscale polymer-MOF hybrids, where the Zr terepthalate UiO-66 nanoMOF was coated with polyaniline (PAN) (UiO-66@PAN) for phototermal therapy (PTT) 161 . The satisfactory photothermal performance of PAN coated NPs was demonstrated upon laser irradiation. Regarding cellular assays, coated NPs did not induce any cytotoxicity in both murine colon cancer CT26 and human colon cancer HTC116 cell lines, but once irradiated, nearly 70% cells were dead. In vivo tests were carried out with both UiO-66@PAN and NIR irradiation, showed a complete tumor regression after 10 days, when compared to the controls (untreated or NPs alone), proving that UiO-66@PAN is a good candidate for PTT based cancer therapy. 161 Recently, Cai and coworkers proposed novel liquid-solid-solution (LSS) method to construct Fesoc-MOF (or MIL-127(Fe)), which is built from 3,3′,5,5′-azobenzenetetracarboxylic acid (H 4 -ABTC) ligand and oxo-centered iron carboxylate trimers of octahedra building blocks. NPs were coated with polypyrole (PPy) (Figure 10). The photothermal effect of Fe-soc-MOF@PPy was very high. The in vitro biocompatibility of coated NPs was confirmed on L929 cells, while their ability to kill cancer cells was tested on murine breast cancer 4T1 cells. NPs before laser irradiation showed no toxicity, while exposure to 808 nm laser led to an inhibition of the growth of cancer cells. The in vivo anti-tumor efficacy was validated on a tumor-bearing Balb/c mice model, by subcutaneous injections of NPs. NIR irradiation (808 nm laser, 1.0W/cm 2 ) was applied for 5 min 1h after injection. Mice treated with Fe-soc-MOF@PPy+NIR exhibited a much smaller tumor compared to the control groups (PBS, PBS+NIR, Fe-soc-MOF@PPy). In addition, analysis of the pathology demonstrated that no damage on other organs was observed.

Furthermore, biodistribution and metabolism experiments were preformed in healthy Balb/c mice (i.v. injections). NPs were however mostly located in liver and spleen, while a high excretion was observed at the second day after injection. 166

Stability in physiological media

One of the key parameters, which needs to be taken into account when designing stealth particles for in vivo applications is their ability to exhibit a long circulation time in physiological media.

One of the easiest methods to study this is the evaluation of size and surface charge of NPs in various physiological-like media (PBS, serum). It is a preliminary step to assess whether or not these particles exhibits might be useful prior to in vivo experiments. The latter are time consuming and very complex; thus this is still very difficult to predict results of in vivo biodistribution and pharmacokinetics of NPs, even a good colloidal stability of the nanoMOFs is established as well as demonstrated in vitro results. Recently, several studies evaluated the size nanoMOFs in physiological media. 117,119,[167][168][169][170] Other studies are trying to shed some light over the mechanisms involved in the degradation process of nanoMOFs. [171][172][173] In 2017, some of us evaluated the impact of the particle size and surface charge dependence of MIL-100(Fe) NPs at different pH, as well as size dependence of NPs after dispersion in serum at different times. It was demonstrated that NPs tend to aggregate at physiological pH (~7.4). At the same time, the initial highly negative surface charge observed at lower pH goes almost to neutral values at the pH of the blood (7.4). Additionally, the size of NPs in rat serum was measured to mimic conditions after i.v. injections. As expected, a high aggregation of NPs was found already after 30 minutes of incubation of NPs in serum (Figure 11). Notably, after 3 h of incubation in serum, the nanoparticles disassembled and recovered their initial size. This phenomenon was explained through a combination of several parameters such as a degradation of the nanoparticles leading to a partial release of the MOF linker as well as the complexation of iron by phosphates from serum, leading to the progressive formation of an inorganic iron phosphate corona as the degradation product. In 2014, Bellido and coworkers reported the study of the colloidal stability of MIL-100(Fe) NPS in different physiological-like media, including saline media, gastrointestinal and serum-like media. Their particle size and surface charge measurements were analyzed in various conditions and are summarized in Table 3. NPs in presence of large amount in electrolytes (NaCl and CaCl 2 ;

naturally present in serum) have strong tendency to aggregate. NPs in PBS were stable only during a short period of time (10 minutes), while the addition of BSA at short times led to an increase of the size by ~20 nm, indicating the formation of a protein corona and a faster increase of the particle size. In addition, the formation of the protein corona was shown to slow down the degradation process. Moreover, authors studied the stability of NPs under oral simulated conditions: (a) gastric medium, (b) intestinal medium (SIF) and (iii) low-ionic-strength (lis) SIF (SIF diluted 1/25) with addition of pancreatin (1 % w/v) or mucin (5 % w/v). NPs were shown to be stable without enzymes or mucin. Moreover, the particle degradation in gastric and intestinal conditions was similar to that in PBS. Addition of pancreatin did not slow down the degradation process, as was observed in presence of albumin, while addition of mucin led to rapid aggregation of NPs, however without any significant degradation process. 170 Li and colleagues, published in 2017 an important study focused on the in-depth understanding of the degradation process of MOFs drug carriers such as MIL-100(Fe) as micro-or nanoparticles. The authors studied the effect of the size and shape distribution after incubation in water or PBS, using DLS and microscopy techniques (TEM, Raman, Mössbauer) (Figure 12).

NPs that were incubated in PBS for 6 h at 37°C showed a rapid degradation, while releasing a large amount of BTC ligand (34±3 wt%). In more diluted conditions (10 times dilution), a similar degradation occurred after 2 days of incubation in PBS. On the contrary, only a negligible amount of linker was released after incubation in water for 2 days. Degradation in PBS led to a change in shape of the NPs; their shape of edges became more round (Figure 12 B), in agreement with an amorphization of the structure. Interestingly, this did not lead to a change in the particle size distribution. Raman spectroscopy of microparticles of MOFs, was used to follow the degradation process. It starts from the outer surface of the particles with the formation of an amorphous shell, while the core of the MOFs remains intact and crystalline. The in aqueous solution exhibited a particle size around 150 nm and a highly negative surface charge (-30 mV) that did not change after 24 h of incubation in 50% fatal bovine serum at 37°C, which indicates a good stability in physiological conditions. Moreover, incubation of NPs in PBS for 14 days at 25°C also neither led to an increase in size nor a change in charge. Authors concluded that this system was suitable for in vitro and in vivo applications due to their excellent biostability, as well as good storage stability, which they further veryfied. 167 Yang and colleagues, proved the good stability of cypate@MIL-53 NPs, modified with PEG and transferrin (Tf) (CMNPs) in various physiological media. Size of CMNPs (250 nm in aqueous solution) was not affected after incubation for 24 h in PBS, serum or in RPMI-1640 cell medium.

However, these results were demonstrated only by DLS measurements and were not analyzed by PXRD to check the crystallinity of the nanoMOFs. However, the particle size is here still slightly too large to be considered for i.v. injections. 168 Before studying the efficiency of ZIF-8 NPs coated with organosilica on the reduction of tumor (HeLa tumor-bearing nude mice), Ren and coworkers proved that the size and surface charge of NPs were intact upon 5 days in PBS or in DMEM + 10% FBS. 169 Difference compared to previous study 173 is probably due to the dense coating presence on the outer surface of NPs.

Before performing in vitro or in vivo experiments, it is important to understand the size and surface charge development of NPs in media mimicking in vivo experimental conditions (physiological conditions). However, these studies emphasize that some differences occur in different groups in terms of colloidal or chemical stability for the same MOF NPs. This is probably due to differences in experimental conditions (e.g. concentration of NPs, stirring, crystallinity/defects of NPs…) as well as the use of different protocols for analysis and/or the use of different equipments. Thus this is very difficult to compare results from one study to another.

This would be very beneficial if in a near future a protocol for DLS analysis of particle size could be proposed as a standard (ex. concentration of NPs, time of sonication, time and number of scans, data analyzing).

Toxicity: from in vitro to in vivo evaluation

There is to date an increasing number of publications demonstrating the lack of significant in vitro toxicity of nanoMOFs based on 2D cell studies using established cell lines. However, these in vitro tests are far from reproducing the in vivo situation. Thus, it is important to emphasize that the lack of in vitro cytotoxicity does not mean that the nanodevice is safe and biocompatible.

In general, toxicity issues should be tackled down based on a more rational safe-by-design approach when considering nanoMOFs for biomedical applications. Most of authors justify the choice of their nanoMOFs based on their constitutive parts, i.e. through the reported metal and ligand individual toxicity. This strategy can be useful mainly to discard the most toxic nanoMOFs, but it is far from being optimal. The available data about metal toxicity refer mainly to medial oral lethal dose, that is the amount of a drug or other substance that, when orally administered to a group of experimental animals, will kill 50 % of the group in a specified time (LD50), and to human exposure to metals present in the environment (water, air, soils and food). Administration of nanoMOFs in the human being as part of a pharmacological treatment implies active penetration of the metal in the organism. In this case, exposure to metals is different than the previously studied ones and will be determined by the intensity (dose), frequency and duration of the treatment, as well as by the administration route that determines how the metal is distributed, accumulated, metabolized and eliminated. Certainly, LD50 value cannot be extrapolated to other administration routes for which biodistribution will be totally different. Furthermore, when one deals with in vivo human toxicity, there is even less data available for most of the constitutive MOFs ligands, such as terephthalic acid in MIL-101 or UiO-66, whose toxicological properties after i.v. administration have not been thoroughly investigated.

As a consequence, to obtain more information about potential toxicity of the nanoMOFs after administration into humans, one shall focus more on reports dealing with in vitro or ex vivo models that specifically mimic an in vivo situation, rather than relying on simple 2D in vitro toxicity tests using established cell lines (Figure 13) 174 . In many cases the cytotoxicity is only performed in cancer cell lines to demonstrate a potential anticancer effect, and not in healthy cells to assess the toxicity of the carrier 132,175,176 , even if Tamames et al. already demonstrated that toxicity of numerous MOFs is higher for healthy cells than for tumor cells 177 . Besides, when healthy cells are used the choice of the cell type is rarely justified based on therapeutic applications 178 . Some efforts have already been made to design more pertinent cytotoxic in vitro studies by Wuttke et al. that explored the toxicity of various nanoMOFs using different primary healthy cell types depending on the proposed medical application 179 . For their use as drug delivery systems, toxicity and inflammatory response on vascular cells and lung cells were tested, whereas for their application as drug delivery implant coatings of dental or nerve guidance tubes, fibroblasts and neural cells were chosen respectively. Furthermore, alternative models to screen toxicology, such as numerous 3D in vitro models, are currently being developed to better reproduce the in vivo conditions of a particular tissue, such as the tumor environment, the liver or the skin [180][181][182][183] . These new models were also reported by Wuttke et al. where the biological response of sensory neurons to the nanoMOFs was monitored using rat neonatal organotypic dorsal root ganglion cultures 179 . It is expected that in a near future these 3D models will replace the monolayer cell cultures to better investigate eventual adverse effect of nanoMOFs.

Concerning the in vivo toxicity tests of nanoMOFs, very few studies have been performed to date.

The first study describing toxicological information in vivo of nanoMOFs was reported by some of us in 2010 90 . In this pioneering experiment, iron carboxylate NPs such as MIL-88A NPs with particle sizes of 150 nm or 500 nm, as well as MIL-88Bt NPs with 50 nm or 140 nm were intravenously administered into Wistar rats via the jugular vein. Animals were followed up to 3 months and several parameters were evaluated such as animal behavior, animal weight evolution, weight changes of different organs, cytochrome P-450 activity, ALT and AST transaminases levels and interkeukine-6 serum concentration. The only significant effect was a slight transitory increase in the spleen and liver weights, attributed to the fast sequestration by the reticuloendothelial organs of the nanoMOFs that went back to normality one to three months after injection. The absence of immune or inflammatory reactions after NP administration supported their lack of toxicity. Moreover, the absence of activation of cytochrome P-450 suggested a direct excretion of the polyacids, in agreement with their high polarity.

In 2013 some of us investigated the in vivo toxicity and biodistribution of three different uncoated iron carboxylate based nanoMOFs, namely MIL-100(Fe), MIL-88A(Fe) and MIL-88B_4CH 3 (Fe), using Wistar rats 124 . First, i.v. administration of increasing doses of these NPs was performed to assess lethal dose 10 (LD10). Animal behavior during 7 days was analyzed and then animals were sacrificed and organs harvested for histological examination. This early study led to very promising results. Indeed, the LD10 was never reached with the studied nanoMOFs, any mortality was caused by nanoMOFs administration or significant toxicity signs were observed. In fact, the highest practicable dose was established based on colloidal stability and not on toxicity issues. Only an expected increase in the oxidative stress was observed but it came back to the control level after one month.

In 2015 a third study systematically describing the in vivo toxicity of a series of 9 different nanoMOFs was reported by Ruyra and colleagues 184 . In vitro cytotoxic test using 2D cell culture were compared to results obtained using the zebrafish embryo in vivo model (Figure 13). On the whole, a strong correlation between the results using both methods was found except for MIL-101(Fe), that was more toxic in vivo in zebrafish than in vitro. Very interestingly, authors concluded that the toxicity of these materials was mainly governed by the release of metal ions during degradation, while they also highlighted the importance of the formation of other species upon degradation. This confirms that when assessing nanoMOF's toxicity more information about degradation mechanisms and degradation products under physiological conditions is required. For a while, it was generally assumed that nanoMOF in vivo degradation resulted in releasing the constitutive metal ion and the ligand. However, this is in contradiction with the findings of Ruyra et al. that have proven that a few nanoMOFs once in contact with cell culture medium became amorphous and underwent structural rearrangements and/or reactions that generated new inorganic species, responsible for adverse effects 184 . In a more recent study, this phenomenon has been described in detail for MIL-100(Fe) nano-or micro-particles, confirming the progressive formation of dense phases (iron oxide and phosphate) associated with a degradation in physiological simulated conditions 171 .

Figure 13.

In vitro and in vivo models employed to assess nanoMOFs toxicity.

Besides the above mentioned studies, few recent reports dealing with in vivo nanoMOFs applications disclosed also brief results concerning in vivo toxicity 129,138,185,186 , but it is far from being enough to establish material biocompatibility.

As an intermediate conclusion, despite promises, there is still a great effort to be carried out to assess the toxicity profiles of nanoMOFs before eventual translation into the clinic. Until that, toxicity remains without any doubt one of the keystones to continue building the bridge to move from bench to bedside.

An important advantage of MOFs compared to other nanomaterials lies in the preparation of

Metal Biomolecule Frameworks (also known as bioMOFs for bioactive MOFs or bioMIL for bioactive MIL) by the incorporation of endogenous molecules or active ingredients as building blocks 107 . This allows reducing the amount of non-active undesirable compounds to be administered at the benefit of an anticipated decrease of toxicity. Numerous active ligands have been proposed to date to produce bioMOFs: peptides (Metal peptide frameworks) 187 , nucleobases 188,189 , carbohydrates [190][191][192][193][194] , porphyrines 114,195 , as well as some active ingredients 196- 201 . Unfortunately, most of these bioMOFs are still at their initial stages of development and very few of them have undergone in vivo preclinical evaluation. In 2014 Lu et al. performed the in vivo administration of a porphyrine-based bioMOF 114 . Using a subcutaneous xenograft model in mice of human head and neck cancer cells SQ20B, this study was the first in vivo proof of concept of using nanoMOFs for photodynamic therapy purposes. Of note, MOFs were only locally administered at the tumor site and in vivo toxicity was not assessed since the authors only measured anticancer activity. Therefore, the in vivo toxicity associated to this bioMOF still needs to be determined.

In addition to the toxicity associated to the nanoMOF itself, toxicity issues can be intimately related to the synthesis process, especially to the solvents and reaction modulators employed to prepare the NPs. Indeed, the synthesis of most bioMOFs require the use of toxic solvents such as dimethylformamide or pyridine 199,200 . Thus, to prepare suitable nanomaterials relevant for clinical applications, alternative "green" synthesis routes are urgently needed as it has already been done for other MOFs 131,139,201 .

Biodistribution, targeting, and pharmacological efficacy

The major aim in using nanocarriers for drug delivery is related to the improvement of the pharmacokinetic profile and biodistribution to allow a better drug targeting towards diseased cells and tissues. Despite the great number of studies claiming that nanoMOFs are promising drug carriers, very few of them focus on demonstrating the ability of these systems to modify and improve drug biodistribution.

The first investigation dealing with the pharmacokinetic and biodistribution of nanoMOFs was reported by Baati et al., in which the long term biodistribution (from 1 to 30 days) of three different uncoated iron carboxylate nanoMOFs was examined in rats (Figure 14) 124 . After i.v.

administration of MIL-100(Fe), MIL-88A(Fe) and MIL-88B_4CH 3 (Fe) NPs, both the iron and the organic linker concentrations were quantified in several complex biological media, including liver, spleen and urine. Iron levels were quantified by atomic absorption spectroscopy (AAS), whereas the linker concentration was determined by specific extraction and high performance liquid chromatography (HPLC) methods 202 .

Figure 14. Iron level in different organs after 1, 7 and 30 days of i.v. administration of glucose solution of 220 mg/kg MIL-100(Fe) NPs in rats. Reproduced from Baati et al. 124 An important reversible accumulation in the organs of the reticulo-endothelial system, liver and spleen, was observed for all the tested nanoMOFs (Figure 15). However, liver accumulation was higher for the MIL-88B_4CH 3 (Fe) NPs, whose constitutive linker, i.e. tetramethyl benzenedicarboxylic acid, exhibits a more pronounced hydrophobic character, which could also explain the slight accumulation of this nanoMOF in the brain that was not observed for any of the other materials. Besides, even if the observed normal breathing of the animals supported the lack of lung toxicity, upon histological examination of the pulmonary tissue of nanoMOFs treated rats, it was observed that a large amount of nanoMOFs was found to aggregate and to accumulate in the lungs. Taken together, these data demonstrated that nanoMOFs distributed in different tissues for a period of time, and as a consequence they could act as prolonged drug delivery systems. However, this first study was focused on the evaluation of nanoMOFs without any drug loaded inside their pores. Recently, we showed how the encapsulation of the anticancer drug Busulfan into MIL-100(Fe) NPs drastically modified the drug pharmacokinetics profile, compared to the commercial formulation Busilvex® 203 . Busulfan detected after MIL100(Fe)-Busulfan administration was much lower than after Busilvex® dosage with a mean area under the curve in a plot of drug concentration in blood plasma vs. time of 2.6 and 25 mg/ml/min respectively. Moreover, it was also demonstrated that clearance of the NPs from the blood was faster in the case of drug loaded nanoMOFs, probably due to changes on the surface of the particles that became more hydrophobic, and therefore, more easily recognized by macrophages in the bloodstream. Increased levels of trimesic acid linker in urine after 24 h, i.e. 2 times higher than after empty NPs injection, were observed. All this indicates that the biodistribution/elimination data of the empty nanoMOFs cannot be extrapolated to drug-loaded nanoMOFs. In fact, depending on the nature of the encapsulated molecule, the biological interaction of the drug delivery system might change. Early biodistribution of MIL-100(Fe) NPs in rats has also been investigated. Interestingly, a reversible accumulation of NPs in the lungs after i.v. administration was observed, which was attributed mainly to a pH triggered decrease of NPs colloidal stability in the blood upon i.v. administration. This unique property of MIL-100(Fe)

NPs hold promise for specific lung targeting. Thus, MIL-100(Fe) NPs loaded with the anticancer drug gemcitabine monophosphate were found to increase the lung concentration of the drug, leading to the efficient treatment of an experimental model of lung metastasis 119 .

Figure 15.

Biodistribution of iron nanoMOFs according to the iron concentration. 75 In other in vivo experiments, biodistribution of coated nanoMOFs was assessed. For example, bioMOF prepared from the therapeutic agent Zoledronate and calcium (CaZol) was modified by incorporating the targeting molecule folate and chains of PEG (Fol-PEG-CaZol) 138 . Folate is a molecule with a high affinity for the folic acid receptor, highly expressed in many human cancers. Future experiments should include a more adequate animal model to evaluate the mechanisms behind 5-FU loaded Cu-BTC antitumor activity. Another example was proposed by Hartlieb et al. 207 where ibuprofen was co-crystallized with γ-cyclodextrins (γ-CD), the carboxylic group of the drug being coordinated to alkali metal cations such as K+. The main target of this porous framework, built up from (γ-CD)6 cubes, was to reduce the time required for maximum uptake of ibuprofen (Cmax) and to increase the half-life of the drug within the body. The benefit in terms of drug bioavailability after oral administration was demonstrated, using a mice model, with a 100% longer half-life in blood samples.

Recent advances in theranostics

Many of the recent advances in the field of MOFs for drug delivery have already been discussed in the precedent sections. In the past few years, another target in the field was to develop more complex systems to integrate therapeutics and diagnostics in a unique theranostic tool 111,136,[208][209][210] .

The potential of iron carboxylates as theranostic agents was first proposed by some of us in 2010, when MRI measurements were performed on Wistar rats after i.v. injection of a suspension of MIL-88A NPs 90 . In this report, it was shown that the iron-based core was responsible for favorable relaxivities and imaging properties. Quercerin (QT) as anticancer agent, in combination with CuS nanoparticles as a PTT agent; this nanosystem was further stabilized with folic acid-bovine serum albumin (FA-BSA). The drugloading capacity of quercetin was found to be much higher than other reported values. QT release was further analyzed under or without NIR irradiation at pH 5.0 or 7.4. Without laser irradiation, QT release was significantly higher at pH 5.0, due to the faster degradation of the nanoMOF. At both pH, QT was release to a larger extent after laser irradiation. Both parameters selectively increased cytotoxicity to cancer cells. Functionalization with FA-BSA then increased the tumor targeting efficiency after intravenous injection of FA-BSA/CuS@ZIF-8-IR820, compared to CuS@ZIF-8-QT. Nevertheless, a strong NP accumulation of the NPs was observed also in liver. However, tumor-bearing mice treated with FA-BSA/CuS@ZIF-8-QU and irradiated with laser demonstrated a high inhibition of the tumor growth. However, the differences between FA-BSA coated and non-coated NPs were shown to be minimal. To ensure the safety of this treatment, an histology analysis was performed. 211 Another theranostic strategy was recently proposed by Sun and colleagues, which combined liposomal formulation of dichloroacetic acid (DCA) with MOF-Fe 2+ NPs (NPs are composed from Fe 2+ and 2-aminoterephtalic acid) (MD@Lip) (Figure 17). DCA is commonly used in therapeutic applications due to its ability to both inhibit the enzyme pyruvate dehydrogenase kinase and decrease mitochondrial membrane potential while enhancing the ROS production in cancer cells. Moreover, MOF-Fe 2+ NPs in the presence of DCA could catalyze the production of H 2 O 2 to generate highly cytotoxic hydroxyl radicals, •OH (Fenton reaction), which further led to cell apoptosis. In vitro experiments using healthy and cancer cells confirmed that MD@Lip led to a high cancer cell death, with negligible effects on healthy liver (L02) cells. An in vivo biodistribution study indicated that NPs were mostly accumulated in tumor and liver; moreover a prolonged blood circulation was detected. Nevertheless, these results were not compared with control groups. More importantly, the in vivo anticancer effect of MD@Lip was confirmed in comparison with DCA coated with liposome (D@Lip) and MOF-Fe 2+ coated with liposome (M@Lip). 167 Cheng and coworkers proposed a biomimetic cascade nanoreactor (Men@GOx@ZIF-8@BDOX)

for tumor targeting starvation chemotherapy, by coating ZIF-8 with tumor cell membrane (4T1 membrane) and glucose oxidase (GOx), while encapsulating BDOX, an H 2 O 2 -sensitive doxorubicin prodrug. The starvation therapy, based on glycolysis processes, is known to change the tumor microenvironment through an enhanced intracellular acid and increased concentration of hydrogen peroxide (H 2 O 2 ). In vivo biodistribution after i.v. injections into 4T1 tumor-bearing mice indicated that Men@GOx@ZIF-8@BDOX did accumulate mostly in the tumor and were longer detected, in comparison with ZIF-8@BDOX, which also accumulated in the lungs.

Moreover, the therapeutic efficacy was evaluated and showed that Men@GOx@ZIF-8@BDOX had the best effect in terms of tumor inhibition growth, compared to the control groups. In addition, authors evaluated in vivo therapeutic effects of these NPs against the heterogeneous CT26 tumor without however any positive results. This could be due to a lack of specific targeting ability of coating membrane regards to the homologous tumor. 212

Conclusions and perspectives

MOFs materials have been developed initially due to their promises for numerous industrial applications such as gas separation/storage or catalysis, due to their tunable porosity and large chemical and structural diversity. In parallel, the field of biomaterials has been under continuous development with numerous natural and synthetic materials being considered for controlled drug delivery and theranostic applications. The temporal convergence of MOF discovery and biomaterials science expansion explains probably why these new hybrid materials, prepared by chemists far from the clinical environment, were soon identified as potential candidates in the search of new and more safe drug nanocarriers 90,220,221 . Since then, in the last decade, studies dealing with applications of MOFs in the pharmaceutical sciences have grown steadily (see Figure 18). Initial reports in this domain have focused on fundamentals that govern the drug loading and release of a large number of therapeutic and theranostic systems, while later reports incorporated basic in vitro characterization including first toxicity studies. Noteworthy, during the past 3-4 years, an increasing number of studies have highlighted the first in vivo benefits of nanoMOFs to treat different diseases in animal models. One can thus expect in the upcoming years exciting new developments concerning the preclinical in vivo evaluation of MOFs 119,129,186,205,[208][209][210]215,222,223 . A key condition of success will nevertheless require furthering strengthening the exchanges between chemists, pharmacists and clinicians, for the creation of multidisciplinary projects dealing with the translation of nanoMOFs into more realistic biomedical solutions.

Figure 18. Representative evolution on the number of articles on metal-organic frameworks for drug delivery since 2008. In Pubmed, the initial research topic of "metal-organic framework" was further refined using the research topic of "drug delivery".

The remaining challenges are, however, still numerous if one targets the use of nanoMOFs as third generation drug delivery systems as summarized by Hee et al. 224 . According to Park's group classification, third generation formulations are modulated delivery systems that must be able to cross both physicochemical and biological barriers (Table 4) 224 . A particular effort will be required to get a rational design approach covering all the development stages, including the choice of the nanoMOFs composition, the development of safe synthesis, the surface modification and formulation conditions, the best route of administration and dose. All those developments have to be made bearing in mind the current treatment for a specific disease and the identification of the best animal model for the in vivo preclinical evaluation. Even if the number of works rapidly increases every year (Figure 18), nanoMOFs use as delivery systems are still at their infancy. They appeared more than 30 years after the first drug delivery systems. For instance liposomes were used in the 60s 227 , polymeric particles in the 70s 228 , and polymeric micelles in the 80s 229 . Thus, results with nanoMOFs have not reached yet the maturity of other nanovectors in terms of pharmaceutical development, and it is difficult to make a rigorous comparison. However, nanoMOFs present two important advantages that could help to overcome limitations to cross biological barriers, such as the gastrointestinal or the blood brain barrier, in contrast to other materials used in drug delivery. First, the great flexibility of nanoMOFs chemistry offers unprecedented possibilities to easily modify the surface of the NPs to modulate biological interactions. Second, nanoMOFs drug loading is in many cases far superior to most of the other studied drug delivery systems, meaning that even with a limited passage, the pharmacological effects would be significantly improved; one shall also consider that despite biodegradation, the kinetics of release of the drug can be controlled through a careful tuning of the host-guest interactions. Finally, when developing new nanoMOFs based delivery systems, one shall also pay attention to the ability to produce nanoMOFs under conditions compatible with the pharmaceutical industry from their scale-up to their integration into adapted pharmaceutical formulations. Use the adequate control groups to prove efficacy (empty carrier, free drug).

•

When multiple parameters are evaluated within a unique delivery system (i.e. surface modification and drug release, co-delivery of drugs) add the corresponding controls (i.e. empty carriers with and without surface modification; free drugs separately and in combination, and loaded drugs separately and in combination).

In conclusion, MOFs are promising materials for biomedical applications but before being able to reach the clinical arena, MOF scientists still have to get inspired from the knowledge acquired during decades on other nanocarriers (organic, hybrids or inorganic) and to solve the limitations of already developed nanomedicines. 117 To summarize, nanoparticles once injected in the blood easily adsorb protein present in the blood, forming "corona protein" on the surface of nanoparticles. This process led to recognition of nanoparticles by macrophages and removal of NPs by liver, spleen or bone marrow (RES organs). For efficient tumor targeting and drug delivery this is highly undesirable. To prolong NPs blood circulation, and thus improve tumor targeting abilities, NPs need to be properly designed. Some parameters of nanoparticles, such as size, shape and surface charge have high impact on success during development of stealth nanoparticles. Usually non-spherical NPs exhibit an increased blood circulation times when compared to the spherical NPs. 8,9 Size of nanoparticles should not extend 200 nm. Above that there is high risk of macrophage recognition of nanoparticles and removal from bloodstream. However, also very small nanoparticles (less than 10 nm) are not favorable in case of tumor targeting. They can easily escape from tumor and be cleared from blood by kidneys. 7 In case of surface charge, nanoparticles with high positively or negatively charged surfaces (≥±25 mV) have better stability in physiological media, due to repulsive forces. Strongly repelled NPs have fewer tendencies to aggregate and there are easily re-dispersed once injected. 9 Furthermore, nanoparticles need to have modified surface with coatings molecules, which are able to prevent adsorption of protein present in the blood and thus prolong NPs blood circulation. Various coatings were proposed for different nanocarriers.

Surface modification of NPs may improve physiological stability of NPs and stealth properties, as well as prolonged degradation profile of NPs. Density of the coating also have important impact on macrophage recognition. 10,14,18 As example, PEG with high density has better shielding ability against protein present in the blood. However, even high PEG coating cannot prevent adsorption of all proteins. All these parameters need to be considered for successful application in biomedicine.

2009,

Introduction

This chapter deals with the use of MIL-100(Fe) nanoparticles for biomedicine purposes. Based on previous data, the behavior of this nanoMOFs-based formulation upon intravenous (i.v.) administration was already reported (see Figure 1). [1][2][3] MIL-100(Fe) NPs are stable in aqueous solution at pH lower than 6.5 (except at the isoelectric point around 4) but when pH gets close to 7.4 (physiological pH) nanoparticles start to highly aggregate. This is explained by a strong change of environment, from the glucose injection solution to blood, which results in a massive aggregation of nanoparticles and leads to a nanoparticle filtration by lungs. However, after a few hours, these aggregates start to disassemble, while progressively releasing encapsulated drugs.

This disaggregation process may be triggered by the degradation that occurs in serum with phosphates in a large amount that disrupt the metal to iron bonds leading to a degradation of the nanoparticles. 1 Recently, some of us did investigate the degradation mechanism of MIL-100(Fe) particles in phosphate buffer and showed that an amorphous iron phosphate shell grows progressively with time around a crystalline MOF core. This corona, however, does not change the size distribution of nanoparticles 4 and probably stabilizes nanoparticles in the blood prior to disassembly of the aggregates. Proteins present in the blood may also impact the degradation, thus favoring NPs' disassembly. 1 However, when aiming to treat other types of cancers, such a special biodistribution with nanoparticles mostly accumulating in the lungs, may be considered as a drawback. Therefore in this chapter, we tried to prevent this lung accumulation, as well as to obtain a prolonged D nanoparticle blood circulation (after i.v. administration). This was done through the grafting of polyethylene glycol (PEG) molecules at the outer surface of nanoMOFs. We selected various methods, reported previously by some of us or other groups, that already demonstrated in all cases excellent colloidal stability as well as preliminary in vitro evidences of lower macrophage capture. PEG molecules were grafted either covalently or non-covalently at the surface of MIL-100(Fe). These coatings have been already described in chapter 1. The first route of coating involves amino-PEG (5kDa) that is covalently attached to unsaturated functional -COOH groups of organic linker through an amidation by EDC hydrochloride and sulfo-NHS; these coated nanoparticles are denoted later as MIL-100@PEG NHS (Figure 3 A). 5 A second approach, involves acryl-PEG (5kDa) grafted non-covalently on MIL-100(Fe) NPs through a radical polymerization (Graftfast). This reaction occurs upon redox activation of an aryldiazonium salt, which initiates the radical polymerization of the acryl-PEG to form a polyphenylene first sublayer and the coating is formed by growing oligoradical chains containing the PEG moieties with the polypherene-like prime layer (Figure 3 B); these coated nanoparticles are denoted MIL-100@PEG Graftfast. 6 The last approach is based on β-cyclodextrin phosphate (CD-Ph) or βcyclodextrin polyphosphate further associated with an adamantane-PEG group (CD-Ph-PEG, 2kDa), covalently attached to the outer surface of MIL-100(Fe) NPs through direct coordination of the phosphate groups on the Lewis iron(III) sites from the outer surface (Figure 3 C). Later, these nanoparticles are denoted MIL-100@CD-Ph and MIL-100@CD-Ph-PEG. 7 PEG. Note that previously these coatings were reported using NPs made through a microwaveassisted hydrothermal route with an average particle size of 130 nm. Now, we had to transpose these routes to the newly developed smaller NPs obtained under conventional heating at low temperature (RT to 60°C) and ambient pressure. Thus, these results are novel although we have followed existing surface modification routes.

Next, we studied in detail the resulting colloidal stability of these coated nanoparticles in water, PBS and serum, mimicking in vivo conditions. Finally, we injected these nanoparticles intravenously in mice and studied in detail their biodistribution and pharmacokinetics. At the end of this chapter, we have performed a scale-up of the room temperature reaction of MIL-100(Fe)

NPs, as well as performed preliminary attempts to obtain smaller nanoparticles.

Contribution

All nanoparticles synthesis and characterizations were done by myself.

3.

Synthesis and characterization of nanoparticles.

Synthesis of MIL-100(Fe) NPs

Bare MIL-100(Fe) NPs were synthesized through a "green" room temperature ambient pressure route following a recipe developed very recently in the group. 8 In a 250 ml vial (e.

Surface modification

A summary of the synthesis conditions of MIL-100(Fe) NPs and coating conditions are presented in Table 1. More details are given in Annex 2.

(a) MIL-100(Fe)@polyethylene glycol Graftfast (MIL-100@PEG Graftfast, 5kDa) (b) MIL-100(Fe)@ polyethylene glycol NHS (MIL-100@PEG NHS, 5kDa) (c) MIL-100(Fe)@β-cyclodextrin-phosphate (MIL-100@CD-Ph) (d) MIL-100(Fe)@β-cyclodextrin-phosphate-PEG (MIL-100@CD-Ph-PEG, 2 kDa).

Table 1. Summary of the synthesis conditions of MIL-100(Fe) NPs (room temperature, 48h of reaction) (A) 8 and different coated NPs synthesis: (B) MIL-100(Fe)@PEG Graftfast, (C) MIL-100(Fe)@PEG NHS, (D) MIL-100(Fe)@β-CD-Ph and (E) MIL-100(Fe)@β CD-Ph-PEG.

Characterization

A set of physico-chemical characterizations was performed to check the optimal formation of MIL-100(Fe) NPs, including X-ray powder diffraction (XRPD), dynamic light scattering (DLS),

ζ-potential, transmission electron microscopy (TEM), scanning electron microscopy (SEM), Thermogravimetric analyses (TGA), N 2 sorption (BET) and Fourier transform infrared (FTIR).

The MIL-100(Fe) crystal structure was confirmed by powder X-ray diffraction (PXRD) analysis.

Patterns were collected at room temperature in the 2° < 2θ < 30° range with a step width of 0.02°.

Fourier transform infrared (FTIR) spectra were collected in the 4000-400 cm -1 range to confirm the absence of any free linker. The porosity of nanoparticles was evaluated by N 2 sorption isotherms at 77 K using approx. 50 mg of sample outgassed at 150 °C under primary vacuum for 6 hours. Thermogravimetric analyses (TGA) were performed in the 25-600 C temperature range under a 3°C/min heating speed and O 2 flow of 20 ml/min, to calculate the wt% of the coating. To measure particle size and surface charge of nanoparticles, Dynamic Light Scattering (DLS) was used; to ensure an optimal nanoparticle dispersion, NPs at a concentration of 1 mg/ml at RT in water were dispersed using an ultrasound tip (10% amplitude for 0.5 min). The size was given as the size distribution corresponding to the hydrodynamic diameter of NPs (% number).

Uncoated MIL-100(Fe) NPs were obtained with a yield close to 60 % (based on iron). PXRD (Figure 5) showed the main characteristic diffraction peaks of the MIL-100 structure, with however slightly broader reflections, in agreement with the presence of smaller crystals. FTIR spectra exhibit the characteristic peaks of MIL-100(Fe) (Figure 6) with, in addition, a small peak at ca. 1700 cm -1 , assigned to carboxylic stretching vibrations (C=O), due to residual traces of unreacted linker, trimesic acid. If this peak is intense, this means that a significant amount of free trimesic acid molecules is still trapped within the pores of the MOF and, therefore, one shall proceed with additional washing. The peaks assigned to the carboxylates connected to the iron (III) sites are observed at 1590 and 1350 cm -1 . Another characteristic peak for MIL-100(Fe) is located at 750-700 cm -1 area and corresponds to the µ 3 O center from the iron trimers. N 2 sorption isotherm at 77K is typical of MIL-100(Fe), with a BET surface area, a BET of 1570 m 2 /g (Figure 7) showing the typical sub-steps characteristics of the MIL-100(Fe) solid, with two types of mesoporous cages. TGA shows the typical weight loss for MIL-100(Fe) in the presence of oxygen (Figure 8 and Figure A2-1) with a 68 wt% loss attributed to the ligand and ca. 32 wt% to the inorganic part.

The calculation method to determine the composition of NPs from TGA is detailed in Annex 3. The hydrodynamic diameter of MIL-100(Fe) NPs in milliQ water (1 mg/ml) at pH 4.5 is 115±3 nm (% Number) and 150±8 nm (% Intensity), while ζ-potential of NPs is -36±2 mV (Table 2).

For an easier comparison, when disclosing our next DLS results, only the hydrodynamic diameter in % number will be compared. We observed here that our NPs are well dispersed, with only a single particle population in agreement with a low Polydispersity Index (PDI= 0.23). TEM images of MIL-100(Fe) NPs before and after surface modification are shown in Figure 9.

Uncoated MIL-100(Fe) presents spherical NPs with a size of 38±8 nm. The particle size of coated NPs is as expected slightly larger and depends on the coating. MIL-100(Fe) coated with polyethylene glycol Graftfast, polyethylene glycol NHS, β-cyclodextrin-phosphate an βcyclodextrin-phosphate-polyethylene glycol have average sizes of 44±5 nm, 45±7 nm, 52±7 nm and 46±6 nm, respectively (Figure 9). Powder X-ray diffraction (PXRD) patterns of uncoated and coated NPs confirmed that the crystalline structure of MIL-100(Fe) NPs was not altered after surface modification whatever the coating (Figure 10). Only small changes in relative intensity, in agreement with different hydration rates, were observed (as one did not expect any of the surface grafted molecules to enter the pores of the MIL-100 NPs). evaluated as a function of pH by DLS, within the 2.5-7.4 pH range (Table 3 and Figure 12). A dispersion of MIL-100(Fe) NPs, at a concentration of 1 mg/mL, was freshly prepared in each case by sonication using an ultrasound tip (10 % amplitude, 20 seconds). The desired pH was adjusted using HCl or NaOH aqueous solutions. All measurements were performed in triplicate for three independent samples. DLS measurements were performed immediately after the redispersion of NPs. DLS measurements took approximately 5-10 minutes.

In agreement with previous results, the hydrodynamic diameter and surface charge of MIL-100(Fe) strongly depend on pH. 1 Uncoated MIL-100(Fe) NPs at pH 5.4 have a size of 106±2 nm with a satisfactory polydispersity index (PDI=0.3) and strong negative ζ-potential (-42±3 mV)

(Table 3 and Figure 12). NPs tend to aggregate at lower pH (2.5), while ζ-potential is still positive (12±2 mV). As expected a high degree of aggregation occurs at pH near the point of zero charge (ca. 3), which corresponds to one of the two pKa values of trimesic acid (3.14). At higher pH, the surface charge of MIL-100(Fe) strongly decreases (-40±5 mV), while NPs are more stable, and size remains small (~110 nm). MIL-100(Fe) NPs start to aggregate again at a pH higher than 6.5. At physiological blood pH (ca. NPs start however to aggregate after 7 h in PBS (with 2 populations of 140±12 (70%) and 320±10 (30%)); after 24 h in PBS, the size of NPs is even bigger (size of 676±14 nm), probably due to a degradation of the framework triggered by the phosphates from PBS. Some coatings have however the ability to protect NPs from this process. Among the different coatings, the best results are obtained for the β-CD-Ph and β-CD-Ph-PEG coatings with, after 7 and 24 h, similar particle sizes (137±3 nm (95%) and 162±5 nm (90%), respectively), however with the appearance of a second peak (size 70±10 nm (5%) and 75±10 nm (10%), respectively). β-CD-Ph-PEG coated NPs remained of the same size (~140 nm) until after 24 h of contact in PBS a second population of NPs is observed (320±10 nm (5%)). In both cases, degradation and/or aggregation of NPs probably occur. Degradation may trigger also a partial disassembly of previously aggregated nanoparticles (MIL-100@β-CD-Ph.) PEG Graftfast and PEG NHS coatings were however not very efficient with aggregation already occurring only after 2-3 h of contact of NPs in PBS. Here the absence of phosphates in the surface molecule is probably the main reason why such coatings are less efficient as it was shown before that MIL-100(Fe) NPs degrades through the progressive replacement of trimesate ligands from the nanoMOF structure by phosphate ions from PBS. 4 In a nut-shell, NPs coated with phosphates molecules (β-CD-Ph and β-CD-Ph-PEG) behaves the best in PBS. In order to mimic conditions used for in vivo experiments, 0.2 ml of NPs (uncoated and coated MIL-100(Fe)) were redispersed in glucose 10% (c=5 mg/ml) by ultrasound tip (10 % amplitude, 20 seconds) and further mixed with 1.5 ml of filtered rat serum (ca. volume of blood in 20gmouse is equal to 1.5 ml) by ultrasound tip (10 % amplitude, 20 seconds). Immediately after mixing with serum, size and ζ-potential were measured in triplicate (c=0.6 mg/ml). Additionally, the impact of the initial concentration of NPs redispersed in glucose 10% (c=1.3, 2.5 and 5 mg/ml) on the NPs size in serum was measured (Figure 14 and Table 5).

The particle size of MIL-100(Fe) in serum for uncoated NPs was of 206±8 nm, while NPs coated with PEG Graftfast, PEG NHS, β-CD-Ph and β-CD-Ph-PEG in serum exhibited sizes of 175±20, 180±20, 162±10 and 178±5 nm, respectively (Table 5 and Figure 14). Compared to more acidic pH, the particle size of NPs in serum increased around 2 times whatever the NPs formulation. If one dilutes the solution by a factor two (c= 2.5mg/ml NPs in glucose 10%) this did not decrease the aggregation in serum. Noteworthy, NPs that have not been washed after the initial reaction exhibit a smaller size in water (ca. 70 nm), compared to washed-NPs (ca. 110 nm). Then, these unwashed NPs were coated and washed only after the coating and their size assessed;

unfortunately, the NPs aggregated the same way as the previous (washed) particles. Only NPs coated with PEG Graftfast before washing and using 2 times diluted conditions led to a slight decrease, but still limited, of NPs size in serum (ca. 150 nm), compared to normal synthesis (ca.

190 nm). We think this particle aggregation is due to a combination of several factors. 5) 14 C-MIL-100(Fe)@ 3 H-Gem MP@β-cyclodextrin-phosphate ( 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph) (βCD-phosphate functionalized radiolabeled (nanoMOFs + drug)) 6) 14 C-MIL-100(Fe)@ 3 H-Gem MP@β-cyclodextrin-phosphate-polyethylene glycol ( 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG) (βCD-phosphate / Adamantane PEG functionalized radiolabeled (nanoMOFs + drug))

Synthesis of 14 C-MIL-100(Fe) NPs

In a 20 ml glass vial, trimesic acid (29,24 mg, 0.14 mmol) and iron (III) nitrate nonahydrate (84.3 mg, 0.21 mmol) were mixed with distilled water (10.54 ml) under magnetic stirring for 2 minutes. Immediately after, 14 C-trimesic acid in ethanol (87.84 µl, 8.78µCi) was added to the mixture. The solution was then left under magnetic stirring at room temperature at a speed of 600 rpm for 48 h. The product was then transferred into Eppendorf tubes (1.5 ml), then centrifuged (14500 rpm, 10 min) and washed once with distilled water, and twice with ethanol by redispersion (ultrasonication tip, 20% amplitude, 20 seconds) and centrifugation (14500 rpm, 10 min). The resulting NPs were kept in absolute ethanol in the fridge (4°C). Before further encapsulation, NPs were washed in milliQ water.

Encapsulation of 3 H-Gemcitabine monophosphate in 14 C-MIL-100(Fe) NPs

In one Eppendorf, 14 C-MIL-100(Fe) NPs (8 mg) (NPs stored in ethanol) were first centrifuged (14500 rpm, 10 min) and washed with milQ water by re-dispersion (ultrasonication tip, 20% amplitude, 20 seconds) and centrifugation (14500 rpm, 10 min). These NPs were then redispersed in milliQ water (1.8 ml). In a second Eppendorf, a powder of Gemcitabine monophosphate (2.88 mg) was prepared. Immediately after the redispersion, NPs solution was added to Gemcitabine monophosphate, and the 3 H-Gemcitabine monophosphate (9.6 µl, 19.2 µCi) solution in ethanol was added to this mixture. NPs were kept under stirring overnight.

Loaded NPs were finally recovered by centrifugation (14500 rpm, 10 min) and washed once in milliQ water. Typically, 3 H-Gemcitabine loaded NPs were prepared one day before the planned injections and surface functionalization was performed on the day of injection.

When injecting free Gemcitabine monophosphate, Gemcitabine monophosphate (2.4 mg) was dissolved directly in glucose 10% aqueous solution (1.6 ml) with the addition of 3 H-Gemcitabine monophosphate in ethanol (8 µl, 16 µCi).

Surface modification of 14 C-MIL-100(Fe)@ 3 H-Gem MP

Different surface modifications were prepared according to the methods described before.

Prepared NPs (8 mg) were washed in water and re-dispersed with glucose 10% (1.6 ml) (sonication 20% amplitude, 10 seconds) and injected during the same day.

In vivo protocol-biodistribution and pharmacokinetics

On the day corresponding to the in vivo experiments, each formulation (6 groups, as mentioned above) was dispersed in glucose 10% by ultrasonication tip (20% amplitude, 10 sec), and was intravenously injected via the tail vein in mice, with a dose of NPs of 50 mg/kg. Considering the average weight of the mouse (ca. 20 g), the concentration of injected NPs was 5 mg/ml. Injected formulations contain 0.5±0.2 µCi of 3 H and 0.06±0.02 µCi of 14 C. Pharmacokinetics (PK) and biodistribution (BD) were studied at different times after injection: 5, 15 and 30 minutes, 1, 3, 6 and 24 h (PK), and 30 minutes, 1, 6 and 24 h (BD). In addition, to minimize the number of animals utilized during these experiments, one mouse was used for two PK points and one BD point. Practically, after i.v. injection, blood was first collected from the facial vein and then once again by cardiac puncture. At the second blood collection, organs were also taken (Figure 15).

Results were calculated based on radioactivity detected in each injected formulation.

Detailed procedures of animal experiments are provided in the Annex 2.

Figure 15.

Schematic representation of the in vivo experiments for biodistribution and pharmacokinetics studies. Each mouse was intravenously injected via tail vein with 0.2 ml NPs in glucose 10% (5 mg/ml). Blood was collected after 5, 15 and 30 minutes, 1, 3, 6 and 24 h after an injection, while organs were collected after 30 minutes, 1, 6 and 24 h after an injection.

In vivo protocol-metabolism study

In order to take into account the elimination pathway of the ligand and of the drug through urines or feces, metabolism at 24 h after injection was studied for 2 groups ( 14 C-MIL-100(Fe)@ 3 H-Gem MP and 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG; 5 animals/group), using metabolic cages.

Preparation of the animal samples for characterization

All samples were analyzed by scintillation counting, for 3 H and 14 C measurement. Solid samples were first dissolved and prepared according to standard methods, described in detail in the Annex 2. Before scintillation counting, samples were mixed with "Ultima Gold" (spleen, liver, blood, and urine) or "Hionic Fluor" (other organs and feces).

Results and discussion

Analysis of radiolabeled BTC and Gemcitabine monophosphate concentration in the tissues (i.e. lungs, liver, spleen, kidneys, heart, fat and muscles) showed that bare MIL-100(Fe) NPs did accumulate mostly in the liver, spleen and lungs, which is in good agreement with previous studies. 2,3 In fact, uncoated MIL-100(Fe) NPs strongly accumulate in the lungs (see top Figure 16), due to a strong aggregation in the blood, leading to a passive lung accumulation due to small capillary vessels that exist in this organ. This phenomenon was already previously described by Simon-Yarza et al. and it was taken advantage of this behavior for the treatment of lung metastasis. 1 PEG NHS, CD-Ph and CD-Ph-PEG surface modifications, respectively). As it is shown in Figure 16, each surface modification method led to a significant reduction of accumulation of NPs in the lungs and was totally cleared from the lungs after 24h. In addition, the lower lung accumulation was observed with the β-CD-Ph-PEG coating, where the lung concentration was almost reduced to zero already at a short time after injection. One can tentatively explain this effect through the highly negative charge of the cyclodextrin phosphate molecules together with the PEG chains.

Moreover, DLS results showed that phosphates present in the blood have minimal effect on nanoparticles coated with phosphates (β-CD-Ph and β-CD-Ph-PEG coated nanoparticles), while nanoparticles coated with other molecules have tendency to aggregate in phosphate environment.

This leads to the best colloidal stability of β-CD-Ph-PEG coated NPs and thus to a lower in vivo blood aggregation that totally ruled out any accumulation into the lungs. As it was described before, the most stable solutions should have strong negative or positive surface charge (ζpotential higher than 30 or less than -30 mV) to prevent from aggregation and maintain a small particle size. Aggregation of nanoparticles may occur due to several reasons. First, the uncoated NPs, once intravenously administrated in the body suffer from strong changes in pH, from acidic (pH 3 to 5) to higher values (pH 7.4) in the blood. Second, high concentration of phosphates, present in the blood, initiates the replacement of the linkers by phosphates from nanoMOF structure and leads to the degradation of NPs and a higher aggregation. Third, proteins present in the serum also trigger the protein corona adsorption on the surface of NPs and led to NPs aggregation. In addition, even if a degradation of the NPs occurs at such neutral pH, it has been shown that the β-CD-Ph-PEG coating is highly stable even under such conditions.

Gemcitabine monophosphate free did accumulate only slightly in the lungs (0.8±0.1 % of injected dose) 30 minutes after injection, while Gem MP accumulation was higher when encapsulated inside the bare NPs (2.3±0.2 % of injected dose). This agrees with a passive lung targeting of the bare nanoMOFs as described before. In the case of PEG Graftfast and PEG NHS NPs, no significant reduction of the accumulation of Gem MP in the lungs (2.2±0.4 and 3.0±0.4 % of injected dose, respectively) was observed. β-Cyclodextrin phosphate modification led, on the contrary, to a decrease of accumulation in the lungs after 30 minutes from the injection (1.2±0.2 % of injected dose), similar to the level of Gem MP free. NPs coated with cyclodextrin phosphate coupled with PEG had the smallest pulmonary accumulation of the drug (0.9±0.2 % of injected dose) after 30 minutes from injection, which is in good agreement with the BTC accumulation data into the lungs. In each case, the accumulation of Gem MP was reduced almost to zero after 24 h from injection. The lower accumulation of 3 H-Gem MP comparatively to 14 C-BTC is probably due to the partially fast release of this drug from the NPs in physiological conditions (described previously by Rodrigues-Ruiz et al.)

Figure 16. Lungs NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injection of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 CMIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top figure represents As seen before, the bare MIL-100(Fe) NPs did strongly accumulate in the liver (22.7±3.3 % of the injected dose for after 30 minutes from injection) (see top Figure 17), which represents a major tissue of the reticulo-endothelial system (RES). The level of the NPs accumulated in the and 37.7±1.9 % of the injected dose 30 minutes after injection for PEG Graftfast, PEG NHS, CD-Ph and CD-Ph-PEG coated NPs, respectively. The liver accumulation of NPs was even higher in the case of CD-Ph-PEG coating: the successful prevention of NPs accumulation in the lungs led to a higher concentration of particle circulating in the blood associated with an enhanced macrophage capture leading finally to an enhanced accumulation in the spleen and liver. We concluded that the cyclodextrin phosphate PEG-coated NPs shall be small enough to avoid any clogging into the lungs, but not efficient to render the NPs stealth enough to avoid liver macrophage capture. One hypothesis to explain this lack of efficiency for the PEG coating is that the size of the NPs was still too big or the PEG coating not dense enough to prevent opsonization, leading to a macrophage capture followed by a fast clearance from the blood. For each formulation, the level of BTC was reduced in the liver after 24 h from an injection.

After 30 minutes, Gemcitabine monophosphate accumulated also highly in the liver for all formulations, which is in good agreement with the BTC accumulation observed in the liver.

After 24 h from injections, the levels of Gem MP decreased almost to zero, due to the fast removal of Gem MP from the body.

Figure 17.

Liver NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top figure represents14 C-BTC accumulations, coming from NPs (ligand in NPs and detached ligand). The bottom figure represents 3 H-Gem MP accumulations, coming from the drug (Gemcitabine monophosphate encapsulated or released).

Bare MIL-100(Fe) NPs did accumulate also in the spleen (1.5±0.5 % of the injected dose for 14 BTC 30 minutes after injection) (see top Figure 18), which is another organ of the RES.

Accumulation in the spleen is not high in terms of % of injected dose, but since this organ is very small, it represents higher accumulation in terms of tissue concentration. PEG Graftfast and PEG NHS coatings did not impact much the accumulation of NPs in the spleen (1.9±0.6 and 2.1±1.0 % of injected dose after 30 minutes from an injection, respectively). Cyclodextrin phosphate and cyclodextrin phosphate-PEG coatings led to a still higher accumulation of BTC into the spleen (3.9±1.9 and 4.1±1.1 % of injected dose after 30 minutes of injection, respectively). Similarly to the liver, NPs coated with CD and CD-Ph-PEG, which had the best reduction in accumulation in the lungs, had also a higher tendency to accumulate into the spleen.

Accumulation of NPs in spleen decreased after 24 h for all formulations Accumulation of Gemcitabine monophosphate showed that coatings did not have much impact. After 24 h, the drug level was reduced to zero whatever the type of formulation due to the degradation of the nanoparticles and excretion of both BTC and gemcitabine.

Figure 18. Spleen NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top figure represents14 C-BTC accumulation, coming from NPs (ligand in NPs and detached ligand). The bottom figure represents 3 H-Gem MP accumulation, coming from the drug (Gemcitabine monophosphate encapsulated or released).

BTC accumulation in the kidneys reached the highest level for bare MIL-100(Fe) NPs and for PEG NHS coated NPs (13.0±5.7 and 10.7±3.4 % of the injected dose at 30 minutes after injection, respectively), corresponding to the fast excretion of BTC and of the drug (Figure 19).

Other coatings led to a lower accumulation in the kidneys. This may be explained by different amount of released BTC and Gem MP from smaller NPs, which led to difference in concentration in kidneys after 30 minutes from the injection. A similar accumulation in kidneys was observed for the pure gemcitabine monophosphate due to urinary excretion pathway. The accumulation of the BTC and Gem MP was then reduced 24 h after injection. Finally, the accumulation of BTC and Gem MP was almost negligible in the heart (Figure 19).

Figure 19. Kidneys and heart

NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top left figure represents 14 Additional tissues, fat, and muscles, were analyzed after 6 h from i.v. injection for a better understanding of the biodistribution (Figure 20). It was observed that accumulation of NPs in the fat remained negligible. BTC and Gemcitabine only slightly accumulated in the muscles.

C-kidneys 3 H-kidneys C-heart 3 H-heart

. Fat and muscles NPs accumulation profile at 6 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top left figure represents Pharmacokinetics assays of BTC and gemcitabine monophosphate showed all a fast clearance of NPs, uncoated or coated, from the blood (Figure 21). BTC concentration of uncoated NPs at 5 minutes from injection reached 16.5±4.0 % of the administered dose in the serum and after 30 minutes, the level of BTC started to decrease down to 9.7±2.1 % of injected dose. This value is much higher compared to the previously reported value 1 , where at the same time concentration of uncoated NPs in the blood was reported to be approximately 1.3 % of the injected dose and after 3 h of 0.1% of the administered dose. This difference is probably due to the size of MIL-100(Fe) NPs. In our case, a room temperature synthesis was considered, instead of a hydrothermal microwave-assisted route, leading to the much smaller NPs (50-70 nm against 150-200 nm for the microwave route). At longer time points, the blood concentration of BTC dramatically decreased for all MIL-100(Fe) NPs formulations. Gemcitabine monophosphate level of the free drug and of drug encapsulated into bare NPs exhibited a similar pharmacokinetic profiles. This level started to decrease after 1 h from injection and was reduced almost to zero 3 h after injection. Unexpectedly, the level of BTC in the blood after the coating of the NPs decreased, whatever the coating, thus NPs were cleared faster from the bloodstream. Similarly, after coating the levels of Gem MP decreased dramatically. The level of Gem MP decreased more rapidly 1 h after injection. 3 h after injection, the level of Gem MP was reduced almost to zero being within the 1.2-2.2 % range of injected dose. In a nutshell, none of these MOFs nanoformulations could be considered as long circulating into the blood stream, despite various PEGylation protocols.

Finally, all the PK and biodistribution results indicate that: (a) it is possible to find a suitable coating to prevent from the accumulation of the NPs in the lungs and (b) none of the coatings was capable to ensure a "stealth" effect in vivo allowing long circulation into the blood compartment.

Figure 21.

Pharmacokinetics profile of intravenously injected formulations (0.2 ml NPs in glucose 10% at concentration 5 mg/ml and dose ~50 mg/kg). (Dark blue) 3 H-Gem MP free, (red) 14 C-MIL-100(Fe)@ 3 H-Gem MP, (green) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet) 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange) 14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). Top figure represents 14 C-BTC accumulations, while bottom figure represents 3 H-Gem MP accumulation.

PK-14 C (%D in whole blood) PK-3 H (%D in whole blood)

We then performed an excretion study at 24 h after injection, with two NPs formulations: uncoated and coated with β-CD-Ph-PEG (Figure 22). This allowed to estimate how much of the injected dose was eliminated from the body and to draw a balance of the injected dose and it was calculated that a total of 60-65 % of the injected dose of BTC was detected in various organs (liver, spleen, lungs, kidneys, heart, fat and muscles), blood, urine, and feces. Most of BTC was eliminated through urines (45.8 and 25.3 % of the injected dose for bare and coated NPs, respectively) and a lower amount through the feces (4.2 and 10.1 % of the injected dose for bare and coated NPs, respectively). In the same way, Gem MP was mainly detected in the urines (52.6 and 37.6 % of the injected dose for bare and coated NPs, respectively) and less in the feces (4.5 and 11.4 % of the injected dose for bare and coated NPs, respectively). It is understandable that hydrophilic small molecules, like BTC and Gem MP molecules were mainly excreted through kidneys filtration via urine. This is also in agreement with our previous biodistribution data. As already discussed before, it is assumed that the drug was released fast after contact of the NPs with the biological tissues, by progressive replacement of BTC and gemcitabine MP by the phosphate from the tissues. Despite their small size, all NPs were still fast captured by macrophages and removed from the bloodstream by the RES organs. BTC and drug were then cleared from the organism mainly by the urine.

While Gem MP was totally cleared from the organism after 24 h, at that time BTC was still present in organs, with the highest accumulation rates in liver, spleen or lungs. Progressive degradation of the nanoparticles leads to a slower linker removal. In a previous study by some of us (Baati et al.) 3 , it was demonstrated that after intravenous injection (studied in Wistar rats) of MIL-100(Fe) (microwave synthesized NPs), all accumulations of NPs in various organs (the highest accumulation of NPs was observed in liver, spleen, and lungs) were decreased to zero after 30 days from injection.

6.

Particle size optimization of MIL-100(Fe) NPs

Motivation

A objective was to optimize the particle size of the nanoMOFs. In fact, as one could see it in the previous sub-chapter, the surface coating with PEG of the nanoparticles of MIL-100(Fe) did not

H-NP-Gem MP-CD-PEG 14 C-NP-Gem MP H-NP-Gem MP

lead to a change in the PK and BD of the drug loaded nanoMOFs. As mentioned in the introduction chapter, getting "stealth" effect depends on several parameters. An important one is the particle size with the smaller the particles the better chances for a given type of nanoparticle and surface coating to avoid macrophage capture. As we observed that in serum conditions our surface coated NPs did undergo a significant increase in particle size due to aggregation, largely exceeding the particle size in pure water. Several synthesis parameters were tried here to further decrease the particle size of the nanoMOFs. Parameters such as the reactant concentration, the time, the temperature, the addition of modulators, as well as microwave irradiation (in contrast to conventional heating) were tuned in order to obtain small enough NPs in serum. Microwave synthesis of nanoMOFs offers many advantages, compared to conventional routes. [9][10][11] In fact, microwave heating permits fast and homogeneous heating of the solution, which in turns facilitate the formation of nuclei throughout the reaction solution. Short reaction time is then required to limit crystal growth. Microwave assisted synthesis is, therefore, a suitable method of choice for MOF nanoparticle synthesis and narrow particle size distribution. Our group previously reported the green synthesis of MIL-100(Fe) under hydrothermal conditions using NPs of ca. 130 nm in size. 12 However, as seen previously, smaller nanoparticles with narrow particle size distribution can also be obtained using conventional heating partly because of a low temperature reaction, low reactant concentrations, and a fast stirring.

Many researchers have reported studies dealing with the formation of MOFs nanocrystals. For instance, Marshall and coworkers proposed a kinetic mechanism, which controls the size of MOF nanocrystals depending on the composition and several parameters such as : (a) linker deprotonation, (b) modulator deprotonation (acidic modulators), (c) linker complexation, and (d) termination. 13 Linker and modulator ligands are in competition for metal ion coordination sites and this process continues until the local concentration of ligands far exceeds the metal ions, thus inhibit particle growth. 13 Modulators are typically monotopic carboxylic acids (e.g. acetic acid, formic acid, benzoic acid) added during nanoMOFs synthesis. [13][14][15][16][17][18] The key role of these additives is to modulate the nanoMOFs size by competing with the ligand for the binding of the metal. The role of Brönsted bases (e.g. KOH, NaOH) is mainly to increase the ligands' solubility and/or change the inorganic condensation. In both cases, the target was to reduce particle growth. 13

Size optimization of MIL-100(Fe) NPs

To obtain NPs with a size lower than 100 nm in physiological media (blood), we first considered three different concentrations of reactants: c1-concentration 1 (standard concentration i.e., from the initial patent 8 ), c2-concentration 2 (diluted, concentration 1 /2), c3-concentration 3 (concentrated , concentration 1 *2). All conditions are summarized in Tables 6 and7.

As the first set of synthetic parameters, additives were used such as acetic acid as a modulator or NaOH both to modify the pH of the reaction and/or to change the solubility of the linker. This was done to impact the nucleation, as well as, the crystal growth in order to tune the particle size.

Adding base leads to easier deprotonation of the ligand making faster the nucleation while acetic acid is a typical inhibitor able to compete with the polycarboxylic ligand. The use of inhibitors allows obtaining either larger particle size or smaller particle sizes depending on the amount used and the system (metal, ligand, solvent) in the study. NaOH was added as a powder to a solution containing trimesic acid and iron (III) nitrate nonahydrate dissolved in distilled water. These reactions were prepared at different concentrations: c1, c2, and c3. The NaOH / linker molar ratio was varied i.e., 0.5, 1, 1.5, 2 or 2.5. Mixtures were kept under magnetic stirring for 48 h (at room temperature). In the case of acetic acid, different molar ratios of acetic acid/iron were chosen: 2.5, 5, 10, 15 and 20. Acetic acid was added to an aqueous solution of iron(III) nitrate nonahydrate dissolved in distilled water in the presence of trimesic acid. These reactions were prepared at different concentrations: c1, c2, and c3. Mixtures were kept under stirring during 48 h (at room temperature).

The impact of co-solvent was also studied. Absolute ethanol was added to the aqueous solutions in order to increase the solubility of the BTC ligand and thus one did expect here to accelerate the nucleation of the nanoMOFs and obtain smaller NPs. Molar ratio ethanol to water: 10:90, The last strategy focused on the use of microwave irradiation at low temperature, 60°C. As described before this method allows homogenous and faster heating and is a powerful method to obtain controlled particle sizes with a low polydispersity. Different concentrations were considered: c1, c2, and c3. As an example, using the same conditions as the patent (c1 concentration), the reactions was set to a power of 1600 W, for 6 minutes, 30 minutes or 60 minutes with 1.5 minutes ramp to 60°C. To stop the reaction (to stop the crystal growth), the reactor containing the nanoparticles was placed in an ice bucket. DLS was used to assess the change in particle size distribution. When nanoparticles were reported to exhibit satisfactory hydrodynamic diameters (below 100 nm in water), we performed PXRD to check the crystallinity of the samples. Then, the particle size was further investigated by SEM/TEM. For the best nanoparticles only, BET and TGA were performed followed finally by DLS analysis of the size of the nanoparticles in serum, following our previously described protocol.

The results are the following. In the case of NaOH addition, only diluted condition (c2) of a ratio of 0.5 NaOH per linker (Table 8) was satisfactory. The size is 77±4 nm in milliQ water (c= 1 mg/ml) was obtained, which is slightly smaller than the one obtained by standard reaction (110±8 nm). The addition of a larger amount of NaOH did not lead to the formation of smaller nanoparticles. Moreover, nanoparticles did not form when a larger amount of NaOH was used.

Similarly, only the addition of the smallest amount of acetic acid (acetic acid/Fe=2.5) led to the formation of MIL-100(Fe). However, the product was obtained as a gel-like material. Larger amounts of acetic acid did lead to the formation of gels together with large particles. As previously described, the addition of small quantities of acetic acid, as a modulator led to the decreased size of nanoparticles. A large amount of acetic acid led to slower complexation of metal-linker and result in the formation of larger particles. 13 Thus, only the MIL-100(Fe) NPs obtained with the addition of 2.5 acetic acic/iron results in a small particle size of 60±4 nm in milliQ water (c= 1 mg/ml) (Table 8). Using mixtures of ethanol and distilled water instead of water as a sole solvent for MIL-100(Fe) formation did not lead to the formation of smaller nanoparticles. Furthermore, only a small amount of ethanol led to the formation of MIL-100(Fe), however as a gel-like product with higher hydrodynamic diameter.

In the case of microwave synthesis, MIL-100(Fe) was formed only when using the longest reaction time (60 minutes), whatever the concentration. Using our patent concentration (c1), the obtained size was 54±4 nm in milliQ water (c= 1 mg/ml), with however a smaller, second size distribution peak of 200±10 nm (2%). Nanoparticles formed under diluted reaction (c2) were larger (91±6 nm in milliQ water (c= 1 mg/ml)) but obtained with a very poor yield. When doubling the concentration (c3) a particle size of 52±3 nm in milliQ water (c= 1 mg/ml) was obtained (Table 8). These smaller nanoparticles exhibit a suitable ζ-potential (below -29 mV) (Table 8). It should be noted that a high ζ-potential leads to higher nanoparticle repulsivity. PXRD patterns of nanoparticles with the smallest particle size are shown in Figure 23, in comparison with the one of MIL-100(Fe) obtained from the standard reaction. All nanoparticles exhibit a satisfactory crystallinity, except those formed with acetic acid (acetic acid/Fe=2.5), that is almost amorphous, in agreement with their gel-like product. NPs made using diluted microwave conditions are also poorly crystalline. Finally, SEM and TEM images were collected for the smallest particles only. As usually observed, the sizes observed by microscopy were smaller than the ones deduced from DLS.

Nanoparticles obtained under diluted condition (c2) with 0.5 NaOH per linker have a particle size of 34±5 nm; that is only a bit smaller than originally obtained nanoparticles (38±8 nm).

Nanoparticles obtained by microwave synthesis in normal concentration, c1 have a similar particle size of 34±6 nm (1 h reaction, 60°C). Only the nanoparticles obtained through microwave synthesis using diluted concentration (c2) exhibit a much smaller size of 22±2 nm (1 h reaction, 60°C). However, a better image would be needed and unreacted linker can also be observed. This indicates that the synthesis reaction was not complete, in agreement with their poor crystallinity and the formation of only a small amount of product. However, nanoparticles obtained by microwave synthesis using a double concentration (c3) possess both a satisfactory smaller size of 25±3 nm (1 h reaction, 60°C) and a good crystallinity. Finally, the particle size of these NPs (uncoated) in serum was smaller than the one of the previously obtained nanoparticles (140-150 nm against 200 nm). Nevertheless, such sizes are still too large to expect in vivo blood long circulation properties. To maintain a small size in serum, these nanoparticles would need further surface functionalization with PEG to improve the colloidal stability and expect a smaller size in serum, suitable to avoid macrophage capture. Due to lack of time, these experiments could not be performed.

To summarize, there are several ways to prepare MIL-100(Fe) nanoparticles, even though the low temperature process allowed for the synthesis of nanoparticle size with narrow size distribution here, our study would have required even smaller particle sizes precluding unwanted aggregation under physiological conditions. The preliminary synthesis optimization tests show however that it was possible to obtain smaller particles with less aggregation in water.

Conclusion and perspectives

Among the large variety of existing nanoMOFs, MIL-100(Fe) NPs are of very high interest for nanomedicine due to their very high drug loading, as well as their biodegradable and non-toxic character. These nanoparticles were previously tested in vivo in mice and rats and showed an unusual behavior with most nanoparticles after intravenous injection being fast cleared from the bloodstream through lung filtration, together with a strong liver and spleen accumulation. rather small observed by electron microscopy (38±8 nm), they have a high tendency to aggregate in aqueous conditions (110±6 nm, uncoated MIL-100(Fe)) and the risk was high to coat not single nanoparticles but aggregates of them. We observed that coating did not affect the particle size in pure water, while uncoated nanoparticles tended to strongly aggregate when pH was close to the one of blood (7.4). The coating was expected to prevent from such aggregation and the most stable particles were those coated with β-CD-Ph-PEG, where particle size was stable in all pH ranges (2.5-7.4). However, in rat serum, all nanoparticles (uncoated and coated) did slightly aggregate (150-200 nm). This is probably due to the strong increase in ionic strength that shields the surface of the NPs as well as to the adsorption of protein from the serum.

In a next step, we prepared MIL-100(Fe) NPs using radiolabelled 14 C-linker and further encapsulated radiolabelled 3 H-Gemcitabine monophosphate. The resulting MIL-100(Fe)@Gem MP NPs were further coated through the different proposed strategies and injected into different groups of animals intravenously. Pharmacokinetics and biodistribution results showed that all nanoparticles (uncoated and coated) were fast cleared from bloodstream by the liver and the spleen. Uncoated nanoparticles, as expected, mostly accumulated in the lungs. However we could demonstrate that the coatings did significantly reduced the pulmonary accumulation, especially after β-CD-Ph-PEG coating. This was likely due to the higher charge of the phosphate CD, as well as, to the dispersive effect of the PEG chains in water, leading to a significant reduction of pulmonary accumulation associated with a better colloidal stability in blood.

However, as shown by PK and BD experiments, the accumulation in the liver and the spleen could not be avoided and it is therefore assumed that CD-Ph-PEG coating was not effective enough to protect NPs from opsonization, macrophage recognition and fast clearance from the bloodstream. In addition, Gem MP was released fast due to rapid exchange with phosphates present in large excess in the blood and other tissues of the body.

The key point to obtain "stealth MOF nanoparticles" would, therefore, to have smaller nanoparticles with a sufficiently dense coating able (a) to avoid aggregation (and lung uptake) as well as (b) opsonization (and liver/spleen uptake). This has so far not been achieved with MIL-100(Fe) NPs. In the last part of this chapter, we tried to synthesize smaller MIL-100(Fe) NPs by modifying synthesis parameters including the use of microwave irradiation. Slightly smaller nanoparticles (22-34 nm) were obtained (microscopy) with a smaller size in aqueous conditions (below 80 nm), leading to promises for a size reduction once coated in blood conditions.

Another possibility would be to lower the concentration of nanoparticles to minimize the aggregation before injection in the blood. Alternative nanoMOFs could also be tried that are more stable in blood conditions and/or other coatings with lower protein adsorption. There are a few examples of nanoMOFs systems where MTX has been successfully encapsulated through the direct synthesis of ZIF's nanoparticles in the presence of MTX. 1,3 However, the use of iron carboxylate nanoMOFs for the encapsulation of MTX had not been reported before. We considered, therefore, the encapsulation of MTX using our recently developed small MIL-100(Fe) NPs (room temperature synthesis) 4 in view of two new collaborative projects, initiated during this PhD (as side-projects), described in detail in this chapter.

Encapsulation and release from MIL-100(Fe)

Experimental procedure

MTX is a poorly soluble drug in water and most polar organic solvents. A green protocol to dissolve in water, on a reproducible manner, MTX had first to be developed. To enhance its concentration, MTX was initially protonated through soaking into an acidic HCl 10 mmol/l aqueous solution using an optimum concentration of 1 mg/ml and then further dissolved in dark conditions, to prevent light sensitive MTX from degradation, after incubating at 55°C during two hours, until all MTX was dissolved. As within two hours at room temperature MTX starts to precipitate again, encapsulation of MTX into the nanoMOF was performed right after the dissolution step. First, two weight ratio of NPs to MTX (1:1 and 1:3) were investigated (Table 1).

10 mg of MIL-100(Fe) redispersed in 0,3 ml milliQ water was added to the MTX acidic solution (cMTX= 1 mg/ml, 10 or 30 ml of HCl 10 mmol/l), and left under sonication (1 minute). The tubes containing the solutions were covered with aluminum foil to ensure dark conditions to the light sensitive drug and were incubated overnight at 37°C, to ensure maximum encapsulation rates and avoid precipitation of unloaded MTX. MOF nanoparticles loaded with MTX were then recovered by centrifugation and washed twice with the same amount of solvent (10 or 30 ml of HCl 10 mmol/l). To quantify the loading of MTX, the supernatant was analyzed by UV-Vis (210-410 nm). Results are presented in the Annex 3 (Figures A2-3). The wavelength for the calculations was selected at 306.6 nm. Calibration curves of MTX in various media (HCl 10 mmol/l, milliQ water and PBS), used for calculation of MTX encapsulation, is represented in Annex 3 (Figures A2-4-A2-7). Note that a particular attention was given to avoid any further precipitation of MTX particularly when the UV analysis was carried out after some delay; in such a case, the supernatants were incubated for two hours at 55°C prior to the UV analysis.

Dried NPs loaded with MTX (MTX@MIL-100(Fe)) were finally analyzed by TGA, also to better quantify the MTX loading.

The MTX release from MTX@MIL-100(Fe) NPs was then studied. The NPs were dispersed in 5 ml of PBS under stirring at 37°C and supernatants were extracted at different times (0, 0.

Results

First of all, all the 'structural' characterizations (PXRD, IR) indicated that the encapsulation of MTX did not lead to any degradation of the MOF framework (Figure 2 and Figure 3). From TGA and UV-Vis analysis, the loading of MTX was estimated. The method used to calculate the encapsulation of MTX from TGA is described in Annex 3 (Figure A2-9). UV-Vis spectrum of MTX in HCl 10 mmol/l and a calibration curve of MTX in HCl are presented in Annex 3 (Figures A2-3 and A2-4). The MTX loading was noticeably higher when a MIL-100(Fe)/MTX 1:3 weight ratio was considered during the encapsulation process with 50±5 wt%, compared to 38±4 wt% when a 1:1 weight ratio was used (UV-Vis method) (Table 2, Figure 4 and A2-9). All methods to calculate the encapsulation of MTX indicated that MTX encapsulation was more effective when more concentrated MTX conditions were considered. N 2 sorption porosimetry was performed only for one example of MTX-encapsulated nanoMOFs (encapsulation in HCl 10 mmol/l in a MIL-100(Fe)/MTX 1:3 weight ratio) (Figure 5).

Encapsulation of MTX was performed by using a small amount of product (10 mg of MIL-100(Fe)). Moreover, the amount of recovered sample was not sufficient to perform any N 2 sorption (NB: typically one shall use a minimum of 30 mg to perform good N 2 sorption). Thus, in order to carry out a proper N 2 porosimetry experiment, a suitable series of MTX loaded nanoMOF was further prepared following the same conditions, by collecting samples with the same encapsulation rate. As expected the pore volume and surface area of the nanoMOFs did strongly decrease after encapsulation in agreement with a partial filling of the mesoporous cages of MIL-100(Fe) (Figure 5).

A

For the next steps, only MTX@MIL-100(Fe) NPs obtained through a MIL-100(Fe)/MTX 1:3

weight ratio was considered. The analysis of the MTX release profile in PBS at 37°C did show that the release of MTX is rather fast with most of the drug released during the first few hours:

78% after only 7 hours, while the rest of MTX is released more slowly, up to 2 days, until all encapsulated MTX is released (Figure 6). PXRD indicates that the NPs degrade rapidly once in PBS explaining mainly why MTX is released rather fast through the degradation of the nanoMOFs (Figure 7). As shown before by some of us, the formation of an iron oxophosphate passivating layer at the outer surface of the nanoMOFs after soaking in PBS probably slows down the degradation of the NPs and thus the release of the remaining MTX molecules from the nanoMOF. 5 However, it has also been shown previously that in spite of the fast degradation of the MOF in PBS, some drugs could be released quite slowly, from a few days (eg AZT-Tp) to weeks (eg Doxorubicin). In such cases, a strong coordination of the drug on the iron open metal sites was the main parameter to explain the more prolonged and controlled release. 6,7 This shows here that MTX exhibits probably only weak interactions with the MOF framework and thus degradation and diffusion drive the release of the drug, being relatively fast. These first results were then further utilized for two different potential applications: encapsulation of MTX using core-shell magnetic nanoMOFs-iron oxide NPs and the toxicity evaluation of molecules in nematode Caenorhabditis elegans (C. elegans) using nanoMOFs.

These results are described in detail in the following parts.

Application to core-shell nanoMOFs -USPIO

Introduction

Recently, some of us reported the preparation of MIL-100(Fe) NPs coated post-synthetically with Ultra-Small nanoparticles of Superparamagnetic of Iron Oxide (USPIO) in order to combine the drug release properties of the nanoMOF with magnetic resonance imaging (MRI) of iron oxides. 8 The presence of maghemite superparamagnetic NPs conferred indeed the nanoMOFs a very high r2 relaxivity value, comparable when 10 wt% of USPIO (γFe 2 O 3 , maghemite) was considered, to the best commercial available systems (i.e. Endorem, Sinerem). Imaging properties of these NPs were then further confirmed in vivo by T2 weighted MRI (mice). This potential theranostic tool improved MRI imaging properties of the nanoMOFs in vivo while conserving the high drug loading/release capacities of the functionalized nanoMOF (doxorubicin).

Recently, Dr. Sene developed a novel MIL-100(Fe)@γFe 2 O 3 nanocomposite, with this time USPIO NPs being embedded directly through direct synthesis into the nanoMOFs crystals. To achieve this, the synthesis of MIL-100(Fe)@γFe 2 O 3 NPs was performed in situ using preformed USPIO in the presence of the precursors of MIL-100(Fe) (iron nitrate, BTC ligand) at room temperature (explained in details in the Annex 2). In previous findings, some of us used doxorubicin as a model drug. Here, we found it interesting to use another challenging drug.

Therefore, we used newly obtained knowledge of Methotrexate encapsulation in MIL-100(Fe), to encapsulate MTX in this core-shell MIL-100(Fe)@γFe 2 O 3 NPs.

Contribution

This core-shell MIL-100(Fe)@ Fe 2 O 3 project was carried out in collaboration with Prof.

Results

Synthesis and characterization

The synthesis of MIL-100(Fe)@γFe 2 O 3 was performed following conditions established by Dr.

Sene (see Table 3) at room temperature over 24 hours. More details about the synthesis conditions and the characterizations are given in the Annex 2.

Table 3. Summary of the synthesis conditions of MIL-100(Fe)@ Fe 2 O 3 . A set of physico-chemical characterization was performed to investigate the optimal formation of the MIL-100(Fe)@γFe 2 O 3 nanocrystals and to quantify the γFe 2 O 3 content, including X-ray powder diffraction (XRPD), transmission electron microscopy (TEM), Thermogravimetric analyses (TGA), N 2 sorption (BET) and Fourier transform infrared (FTIR).

MIL-100(Fe

Powder X-ray diffraction (PXRD) patterns of MIL-100(Fe) and MIL-100(Fe)@γFe 2 O 3 NPs confirmed that the crystalline structure of MIL-100(Fe) NPs was not altered by γFe 2 O 3 (Figure 8).

Moreover, MIL-100(Fe)@γFe 2 O 3 NPs were formed more rapidly than the bare nanoMOFs, after 24 h instead of 48 h, respectively. TGA measurement allowed us to determine the amount of γFe 2 O 3 . TGA graphs of MIL-MIL-100(Fe)@γFe 2 O 3 -5wt%, and MIL-100(Fe)@γFe 2 O 3 -10wt% are presented in Figure 9. TEM images are shown in Figure 10. MIL-100(Fe)@γFe 2 O 3 presents a spheroidal shape, similar to the one of MIL-100(Fe) alone, with the particle size of 72±10 nm and 67±12 nm for MIL-100(Fe)@γFe 2 O 3 -5wt% and MIL-100(Fe)@γFe 2 O 3 -10wt%, respectively. These particles consist in the the assembly of MIL-100(Fe) and γ-Fe

MTX encapsulation and release

First of all, structural characterizations (PXRD, IR) indicated that the encapsulation of MTX did not lead to any degradation of the MOF framework (Figure 12 and Figure 13). Transmitance (%) wavenumber (cm -1 ) MIL-100(Fe) MTX@MIL-100(Fe)@γFe 2 O 3 -10wt% 1702 cm -1

Figure 13. FTIR of MIL-100(Fe) (black) and MTX@MIL-100(Fe)@γFe 2 O 3 (red). MTX was encapsuted in HCl 10 mmol/l at MIL-100(Fe)/MTX 1:1 weight ratio.

MTX encapsulation with the MIL-100(Fe) and MIL-100(Fe)@ Fe 2 O 3 (5 wt% or 10 wt%) was performed according to the previously described method, where 10 mg NPs was mixed overnight with 30 mg of MTX earlier dissolved in 30 ml of HCl 10 mM (at 37°C, dark conditions).

Nanoparticles were then washed twice with the same amount of solvent (HCl 10 mmol/l) by centrifugation (14500rpm, 10 minutes) and redispersion (sonication batch). Supernatants were then investigated by UV-Vis for calculation of the encapsulated MTX (indirect method), while dried powder of MIL-100(Fe) was analyzed by TGA for direct calculation of encapsulated MTX.

After MTX encapsulation, release of MTX from MTX@NPs (NPs: MIL-100(Fe), and MIL-100(Fe)@γFe 2 O 3 -10 wt%, 10 mg) in PBS (5 ml) was measured at different times (0, 0. MTX was encapsulated in 50±1 wt%, 26±2 wt%, and 30±2 wt% in MIL-100(Fe), MIL-MIL-100(Fe)@γFe 2 O 3 -5wt%, and MIL-100(Fe)@γFe 2 O 3 -10wt%, respectively (UV-Vis method).

These results were obtained in triplicate (Table 4). The loading capacity of the nanocomposites was thus slightly lower than that of the bare nanoMOF, partially due to the presence of 5-10 wt % of dense inorganic NPs that do not contribute to the encapsulation of MTX. The diffusion and thus the encapsulation of MTX in the MIL-100@γFe 2 O 3 can be hampered by the presence of maghemite.

Table 4. Encapsulation rates of MTX in NPs: MIL-100(Fe), MIL-100(Fe)@γFe 2 O 3 -5wt%, and MIL-100(Fe)@γFe 2 O 3 -10wt%, calculated from UV-Vis and TGA results.

Name of sample Encapsulation (wt%)-

UV-Vis

Encapsulation (wt%)-TGA theoretical MTX@MIL-100(Fe) 50±1 53±2

MTX@ MIL-100(Fe)@γFe 2 O 3 -5wt% 26±2 30±5

MTX@MIL-100(Fe)@γFe 2 O 3 -10wt% 30±2 32±3

The release profile of MTX from MIL-100(Fe)@γFe 2 O 3 (10 wt%) was similar that from bare MIL-100(Fe) NPs, where most of the MTX (80%) was released within a few hours in PBS while the remaining MTX was slowly released upon 2 days (Figure 14). This indicated that the release profile of MTX was not altered by the presence of 10wt% maghemite. These results were repeated recently by Heng Zhao from our group. However, when compared with new data (Figure A2-8), the release profile obtained previously presented a second slope with a slower releasing drug. In contrast, in the case of a study performed very recently by Heng Zhao, MTX was entirely released within less than 10 hours. Besides, in the previous study, the total incubation time was calculated, excluding the time needed for centrifugation and re-dispersion of NPs in PBS (~30 minutes), however recently (Heng's data) this time was also included to calculate the total incubation time. This indicates that the release profile obtained recently is even faster than the previous one. The difference is probably caused by the use of a higher volume of PBS by Heng Zhao (15 ml instead of 5 ml, as done previously). The degradation of NPs was probably accelerated by the higher content of phosphates and thus fastened the release of encapsulated MTX. The second slope with a slower release of the drug is likely related to partial pore-blocking of the smaller cages of MIL-100(Fe), which might lead to a slower exchange with phosphates and thus a slower release of the drug. In addition, previously, prior to the release study MTX@NPs were dried (vacuum oven, RT), while recently the release profile of MTX was prepared directly from wet MTX@NPs. Moreover, dried NPs have a higher tendency to aggregate once re-dispersed in PBS, and thus the release may be slightly altered compared to the wet samples. HPLC was also considered for more accurate quantification, but due to technical problems, this was done only recently by Heng Zhao in our team. However, the results from HPLC are on the whole comparable with those obtained by the UV method.

MTX@MIL-100(Fe) MTX@MIL-100(Fe)@γFe 2 O 3 -10wt%

Figure 14. The release profile of MTX in PBS (37°C) from MTX@MIL-100(Fe) (red), and MTX@MIL-100(Fe)@γFe 2 O 3 -10wt% (blue), calculated from UV-Vis.

Conclusion

Successful preparation of magnetic MIL-100(Fe)@γFe 2 O 3 nanocomposites, following a previously developed protocol, was carried out. Previously obtained MIL-100(Fe) nanoparticles functionalized with maghemite, were reported to have a very high r2 relaxivity and satisfactory in vivo imaging properties (T2 weighted MRI). These newly obtained core-shell nanoMOFs@maghemite nanoparticles also have excellent magnetic properties, even better than previously observed for nanoMOFs with 10 wt% of maghemite. Based on these results, the newly obtained core-shell NPs@maghemite is also very promising. MTX was loaded in the MIL-100(Fe)@γFe 2 O 3 (5 wt% or 10 wt%) with encapsulation rates of 26±2 wt% and 30±2 wt%, respectively, slightly lower than in the case of the pure MIL-100(Fe) (50±5 wt%). MTX was then released mostly within a few hours of incubation in PBS, the rest being released more slowly during two days. Recently, HPLC was proposed to calculate MTX encapsulation and release, as a more accurate technique. Then, in vitro cell culture tests (cytotoxicity, activity) will be performed. If successful, finally, animal testing will be planned. elegans, including evaluation of the toxicity effect in the worm reproductive system using unloaded nanoMOFs (MIL-100(Fe)) as well as MTX-loaded nanoparticles.

Results

Scale-up synthesis of MIL-100(Fe) NPs

The first objective was to scale-up the low temperature synthesis of MIL-100(Fe) NPs, in order to avoid reproducibility issues with the nanoMOFs synthesis and encapsulation of MTX, and to minimize the time to prepare and characterize the NPs. This is a very important parameter when one considers the use of such nanomaterials for a given application.

The large scale synthesis of MIL-100(Fe) NPs was performed in a 30 liters batch reactor. This represents an increase in the scale by more than 200 times in comparison with the smaller laboratory scale synthesis. Trimesic acid (55.5 g) and iron (III) nitrate nonahydrate (160 g) were introduced in distilled water (20 l) under mechanical stirring at a speed of 350 rpm. The temperature was set at 25°C (external temperature i.e., solution temperature). IR analysis was used to follow the crystallization process and the reaction was finally stopped after 7 days (Figure 16). The product was first recovered by centrifugation using large bottles of 1 liter at 9000 rpm for 15 min. This process was repeated 4 times until all NPs were centrifuged. The solid was then redispersed in water and then recovered by centrifugation. To remove traces of free ligand, a second activation step was carried out in ethanol following the same process. The final product was kept wet in absolute ethanol. NPs was close to 1500 m 2 /g (Figure A2-13). The particle size deduced from TEM analysis is 32±7 nm (Figure 17). DLS gave however a slightly larger hydrodynamic diameter of 104±8 nm.

All these results were in good agreement with those obtained from the smaller scale synthesis.

Encapsulation and release of MTX

Nanoparticles synthesized according to the small-scale route (described in chapter 2) were used only for the preliminary tests. Later, only NPs prepared in large scale synthesis, described above, were used.

Encapsulation of MTX was prepared each time one day before the test with worms.

Encapsulation of MTX was slightly adapted from the previously described protocol as different format of MTX was used. MTX in liquid condition for medical purpose (MTX for injection, in slightly basic solution, 100 mg/ml in 0.01 M NaOH) from Dr. Jean-Claude Alvarez in hôpital Raymond-Poincaré de Garches AP-HP was obtained, which helps to solve MTX dissolution issues. Briefly, an equivalent of 30 mg of MTX (or 10 mg) was mixed either with 30 ml of milliQ water or 30 ml of HCl 10 mmol/l. Next, 10 mg of MIL-100(Fe) NPs were added to the MTX mixture and well mixed (ultrasonication, 1 minute). NPs were encapsulated at 37°C overnight in the dark conditions, to avoid degradation of the drug. Finally, after encapsulation was completed, NPs were washed three times with 30 ml of milliQ water or 30 ml of HCl 10 mmol/l (the same solvent as for encapsulation) (Table 5). The supernatant was analyzed on the same day for the indirect quantification of MTX encapsulation by UV-Vis and dried samples for direct quantification of MTX by TGA. MTX-encapsulated samples were also analyzed by PXRD and IR. 5. Summarized conditions of MTX@MIL-100(Fe) encapsulation (encapsulation in HCl 10 mmol/l or milliQ water).

First of all, PXRD and IR analysis both indicated that the encapsulation of MTX did not lead to any degradation of the MOF framework (Figure 18 and Figure 19). Transmitance (%) wavenumber (cm -1 )

MTX@MIL-100(Fe) in miliQ water MTX@MIl-100(Fe) in HCl MIL-100(Fe)

Figure 19. FTIR of MIL-100(Fe) (black) and MTX@MIL-100(Fe), after loading with with HCl 10 mmol/l (red), or after loading in milliQ water (blue)

The solution of MTX was easily dissolved in both milliQ water and HCl solution. Noteworthy, encapsulation of MTX in aqueous conditions did not lead to the formation of any MTX precipitate, which occurred for a few samples, while encapsulating in HCl. However, encapsulation rates of MTX were higher for encapsulation in HCl (50±5 wt% and 33±3 wt% for encapsulation in HCl and milliQ water, respectively; calculated from UV method). Note that calibration curves of MTX in HCl or milliQ water are presented in Annex 3 (Figure A2-6 and A2-7). The encapsulation of MTX was further confirmed by the TGA method (Figure 20, Table 6). Calculations of MTX loading by TGA is shown in Annex 3 (Figure A2-10).

Although the encapsulation rate of MTX was higher in HCl condition, together with CeleScreen, we decided that aqueous conditions were more suitable for this application, since the encapsulation rates are not the crucial point in this study and aqueous condition gave higher reproductively of MTX encapsulation rates. Since we have decided to work with aqueous loading protocols, thus, only MIL-100(Fe) NPs loaded with MTX under aqueous conditions (nanoMOFs/MTX 1:3 weigh ratio, MTX@MIL-100(Fe)) were used for the release study.

To perform a release study mimicking the condition used for the tests with worms, 8 mg of MIL-100(Fe) with MTX was dispersed in 40 ml of NGM (without agar) or in 40 ml of milliQ water.

We used liquid NGM instead of NGM with agar, since the quantification of the released MTX in solid condition of NGM agar was technically tedious. The released study was performed in dark conditions, to prevent light-sensitive MTX from degradation. The mixture was incubated for different times (5 and 40 minutes, 1, 2, 3, 4, 5, 23 and 30 h, RT). Samples of MTX@MIL-100(Fe) in solution were taken and centrifuged. The supernatants were finally analyzed for MTX quantification by UV spectroscopy.

The release profiles of MTX from MIL-100(Fe) showed that MTX was only slightly released in milliQ water (up to 6 %) while MTX release in NGM occurred at an early time of incubation (19% after 5 h of incubation) and reached 38% after 30 h of incubation (Figure 21). However, due to technical problems with the calibration curve of MTX in NGM solution, these results still need to be confirmed. As mentioned in previous sub-chapter, a new PhD candidate at IMAP, Heng Zhao, is currently trying to improve MTX quantification method using HPLC, instead of UV-Vis, which is expected to reduce reproducibility issues encountered while using UV method.

We speculate that a higher release in NGM is related to the presence of phosphates in the media.

NGM is however a biological media that is more acidic pH than PBS (6.4 instead of 7.4). As it is known that iron carboxylate nanoMOFs degrade faster when the pH increases, this might explain why MTX is released slower, and only partially, in NGM compared to PBS. We also suspect that conditions used here for the study in NGM (stirring, liquid form of NGM solution without agar) naturally triggered a higher degree of release of MTX from nanoMOFs, compared to release in gel-like solid form of NGM, used for real experiments. In conclusion, the release of MTX from nanoMOFs in water or NGM was satisfactory for test in C. elegans. The fact that more than half of MTX remained encapsulated inside nanoMOFs, even after long incubation with NGM, gave high expectations that encapsulated MTX entrapped in nanoMOFs was going to be successfully consumed by worms and would cause the desired effect.

Evaluation of the toxicity effect of MIL-100(Fe) on C. elegans

Prior to any practical use of the nanocarriers, it is of high societal relevance to investigate their possible toxic effects. MIL-100(Fe) has been already demonstrated to have low toxicity upon high dosage in mice study, 18,19 here, however, we focused only on the toxic effect on the reproductive system. Next, to determine the effect of MIL-100(Fe) nanoparticles per se on worm reproduction, L4

worms were exposed to MIL-100(Fe) overnight, and 10 gravid adults were transferred to assay plate on the next day for 5 hours. After 5 hours, the gravid adults were removed and the number of laid egg and hatched larvae were counted (Figure 23). Worms in the presence of E. coli OP50 alone (OP50) or E. coli OP50 mixed with MIL-100(Fe) (MOFs-OP50) were shown no significant difference in terms of laid eggs or hatched larvae (Figure 22). This indicates that MIL-100(Fe) is not causing any detrimental effect on worm reproduction. These results justified the selection of such nanoMOFs as nanocarrier for this proof-of-concept.

Evaluation of the toxicity effect of MTX@MIL-100(Fe) on C. elegans reproductive system

MTX is known to elicit activity on reproduction, therefore, the effect of MTX encapsulated inside the nanoMOFs on C. elegans reproductive system was measured, by quantifying the number of laid eggs and hatched larvae. MTX loading was optimized regarding the nanoMOF/MTX weight ratio, the solvent, time, temperature and stirring speed. The loading was done by simple impregnation in milliQ water for overnight at 200 rpm stirring in dark conditions with nanoMOFs/MTX 1:3 weight ratio.

Next, L4 stage worms were fed with MIL-100 alone or MTX@MIL-100 mixed with E. coli OP50 in 1:1 ratio overnight and 10 gravid adults were transferred to assay plate on the next day for 5 hours. The gravid adults were removed and left for 5 days. It was not possible to determine the number of egg laid because of the high aggregation of MIL-100(Fe), due to its poor colloidal stability in NGM (phosphates, high concentration of ions), therefore, only number of hatched larvae was quantified. Moreover, to confirm the effect on the reproductive system is solely attributed to MTX encapsulated inside MIL-100(Fe) but not the free MTX due to sample preparation or release from MIL-100(Fe), the supernatant of MIL-100 or MTX@MIL-100 was collected and mixed with E. coli OP50 to feed the worm (Figure 24 and Figure 25).

A significant decrease was observed in the number of hatched larvae, with almost 90% reduction in MTX@MIL-100, compared to MIL-100 alone as a control (Figure 25). The experiment was repeated 3 times and the result was found to be reproducible. Moreover, no reduction of the number of hatched larvae was observed in presence of supernatant with free MTX indicating that only MTX encapsulated in MIL-100(Fe) was physiologically administrated into the worm (consumption with food), causing the effect on the worm reproductive system.

Conclusions

In this part, we collaborated with CeleScreen, which aims to help cosmetic or pharmaceutical companies to test the toxicity effect of their novel ingredients using C. elegans as a testing model.

This method may surpass simple in vitro experiments, as well as overcome limitations (e.g. strict law) with animal testing (rodents). First, we succeeded to scale up the room temperature synthesis of MIL-100(Fe) NPs, however requiring a slightly longer reaction time (7 days).

Noteworthy, the total preparation time of scale up synthesis of MIL-100(Fe) was reduced (less unreacted acid leads to fewer washing steps). A large batch of NPs allows using the same batch of NPs during multiple tests and to avoid reproductively issues, sometimes presence while using many different batches of NPs. We slightly modified the protocol of the encapsulation of MTX in MIL-100(Fe) to adapt the different format of MTX used (liquid MTX used for injections). The release profile of MTX from MTX@MIL-100(Fe) in water was negligible, while the release in NGM (common worm's medium) was slightly higher, although lower than in PBS. The release in NGM was, however, still satisfactory, especially when real worms' conditions are taken into account (gel-like solid nature of NGM agar and no stirring). Remarkably, MIL-100(Fe) alone did not affect the worms' general growth and development, as well as reproductive system. Only MTX-loaded NPs showed a change in the reproduction of the worms, while free MTX present in supernatant had no effect on C. elegans reproduction. Moreover, it was demonstrated that the effect was caused only by the MTX encapsulated in the NPs, and not free MTX, after release.

Nevertheless, we proposed a strategy to be taken in order to overcome the problem with counting eggs due to high aggregation of MIL-100 in media. For instance, one could consider the functionalization of nanoMOFs outer surface with biomolecules known to improve NPs colloidal stability in solvents with high phosphate content (e.g., heparin, PEG). Moreover, functionalization of nanoMOFs surface with bacteria, may attract worms to increase consumption of NPs and thus increase the effect of the drug.

In the last decade, Metal-Organic Frameworks nanoparticles (nanoMOFs) have been extensively studied in the field of nanomedicine mainly for controlled drug delivery and theranostics. Some of these materials have demonstrated an excellent colloidal stability in water or PBS. However, their stability in physiological media is still problematic. Thus, this is difficult to predict the effect of NPs after administration in the organism. In particular, there is still a strong need to perform studies to get a deeper understanding of this dependence.

Nanoparticles, once injected intravenously in the body, easily adsorb proteins present in the blood to form "corona protein" at their outer surface. Macrophages then fast recognize these nanoparticles and remove them from the bloodstream by liver, spleen, or bone marrow (i.e. the so-called reticulo-endothelial system, RES organs). Fast elimination of nanocarriers from bloodstream and accumulation in RES organs is a typical limitation for successful tumor targeting and drug delivery. Therefore, a strong effort has been devoted so far to optimize these nanoparticulate systems (size, shape, or surface charge). It is currently admitted that nanoparticles with sizes below 200 nm are more likely to escape from macrophage recognition.

Moreover, nanoparticles with high positively or negatively charged surfaces (≥±25 mV) have a better stability in physiological media, due to repulsive forces. Strongly repelled NPs have fewer tendencies to aggregate and are easily re-dispersed once injected. Many strategies have been proposed to modify the surface of the nanoparticles with biomolecules to reduce protein adsorption such as PEGs. It was proved that the nanoparticles with a high PEG coating density exhibit a longer circulation time in the bloodstream ("stealth" effect).

In particular, some of us investigated the biodistribution and pharmacokinetics of the MIL-100(Fe) NPs after intravenous administration in rodents (rats and mice). Such biodegradable nanoparticles of mesoporous iron carboxylates have a high potential in biomedicine. They exhibit no toxic effect after intravenous injection up to a high dosage of nanoMOFs. However, MIL-100(Fe) NPs tend to aggregate once in physiological media, and after intravenous administration, it was shown that these nanoparticles massively accumulate in the lungs and to a lower extent in the liver and spleen. Recently, some of us took benefit from this to use nanoMOFs passive lung targeting to treat lung metastasis. However, if one targets other types of cancer, this is considered as a main disadvantage.

Based on these observations, this thesis ambitioned to develop new engineered nanoMOF carriers that do not suffer from a passive lung targeting while allowing in vivo stealth properties to avoid liver and spleen capture. water, presumably due to the presence of a high concentration of ions and/or proteins. The best colloidal stability in all media was obtained with β-cyclodextrin-phosphate-polyethylene glycol coated nanoparticles. This was probably due to the concomitant presence of phosphates enabling a highly negative surface charge as well as the PEG brushes that would minimize through steric hindrance a two strong aggregation of the NPs.

The crucial part of this PhD thesis was in vivo study of coated nanoparticles, studied in mice.

First we observed that the various coatings had a slightly different effect on overall biodistribution and pharmacokinetics. The best results in the reduction of the lung nanoparticle accumulation were observed for MIL-100(Fe) NPs coated with β-cyclodextrin-phosphatepolyethylene glycol. This effect is once again likely due to the high negative charge of the phosphate CD, which makes these nanoparticles exhibit a better stability in physiological media.

However, whatever the coating, the engineered nanoparticles all highly accumulated in the liver and spleen resulting in a drastic change in overall biodistribution. This indicated that these coated nanoparticles if they are small enough to prevent from being clogged in the lungs, are still too big once in blood to escape from macrophage recognition.

Therefore, the next objective was to develop nanoparticles with a smaller initial size (in water).

Thus, nanoparticles, even after aggregation in blood, might still be small enough to prevent from macrophage recognition. Several synthesis parameters were adjusted, including the addition of the modulator, or the use of microwave irradiation. Slightly smaller nanoparticles (22-34 nm, compared to 40 nm standard synthesis) were obtained (microscopy) with a smaller size in the aqueous conditions (below 80 nm). Thus, after further surface modification and using more diluted conditions for coating, one might expect that such NPs could have a higher chance to significantly minimize their aggregation in the blood and thus exhibit stealth properties.

Moreover, we successfully scale-up the new patented room-temperature synthesis of MIL-100(Fe) NPs.

In the last part of the thesis, we focused on the encapsulation of the anticancer drug Methotrexate The study of the MIL-100(Fe)@γFe 2 O 3 systems was previously initiated by some of us through decorating the outer surface of MIL-100(Fe) NPs with maghemite (γFe 2 O 3 NPs). This conferred nanoMOFs with a very high r2 relaxivity, comparable to the commercially available systems.

Moreover, the imaging properties of these NPs were further confirmed in vivo by T2 weighted MRI (mice). To avoid any potential leaching of the USPIO upon MOF degradation that occurs rapidly in vivo, another approach was addressed where core-shell MIL-100(Fe)@maghemite NPs produced in situ (Dr. Sene), with maghemite as a core and MIL-100(Fe) as a shell. We succeeded in loading MTX with a high encapsulation rate (up to 36 wt%) into these nanoparticles. The release profile of MTX in PBS was similar to the one of MIL-100(Fe) alone.

However, a precise quantification of the release through other techniques is still required, which will be soon carried out by another team member. The next steps of this project shall include in vitro and in vivo experiments of MTX@MIL-100(Fe)@γFe 2 O 3.

For the second potential application, we collaborated with the CeleScreen start-up whose objectives are to provide cosmetic, or pharmaceutical companies to simplify and cheaper toxicity tests using C. elegans worms to assess the toxicity of their novel ingredients. This method may surpass simple in vitro experiments, as well as overcome limitations (e.g. strict law) with animal testing (rodents). First, we verified that MIL-100(Fe) was totally friendly to the worms through first tests with the worms, causing neither any toxicity, nor altering their reproduction cycle, indicating MIL-100(Fe) was a very promising candidate. Then, the MTX was successfully encapsulated into the MIL-100(Fe) NPs. MTX@MIL-100(Fe) could successfully inhibited the larvae hatching of the worms. These satisfactory results were further confirmed to be the effect of encapsulated MTX, in comparison with no toxicity for the free MTX solution. In terms of perspectives, other drugs are to be considered for encapsulation into MIL-100(Fe) or other non toxic MOFs. Besides, surface modification of MIL-100(Fe) has been considered to achieve a better colloidal stability and minimize aggregation in NGM agar, which may allow an easier egg calculation. Moreover, decorating nanoMOFs with bacteria (worms' food) might also increase the consumption of MIL-100(Fe) by worms and thus increase the effect of the encapsulated drug. (MIL stands for Materials from Institut Lavoisier). Its nanoparticles (nanoMOFs) can load a high amount of challenging drugs with a controlled release in physiological conditions. However, these nanoMOFs tend to aggregate in physiological media and accumulate in organs after intravenous administration. Passive lung targeting of these NPs was previously used for successful lung cancer treatment in mice. However, when the distribution in other organs or tissues is needed, this NP behavior may be considered as a drawback.

In this work, we focused on MIL-100(Fe) NPs. We studied the impact of different surface modifications routes (covalent and non-covalent) on the stability of NPs in physiological media (to mimic in vivo conditions) and the ability of the coatings to tune the in vivo biodistribution and pharmacokinetics of NPs after intravenous administration in mice. We selected various surface modifications strategies: PEG (NHS route), Graftfast PEG, β-Cyclodextrin-Phosphate and β-Cyclodextrin-Phosphate-PEG, which were already reported and showed positive in vitro behavior, but were never studied in vivo. Particularly, β-Cyclodextrin-Phosphate-PEG coated NPs showed the best effect on reduction of lung NPs accumulation. However, the liver and spleen accumulation was higher than uncoated NPs together with a short circulation in blood.

These NPs have strong negative surface charge, thus they tend to repel each other and do not form aggregates. Thus we can conclude that β-Cyclodextrin-Phosphate-PEG modified NPs in blood are small enough to escape accumulation process in lungs. However, they could not escape macrophage recognition and fast elimination from the bloodstream by the reticulo-endothelial system (RES) organs, such as liver or spleen, probably due to the still too large particle size after intravenous administration and a not enough effective coating. Moreover, we put some effort to were dissolved directly with a glucose 10% aqueous solution (1.6 ml) and 3 H-Gemcitabine monophosphate (8 l, 16 Ci) was then added to the mixture.

Surface modification of 14 C-MIL-100(Fe)@ 3 H-Gem MP

When considering the need of 6 mg of nanoMOF to be injected within a single day, 8 mg of nanoMOFs were prepared; an additional amount was used for analysis of injected formulations plus a slight excess in case of problem during injections.

In vivo protocols

Intravenous injection via tail vein

NPs formulations were intravenously injected in the mouse by lateral tail vein. For an easier injection mice were immobilized using a cylindrical restraint tube (see Figure A1-1). This kind of tube is appropriate for conscious mice or rats, except for hairless or nude animals. For such type of tubes, only the tail is dangling outside the restraint tube, which make injections easier. It is possible to rotate the tube for the best position of the tail for the injection. In case the tube is fixed to the platform, it is also possible to change the position of the mouse within the tube or slightly rotate the tail. The key parameter for a successful injection is to worm the tail of the mouse, to ensure a vasodilation of the tail vessels. This made blood vessels on the tail dilate and become better visible. Injections could be done through the lateral tail veins, which are located on the sides of the tail. The vessels are very thin and superficial, thus injections typically require some special skills. The needle should slide into the tail parallel to the tail. When the needle is in the right position inside the vein, injection was very easy, then no resistance was observed.

Material could be then injected slowly, with a constant motion, to avoid any rupture of the vessels. We could also observe the flow of the injected formulation, through a change of the color of the vein, from dark blue to light, pale color, to almost transparent. This effect was observed as the blood was pushed out by the injected formulation. Another proof of successful injection was the blood droplet observed after injection, at the injected point.

Blood and organ collection

After a given time, blood was collected from conscious mice via facial vein by the puncture (see Figure A1-2). A collection of around 200 µl of blood was possible from one mouse, which was next stored with the addition of heparin to prevent coagulation. Samples were centrifuged and stored in a freezer for analysis (-20°C). Before the second blood collection, the solution of Dolethal (sodium pentobarbital) was injected intraperitonealy. A few minutes after injection mice lose consciousness. Then, the second blood collections were done by heart puncture, by opening the chest of the mice and collecting the blood directly from the heart with syringe and needle. Blood was then stored in eppendorf tubes with the addition of heparin. After blood collection from the heart, we could collect organs for further analysis such as lungs, liver, spleen, kidneys, heart, fat and muscles. Fat and muscles were collected only at 6 h after injection, as the additional organs for a better understanding of the overall biodistribution.

samples and an additional 0.1 ml of solution of hydrogen peroxide 30% was added. Then, samples were well mixed. This process was repeated 4-5 times. Samples were kept with open caps for removal of the gas bobbles. Samples were then placed again in the oven for another 1 h at 60°C and after being cooled down, 10 ml of solution of "Ultima Gold" was added to the blood samples. Samples were finally well mixed and after 10 minutes considered for scintillation counting.

Preparation and characterization of the urine 1 ml of urine was mixed with a 10 ml of solution "Ultima Gold" and well mixed. Samples were then analyzed by scintillation.

Preparation and characterization of the feces

Feces were dried and frozen before characterization. Approximately 20 mg of feces were introduced into 20 ml plastic vials and then soaked with 0.2 ml of water during 30 minutes. Next, 1 ml of solution "Soluene-350" was added to the samples and were kept in the oven for 1-2 h at 50 C. Then, 0.5 ml of propan-2-ol was added to the samples and kept in oven at 50°C for another 2 h. Once the samples were cooled down, 0.2 ml of hydrogen peroxide 30% was added dropwise to the samples and gently mixed. Samples were kept open for approximately 10 minutes to enable gas evacuation. Finally, 10 ml of solution "Hionic-Fluor" was added to the samples and well mixed. After 10 minutes the samples were analyzed by scintillation counting.

C. elegans tests

During the tests with C. elegans, nanoparticles (1 mg) were mixed with a food source of worms (E. coli OP50, 1 mg in 50 µl). This mixture was then added to the plates with NGM agar to feed worms. NGM is the medium commonly used in worms' experiments, containing: NaCl, peptone, ) and/or constitutive organic ligands (carboxylates, phosphonates, azolates, etc.) leading to thousands of MOFs with unique features [5][6][7][8] (and references therein). MOFs exhibit therefore highly porous structures that span over a large range of pore sizes (micro-or mesopores) or pore shapes (cages, channels, etc.), and possess either rigid or flexible frameworks. One can also further tune their polar/apolar character through the use of polar or apolar organic functionalities, often carried out through direct synthesis or postsynthesis modification on the ligand or grafted on the metal sites, which strongly impacts the sorption properties of the solids. [9] As a consequence, a large number of potential applications of MOFs have been proposed to date such as gas adsorption/storage or separation, [10][11][12] catalysis, [13,14] energy, [15][16][17][18] optical properties, [19] sensing, [20][21][22] and biomedicine, [23,24] among others. [8] Downsizing materials to the nanoscale is also a suitable method to tackle down new applications compared to the ones of their bulk analogues, [25][26][27][28][29][30][31][32][33] such as in biomedicine, whereas applications of nanoparticles of metal-organic frameworks (nanoMOFs) is a rapidly developing topic of interest. Nanomaterials, due to their smaller particle size, can indeed improve the drug delivery performances for the treatment of several diseases. [24,34] Moreover, the high and regular porosity and the unique combination of well-dispersed metal sites and organic groups within the framework of nanoMOFs, combined with the low toxicity of polycarboxylic acids and selected metals (Fe, Zn, Ca, etc.), make these porous solids appealing nanocarriers for the controlled release of drugs. [9] Considering the drug delivery requirements, the best nanovectors are the ones that fulfill the following conditions [35][36][37][38][39][40] : 1) high drug entrapment (payload and efficiency), 2) controlled drug release without "burst" effect, 3) ability to target diseased cells and tissues in a highly selective manner, 4) lack of toxicity through a progressive degradation and the absence of accumulation in the body, 5) an easy engineering of the outer surface of the nanoparticles (NPs) for an improved in vivo stability and/or biodistribution (the distribution and accumulation in the different organs and tissues), and 6) possibility of NPs to be detected by imaging techniques.

5
In the past few years, numerous studies have demonstrated the great potential of nanoMOFs at the preclinical level for biomedical applications. Many of them were reported very recently In the past few years, numerous studies have demonstrated the great potential of nano particles of metal-organic frameworks (nanoMOFs) at the preclinical level for biomedical applications. Many of them were reported very recently based on their bioactive composition, anticancer application, or from a general drug delivery/theranostic perspective. In this review, the authors aim at providing a global view of the studies that evaluated MOFs' biomedical applications at the preclinical stage, when in vivo tests are described either for pharmacological applications or for toxicity evaluation. The authors first describe the current surface engineering approaches that are crucial to understand the in vivo behavior of the nanoMOFs. Finally, after a detailed and comprehensive analysis of the in vivo studies reported with MOFs so far, and considering the general evolution of the drug delivery science, the authors suggest new directions for future research in the use of nanoMOFs for biomedical applications.

Introduction

Metal-organic frameworks (MOFs) or porous coordination polymers were first reported in the late 1980s [1] and this domain started to expand continuously since the end of the 1990s. [2][3][4] The structure and composition of these crystalline hybrid solids can easily be tuned through the almost infinite possible Adv. Mater. 2018, 1707365 based on their bioactive composition, [41] anticancer application, [7,[42][43][44][45][46][47][48][49] or from a general drug delivery/theranostic perspective. [50] In this review, we aim at providing a global view of the studies that evaluated MOFs' biomedical applications at the preclinical stage, when in vivo tests are described either for pharmacological applications or for toxicity evaluation (Table 1). Of note, most of the nanoMOFs reaching a preclinical in vivo evaluation are those based on Fe carboxylates or Zn azolates. We first describe the current surface engineering approaches that are crucial to understand the in vivo behavior of the nanoMOFs. Finally, after a detailed and comprehensive analysis of the in vivo studies reported with MOFs so far, and considering the general evolution of the drug delivery science, we suggest new directions for future research in the use of nanoMOFs for biomedical applications.

Surface Modifications of NanoMOFs

So far, the great number of studies describing the potential of MOFs to be used for pharmaceutical applications contrasts with the reduced number of in vivo pharmacological efficacy studies performed to date. [51] There are several reasons why most of the physicochemical and in vitro studies dealing with MOFs in the biomedical field have not been systematically evaluated in vivo yet. Some are related to MOFs' design development itself, such as colloidal stability issues and/or lack of organ or tissuetargeting properties. [51] These challenges are already being overcome by several surface coating strategies. Similarly to what has already been demonstrated for other nanocarriers, it is possible to tune pharmacokinetics and biodistribution by incorporating motifs onto the outer surface of the nanoMOFs, and the aim of this section is to give an overview of the latest advances in this field (the most recent works are summarized in Table 1).

First, it must be noted that for isotonicity, intravenous (i.v.) administration of nanomaterials needs dispersion in an isotonic solution such as 0.9% saline or 5% glucose. Many nano-MOFs, particularly once dried, cannot be dispersed yet in these media and/or are not stable enough and tend rapidly to form aggregates, making i.v. administration not possible due to risk of embolization. In some cases, nanoMOFs could nevertheless be administered at (low) doses at which particle aggregation does not occur. For instance, Baati et al. did not observe any stability problems for low doses of iron(III) carboxylates MIL-100(Fe), MIL-88A(Fe), and MIL-88B_4CH 3 (Fe) (MIL stands for Materials of Institut Lavoisier) when dispersed in glucose solution. [52] But colloidal stability of nanoMOFs still remains a major issue that makes difficult in vivo administration. [53] Selective functionalization of the external surface of nanoMOFs is thus required in view of their biomedical applications.

Apart from the colloidal stability, one major objective of the surface engineering of nanoMOFs is to also allow blood long circulating properties (i.e., "stealthness") of the NPs (Figure 1). Indeed, surface functionalization of NPs with hydrophilic polymers reduces opsonization, a process that involves the interaction between some proteins from the blood and the NPs, resulting in their recognition and elimination by the macrophages to the liver, spleen, and the bone marrow (i.e., the so-called reticulo-endothelial system). [51] Surface modification strategies of NPs and liposomes have indeed been proposed several years ago and polyethylene glycol (PEG) is without any doubt the best candidate. The PEGylation of polymeric NPs was first proposed in 1994 by Gref and co-workers [54] and a great number of studies dealing with PEGylation of other types of NPs have been carried out since that time. [55,56] strategies have been proposed so far to modify the surface of nanoMOFs using a large variety of molecules, such as nucleic acids, [62,63] polymers, [44,[64][65][66] cyclodextrins (CDs), [61] or lipids, [67,68] resulting into various degrees of improvement of the particles properties. One can classify these approaches into three categories: covalent, [61,63,64,[69][70][71][72][73][74][75][76][77][78][79] noncovalent, [59,66,[80][81][82][83][84][85] (Table 2) or core-shell, [20,60,67,[86][87][88][89][90][91][92][93] as detailed below.

Covalent Route

Recently, Zimpel et al. have reported the covalent attachment of functional polymers, such as amino-PEG and a derived oligoamino amide named Stp10-C and constituted by two terminal groups, a primary amine and a thiol, connected via a repetitive diaminoethane motif with proton-sponge characteristics, grafted on the external surface of the mesoporous biocompatible iron carboxylate MIL-100(Fe) NPs (Figure 2). [64] This resulted in a successful improvement of the colloidal stability of the functionalized NPs together with satisfactory cellular uptake of MIL-100(Fe) NPs functionalized with Stp-10C by murine neuroblastoma (N2A) cells. Authors showed nevertheless a negative impact of the polymer coating on MIL-100(Fe) magnetic resonance imaging (MRI) properties. Concerning PEG coating, no information was provided about its influence on macrophage uptake and thus it is not yet possible to conclude about the stealth properties of this system. [64] Agostoni et al. reported the use of cyclodextrin-phosphates (CD-P) covalently attached to the outer surface of MIL-100(Fe) NPs through direct coordination of the phosphate groups to the Lewis iron(III) sites from the outer surface of the NPs. Cyclodextrin molecules being bulkier than the microporous windows of MIL-100(Fe) thus avoided any penetration inside the matrix of the nanoMOFs. By this method, 17 wt% of CD-P was found covalently attached to the nanoMOFs after 24 h of incubation. Upon surface functionalization, the phosphate groups led to a more negative surface charge (ζ -17 mV against -35 mV for uncoated and coated NPs, respectively) and thus a better electrostatic stabilization and a reduced aggregation of the NPs. [61] Moreover, authors demonstrated the good stability of CD-P coating in aqueous solution, where three washings did not lead to any CD-P leaching, while three washings with phosphate buffer saline (PBS) led to only 7% CD-P detachment. Similarly, it was shown that less than 10% of the total CD-P coating was released from the nanoMOFs after 24 h of incubation in PBS Adv. Mater. 2018, 1707365 [64] Copyright 2016, American Chemical Society. or in cell culture media. Loading and release of the antiretroviral azidothymidine-triphosphate was also not affected by MOF coating; in all cases drug cargo was almost entirely delivered after 24 h, under physiological simulated conditions (PBS, 37 °C). Importantly, the coating with CD-P did not cause any significant toxic effects on the cell lines (J774, MCF7 and LP-1). In addition, cellular uptake of MIL-100(Fe) NPs functionalized with mannose-bearing CD-P derivative by the human retinoblastoma cell line Y79 exhibited more than twice higher penetration inside cells as compared to uncoated NPs. [61] The benefit of such a covalent coating in vivo remains however still to be demonstrated. Morris and colleagues, took advantage of the click chemistry route to covalently functionalize the microporous zirconium carboxylate UiO-66-N 3 NPs with oligonucleotides, [62] by a strain promoted click reaction between DNA append with dibenzylcyclooctane and azide-functional UiO-66-N 3 . As a result, colloidal stability in NaCl was meaningfully improved and cellular uptake (HeLa-human cervical cancer cells) was significantly higher, whereas cell viability was not altered.

A more sophisticated system was developed by Wang et al. who prepared surface modified by the mesoporous iron(III) amino-benzenedicarboxylate MIL-101(Fe) NPs loaded with Dox. After the synthesis of the NPs, to avoid premature drug release through systemic circulation, CD and PEG chains were added by a one-pot, and organic solvent-free "green" postsynthetic procedure based on click chemistry and host-guest interactions forming PEG-RGD-β-CD-SS-MIL-101(Fe) NPs. [69] This led to a better stability in PBS and prevented from fast degradation compared to uncoated particles. Surface-modified nanoMOFs presented negligible uptake in α v β 3 integrin negative noncancerous COS7 (African green monkey kidney) cells, while NPs were internalized in α v β 3 integrin expressing HeLa (human adenocarcinoma) cells. In vivo antitumor efficacy in hepatoma H22 tumor-bearing mice was then demonstrated through tumor inhibition effect for coated NPs loaded with Dox and for free Dox, while no side effects were observed in coated NPs. Nevertheless, the main inconvenient lies here in the lack of demonstration of the effectiveness of the coating since the NPs were locally administered by subcutaneous injection, near the tumor.

Schmitt et al. have reported functionalized and hierarchically structured MOFs embedding magnetic core particles (mag-MOFs) through a layer-by-layer synthesis route. [73] Core/multishell particles were obtained following several steps: i) MOF growth of Cu(BA-TPDC) (BA-TPDC-bis(azidomethyl)-terphenyldicarboxylic acid) around a magnetic core, ii) click reaction of blue dye, iii) MOF growth of Cu(TPDC) (TPDC-terphenyldicarboxylic acid), iv) MOF growth of Cu(BA-TPDC), and v) finally a click reaction of the red dye was performed. In the next step, the SURMOF final product was converted into dye-loaded SURGEL (the surface-grafted gels) capsules around the magnetic core. The advantage of SURGELs results from the possibility to control the release of the dye molecule, depending on the pH of the environment. The drug loading of the MOF inner layers was carried out followed by a coating with a protective polymer layer. The resulting magMOF particles could be converted to magGEL by ethylenediaminetetraacetic acid (EDTA) treatment. Different kinetics of release were reported as a function of the pH with a maximum of release reached at pH 11 after 2 h. This product was considered for the oral administration of drugs [73] with however neither in vitro nor in vivo assays to assess the benefit of this strategy. In the same study, the growth of SURMOF films on gold-coated mica substrate was described followed by their conversion into SURGELs via click chemistry with arginineglycine-aspartic acid (RGD) to favor cell adhesion via specific interaction with the integrin receptors of the cell membrane. GRD is well known to be able to induce adhesion of osteoblasts onto implanted surfaces and to improve bone formation. The proof of concept has been performed in vitro using osteoblastlike CAL72 cell line. The cells were seeded onto the functionalized SURGEL substrates. Finally, no toxicity was observed until 24 h and more adherent cells were observed onto functionalized SURGELs. Additionally, microfluidic shear force assay was performed which confirmed that CAL72 cells interacted stronger at the surface of functionalized SURGELs as compared to the nonfunctionalized counterpart. [74]

Noncovalent Approach

Chen et al. prepared an intrinsically radioactive microporous Zr dicarboxylate MOF UiO-66 ( 89 Zr-UiO-66) functionalized with pyrene-derived polyethylene glycol (Py-PGA-PEG) taking advantage of "click modulation." [59] The grafting on the external surface of the NPs was ensured here via strong π-π interactions between the organic moieties of UiO-66 NPs and the pyrene molecules. This strategy was also used to further attach a tumor-targeting ligand named F3 on the surface of the MOF (UiO-66/Py-PGA-PEG-F3). Additionally, NPs were loaded with Dox for the drug delivery experiments. Noteworthily, the kinetics of nanoMOFs' degradation and release of the drug were slowed down, probably due to partial blocking of the pore windows of the MOFs by the polymer, hampering the diffusion of phosphates from PBS into the pores. Moreover, drug release kinetics was dependent on the pH of the environment, and once NPs arrived at the extracellular region of the tumor, release of cargo was sped up. Despite these facts, burst release was not totally avoided with around 19% of Dox delivered within 30 min. Regarding cellular experiments, in vitro tumor cell uptake performed in the MDA-MB-231 triplenegative breast cancer cell line and L929 fibroblast showed higher cellular uptake in the MDA-MB-231 cell line compared to the L929 cell line. Empty NPs did not show any cytotoxic effect on MDA-MB-231 cells, while NPs loaded with Dox succeeded to inhibit MBA-MB-231 cell growth. Toxicity of UiO-66/Py-PGA-PEG was studied in Balb/c mice through i.v. administration of 10 and 50 mg kg -1 NPs. After 7 and 30 d no sign of toxicity was reported based on histological examination and evaluation of biochemical parameters. Finally, authors evaluated biodistribution by in vivo positron emission tomography imaging in a xenograft subcutaneous tumor model in mice and showed that 89 Zr-UiO-66/Py-PGA-PEG-F3 was faster accumulated in MDA-MB-231 tumors compared to 89 Zr-UiO-66/Py-PGA-PEG demonstrating that F3 ligand caused active targeting. Note that for further experiments a reduction on the size of the nanoMOFs would lead to better efficacy results since the large size of coated NPs (≈250 nm) prevented from effective internalization into tumor vasculature cells. Furthermore, blood circulation halflife of less than 2 h was reported suggesting that surface PEG density was not efficient to provide sufficient stealth properties Adv. Mater. 2018, 1707365 to the nanoMOFs. In a first attempt to prepare MOFs for oral administration, MIL-100(Fe) NPs were coated by the bioadhesive polysaccharide chitosan (CS). [80] The surface-modified NPs exhibited an improved chemical and colloidal stability under oral simulated conditions. In vitro, permeability through a model of intestinal barrier and cytotoxicity were improved, in comparison with the noncoated nanoMOFs. The viability and integrity of the intestinal barrier were investigated using an in vitro model of polarized Caco-2 monolayer cells. Systemic and mucosal immune responses were also studied, through complement activation tests and by cytokine profile, resulting in the absence of any complement activation for both coated and uncoated NPs, while cytokine production decreased from one to two orders of magnitude for the CS-coated NPs. [80,81] Another noncovalent approach was reported by Bellido et al. through heparin coating of MIL-100(Fe) NPs. [82] The colloidal stability of these NPs was improved in water and PBS, while reduced cell recognition was also observed in vitro for the heparin-coated nanoMOFs using a macrophage cell line (J774.A1) at short time of incubation (up to 4 h), without any complement activation and reactive oxygen species production. Furthermore, surface modification preserved NP encapsulation capacities, as demonstrated with the active caffeine, while a decrease in the release kinetics compared to uncoated NPs was observed. However, the eventual benefit of these NanoMOFs in terms of pharmacokinetics and biodistribution still needs further demonstration.

Another example of noncovalent approach to decorate MOF NPs was developed recently by Cai et al. who successfully developed surface engineered MIL-100(Fe) with hyaluronic acid and indocyanine green with improved colloidal stability and good cellular uptake by cancer cells using the CD44-positive MCF-4 cell line. [66] Qu and co-workers described the ability of noncovalent functionalized iron carboxylate MIL-101(Fe) with unmethylated cytosine-phosphate-guanine oligonucleotydes (CpG ODNs) (MIL-101(Fe)-CpG nanoconjugates) to enhance the immune response. In addition, T 2 -magnetic response imaging ability was tested in vitro and in vivo. [85] CpG ODNs were adsorbed onto the MOFs by π --π interactions between the CpG ODNs and the terephthalic acid organic ligands. Regarding cellular experiments, the obtained nanoconjugates did not show any toxicity on RAW264.7 cells, even at rather highest concentration (200 µg mL -1 ). After cell internalization, the nanoconjugates interacted with TLR9 and triggered the secretion of cytokines. Furthermore, MIL-101(Fe)-CpG nanoconjugates displayed a higher immune response comparatively to CpG ODNs alone, both in vitro and in vivo. Moreover, T 2 -weighted MR images of the tumor-bearing mice, before and after subcutaneous injection of MIL-101(Fe)-CpG nanoconjugates, showed a strong signal at the site of injection, whereas the tumor position was well visible. Due to T 2 -MRI ability of the nanoconjugates, it was proposed to use these nanoconjugates to track the labeled immune cells and to monitor in vivo the CpG-ODN-based drugs or vaccines. [85]

"Core-Shell" Strategies

Besides covalent and noncovalent surface modification, another strategy related to the so-called "core-shell" approach consists in the covering of the nanoMOF core by a dense coating, identified as the shell of the system.

Rieter et al. were the first to consider silica coating to form core-shell systems made of nanoMOFs to avoid a too fast degradation of the NPs in body fluids. [20] Further functionalization of silica-coated NPs was also performed using diopicolinic acid. Authors described a general method to obtain variable thickness of silica shells on the lanthanide dicarboxylate Ln(BDC) 1.5 (H 2 O) 2 , where Ln = Eu 3+ , Gd 3+ , or Tb 3+ , and BDC = 1,4-benzenedicarboxylate. They also loaded Pt-based drugs into nanoscale coordination polymers, constructed from Tb 3+ ions and DSCP (c,c,t-(diamminedichlorodisuccinato)) Tb 2 (DSCP) 3 (H 2 O) 12 , stabilized with silica shell. [86] More recently, Wuttke et al. reported a "core-shell"-like approach, where nanoMOFs were encapsulated into a lipid bilayer shell. [67] They demonstrated that MIL-100(Fe) and the chromium dicarboxylate MIL-101(Cr) NPs, coated with a lipid bilayer consisting of (1,2-dioleoyl-snglycero-3-phosphocholine), exhibited a better colloidal stability than the uncoated nanomaterial. Noteworthily, the integrity of the lipid bilayer was confirmed and led to a higher uptake by cancer cells (T24 bladder carcinoma cells), without cytotoxic effect of empty-coated NPs. The same team reported very recently the coating of MOF NPs with exosomes via lipid fusion (Figure 3). [87] This method ensured the preparation of an exosome delivery system with unprecedented loading efficiency. The resulting exosomecoated NPs showed no burst leakage with an efficient release of their cargo into the cells.

Another core-shell strategy for surface modification of MOFs was proposed by Li et al. [60] A biomimetic theranostic oxygen (O 2 ) meter (cancer cell membrane@Pt(II) porphyrinic-nanoMOF (mPPt)) was constructed for cancer targeting and phosphorescence image-guided photodynamic therapy (PDT). Pt(II) porphyrinic nanoMOF was formed by O 2 -sensitive Pt(II) meso-tetra(4-carboxyphenyl)porphyrin and Zr 6 clusters. Cancer cell membrane (from 4T1 cells) was selected as a target for MOFs' surface modification, to increase cancer targeting due to homotypic targeting and immune escape abilities. Moreover, mPPt were loaded with photosensitizers. This led to an increase in particle size from 108.5 to 150.5 nm while the ζ potential decreased from 24.5 to -28.5 mV for the coated nanoparticles in comparison with the bare ones. Singlet oxygen ( 1 O 2 ) production ability, as the main cytotoxic species associated with PDT, was evaluated in vitro in 4T1 cells under various O 2 atmospheres.

Adv. Mater. 2018, 1707365 Figure 3. Schematic illustration of the cell uptake of exosome-coated nanoMOFs and release mechanism of the cargo. Reproduced with permission. [87] Copyright 2017, American Chemical Society.

An increase of O 2 levels was observed in cells with mPPt and the highest 1 O 2 production in 21% O 2 atmosphere. The cell uptake was determined after incubation with 4T1 cells and heterogeneous COS7 normal cells, as the controls. mPPt were easily recognized by 4T1 cells, which was not the case with COS7 cells. On the other hand, the escape of mPPt from RAW264.7 murine macrophages was demonstrated, whereas uncoated NPs were captured by macrophages. In addition, in vitro PDT led to a significant cytotoxicity against 4T1 cells in 21% O 2 , under O 2 sensor. Fast and accurate response of mPPt toward O 2 fluctuation was also observed by in vivo imaging, and an improved anticancer activity was noted after treatment with mPPt of BALB/c mice with subcutaneously injected 4T1 cells. [60] Wang and colleagues also reported nanoscale polymer-MOF hybrids, where the Zr terepthalate UiO-66 was coated with polyaniline (PAN) (UiO-66@PAN) for phototermal therapy (PTT). [91] The photothermal performance of PAN-coated NPs was investigated and the temperature raised until 57.2 °C at a concentration of 100 µg mL -1 UiO-66@PAN upon laser irradiation, which was sufficient for the efficient killing of malignant cells. Regarding cellular assays, coated NPs did not induce any cytotoxicity in both murine colon cancer CT26 and human colon cancer HTC116 cell lines, but once irradiated, nearly 70% cells were dead. In vivo tests carried out with both UiO-66@PAN and NIR irradiation showed a complete tumor regression after 10 d, when compared to the controls (untreated or NPs alone), proving that UiO-66@PAN was a good candidate for PTT-based cancer therapy. [91]

Toxicity: From In Vitro to In Vivo Evaluation

There is to date an increasing number of publications demonstrating the lack of significant in vitro toxicity of nanoMOFs based on 2D cell studies using established cell lines. However, these in vitro tests are far from reproducing the in vivo situation. Thus, it is important to emphasize that the lack of in vitro cytotoxicity does not mean that the nanodevice is safe and biocompatible. In general, toxicity issues should be tackled down based on a more rational safe-by-design approach when considering nanoMOFs for biomedical applications. Most of the authors justify the choice of their nanoMOFs based on their constitutive parts, i.e., through the reported metal and ligand individual toxicity. This strategy can be useful mainly to discard the most toxic nanoMOFs, but it is far from being optimal. The data that we possess about metal toxicity refer mainly to medial oral lethal dose, that is the amount of a drug or other substance that, when orally administered to a group of experimental animals, will kill 50% of the group in a specified time (LD50), and to human exposure to metals present in the environment (water, air, soils, and food). Administration of nanoMOFs in the human being as part of a pharmacological treatment implies active penetration of the metal in the organism. In this case, exposure to metals is different than the previously studied ones and will be determined by the intensity (dose), frequency, and duration of the treatment, as well as by the administration route that determines how the metal is distributed, accumulated, metabolized, and eliminated. Certainly, LD50 value cannot be extrapolated to other administration routes for which biodistribution will be totally different. Furthermore, when one deals with in vivo human toxicity, there is even less data available for most of the constitutive MOF ligands, such as terephtalic acid in MIL-101 or UiO-66, whose toxicological properties after i.v. administration have not been thoroughly investigated.

As a consequence, to obtain more information about potential toxicity of the nanoMOFs after administration into humans, one shall focus more on reports dealing with in vitro or ex vivo models that specifically mimic an in vivo situation rather than relying on simple 2D in vitro toxicity tests using established cell lines (Figure 4). [94] In many cases the cytotoxicity is only performed in cancer cell lines to demonstrate a potential anticancer effect, and not in healthy cells to assess the toxicity of the carrier, [62,95,96] even if Tamames et al. already demonstrated that toxicity of numerous MOFs is higher for healthy cells than for tumor cells. [97] Besides, when healthy cells are used the choice of the cell type is rarely justified based on therapeutic applications. [98] Some efforts have already been made to design more pertinent cytotoxic in vitro studies by Wuttke et al. who explored the toxicity of various nanoMOFs using different primary healthy cell types depending on the proposed medical application. [99] For their use as drug delivery systems, toxicity and inflammatory response on vascular cells and lung cells were tested, whereas for their application as drug delivery implant coatings of dental or nerve guidance tubes, fibroblasts and neural cells were chosen, respectively. Furthermore, alternative models to screen toxicology, such as numerous 3D in vitro models, are currently being developed to better reproduce the in vivo conditions of a particular tissue, such as the tumor environment, the liver, or the skin. [100][101][102][103] These new models were also reported by Wuttke et al. where the biological response of sensory neurons to the nanoMOFs was monitored using rat neonatal organotypic dorsal root ganglion cultures. [99] It is expected that in a near future these 3D models will replace the monolayer cell cultures to better investigate eventual adverse effect of nanoMOFs.

Concerning the in vivo toxicity tests of nanoMOFs, a very few studies have been performed to date. The first study describing toxicological information in vivo of nanoMOFs was reported by some of us in 2010. [24] In this pioneering experiment, iron carboxylate NPs such as MIL-88A NPs with particle sizes of 150 or 500 nm, as well as MIL-88B_4CH 3 NPs with 50 or 140 nm, were intravenously administered into Wistar rats via the jugular vein. Animals were followed up to 3 months and several para meters were evaluated such as animal behavior, animal weight evolution, weight changes of different organs, cytochrome P-450 activity, ALT and AST transaminases levels, and interkeukine-6 serum concentration. The only significant effect was a slight transitory increase in the spleen and liver weights, attributed to the fast sequestration by the reticuloendothelial organs of the nanoMOFs that went back to normality 1-3 months after injection. The absence of immune or inflammatory reactions after NP administration supported their lack of toxicity. Moreover, the absence of activation of cytochrome P-450 suggested a direct excretion of the polyacids, in agreement with their high polarity.

In 2013, some of us investigated the in vivo toxicity and biodistribution of three different uncoated iron carboxylate-based nanoMOFs, namely, MIL-100, MIL-88A, and MIL-88B_4CH 3 , using Wistar rats. [52] First, i.v. administration of increasing doses of these NPs was performed to assess lethal dose 10 (LD10). Animal behavior during 7 d was analyzed and then animals were sacrificed and organs harvested for histological examination. This early study led to very promising results. Indeed, the LD10 was never reached with the studied nanoMOFS, and any mortality was caused by nanoMOF administration or significant toxicity signs were observed. In fact, the highest practicable dose was established based on colloidal stability and not on toxicity issues. Only an expected increase in the oxidative stress was observed but it came back to the control level after 1 month.

In 2015, a third study systematically describing the in vivo toxicity of a series of nine different nanoMOFs was reported by Ruyra and colleagues. [104] The in vitro cytotoxic test using 2D cell culture was compared to results obtained using the zebrafish embryo in vivo model (Figure 4). On the whole, a strong correlation between the results using both methods was found except for MIL-101(Fe), which was more toxic in vivo in zebrafish than in vitro. Very interestingly, authors concluded that the toxicity of these materials was mainly governed by the release of metal ions during degradation, while they also highlighted the importance of the formation of other species upon degradation. This confirms that when assessing nanoMOF's toxicity more information about degradation mechanisms and degradation products under physiological conditions is required. For a while, it was generally assumed that nanoMOF in vivo degradation resulted in releasing the constitutive metal ion and the ligand. However, this is in contradiction with the findings of Ruyra et al. who have proven that a few nanoMOFs once in contact with cell culture medium became amorphous and underwent structural rearrangements and/or reactions that generated new inorganic species, responsible for adverse effects. [104] In a more recent study, this phenomenon has been described in detail for MIL-100(Fe) nano-or microparticles, confirming the progressive formation of dense phases (iron oxide and phosphate) associated with a degradation under physiological simulated conditions. [105] Besides the above-mentioned studies, a few recent reports dealing with in vivo nanoMOFs' applications also disclosed brief results concerning in vivo toxicity, [59,68,106,107] but it is far from being enough to establish material biocompatibility.

As a conclusion, despite promises, there is still a great effort to be carried out to assess the toxicity profiles of nanoMOFs before eventual translation into the clinic. Until that, toxicity remains without any doubt one of the keystones to continue building the bridge to move from bench to bedside.

An important advantage of MOFs compared to other nanomaterials lies in the preparation of metal biomolecule frameworks (also known as bioMOFs for bioactive MOFs or bioMIL for bioactive MIL) by the incorporation of endogenous molecules or active ingredients as building blocks. [41] This allows reducing the amount of nonactive undesirable compounds to be administered at the benefit of an anticipated decrease of toxicity. Numerous active ligands have been proposed to date to produce bioMOFs: peptides (metal peptide frameworks), [108] nucleobases, [109,110] carbohydrates, [111][112][113][114][115] porphyrines, [48,116] as well as some active ingredients. [117][118][119][120][121][122] Unfortunately, most of these bioMOFs are still at their initial stages of development and a very few of them have undergone in vivo preclinical evaluation. In 2014, Lu et al. performed the in vivo administration of a porphyrine-based bioMOF. [48] Using a subcutaneous xenograft model in mice of human head and neck cancer cells SQ20B, this study was the first in vivo proof of concept of using nanoMOFs for photodynamic therapy purposes. Of note, MOFs were only locally administered at the tumor site and in vivo toxicity was not assessed since the authors only measured anticancer activity. Therefore, the in vivo toxicity associated with this bioMOF still needs to be determined.

In addition to the toxicity associated with the nanoMOF itself, toxicity issues can be intimately related to the synthesis process, especially to the solvents and reaction modulators employed to prepare the NPs. Indeed, the synthesis of most bioMOFs requires the use of toxic solvents such as dimethylformamide or pyridine. [120,121] Thus, to prepare suitable nanomaterials relevant for clinical applications, alternative "green" synthesis routes are urgently needed as it has already been done for other MOFs. [61,69,122]

Biodistribution, Targeting, and Pharmacological Efficacy

The major aim in using nanocarriers for drug delivery is related to the improvement of the pharmacokinetic profile and Adv. Mater. 2018, 1707365 biodistribution to allow a better drug targeting toward diseased cells and tissues. Despite the great number of studies claiming that nanoMOFs are promising drug carriers, a very few of them focus on demonstrating the ability of these systems to modify and improve drug biodistribution.

The first investigation dealing with the pharmacokinetic and biodistribution of nanoMOFs was reported by Baati et al., in which the long-term biodistribution (from 1 to 30 d) of three different uncoated iron carboxylate nanoMOFs was examined in rats (Figure 5). [52] After i.v. administration of MIL-100, MIL-88A, and MIL-88B_4CH 3 NPs, both the iron and the organic linker concentrations were quantified in several complex biological media, including liver, spleen, and urine. Iron levels were quantified by atomic absorption spectroscopy, whereas the linker concentration was determined by specific extraction and high-performance liquid chromatography methods. [123] An important reversible accumulation in the organs of the reticulo-endothelial system, liver and spleen, was observed for all the tested nanoMOFs (Figure 6). However, liver accumulation was higher for the MIL-88B_4CH 3 NPs, whose constitutive linker, i.e., tetramethyl benzendicarboxylic acid, exhibits a more pronounced hydrophobic character, which could also explain the slight accumulation of this nanoMOF in the brain that was not observed for any of the other materials. Besides, even if the observed normal breathing of the animals supported the lack of lung toxicity, upon histological examination of the pulmonary tissue of nanoMOF-treated rats, it was observed that a large amount of nanoMOFs was found to aggregate and to accumulate in the lungs. Taken together, these data demonstrated that nanoMOFs distributed in different tissues for a period of time, and as a consequence they could act as prolonged drug delivery systems. However, this first study was focused on the evaluation of nanoMOFs without any drug loaded inside their pores. Recently, we showed how the encapsulation of the anticancer drug Busulfan into MIL-100(Fe) NPs drastically modified the drug pharmacokinetics profile compared to the commercial formulation Busilvex. [124] Busulfan detected after MIL-100(Fe)-Busulfan administration was much lower than after Busilvex dosage with a mean area under the curve in a plot of drug concentration in blood plasma versus time of 2.6 and 25 mg mL -1 min -1 , respectively. Moreover, it was also demonstrated that the clearance of NPs from the blood was faster in the case of drug-loaded nanoMOFs, probably due to changes on the surface of the particles that became more hydrophobic, and therefore, more easily recognized by macrophages in the bloodstream. Increased levels of trimesic acid linker in urine after 24 h, i.e., two times higher than after empty NPs injection, were observed. All this indicates that the biodistribution/elimination data of the empty nanoMOFs cannot be extrapolated to drug-loaded nanoMOFs. In fact, Adv. Mater. 2018, 1707365 Reproduced with permission. [52] Copyright 2013, Royal Society of Chemistry. Figure 6. Biodistribution of iron nanoMOFs according to the iron concentration. Reproduced with permission. [9] Copyright 2012, American Chemical Society.

depending on the nature of the encapsulated molecule, the biological interaction of the drug delivery system might change. Early biodistribution of MIL-100(Fe) NPs in rats has also been investigated. Interestingly, a reversible accumulation of NPs in the lungs after i.v. administration was observed, which was mainly attributed to a pH-triggered decrease of NP colloidal stability in the blood upon i.v. administration. This unique property of MIL-100(Fe) NPs hold promise for specific lung targeting. Thus, MIL-100(Fe) NPs loaded with the anticancer drug gemcitabine monophosphate were found to increase the lung concentration of the drug, leading to the efficient treatment of an experimental model of lung metastasis. [125] In other in vivo experiments, biodistribution of coated nano-MOFs was assessed. For example, bioMOF prepared from the therapeutic agent zoledronate and calcium (CaZol) was modified by incorporating the targeting molecule folate and chains of PEG (Fol-PEG-CaZol). [68] Folate is a molecule with a high affinity for the folic acid receptor, highly expressed in many human cancers. Folate-NP conjugates tightly bind the folate receptor and trigger cellular uptake by endocytosis. [126] The biodistribution after single i.v. tail injection was evaluated both in healthy mice and in a xenograft subcutaneous tumor model. Unfortunately, in the study with healthy mice only one group of animals was treated by surface-modified nanoMOFs and therefore authors could not compare the biodistribution results with uncoated nanoMOFs, making impossible to conclude about the benefits of PEGylation in this model, in which 60% and 20% of the administered dose were found in liver and kidney, respectively. In the experiment using the xenograft tumor animal model, three groups were included and were treated with PBS (control group), Ca-Zol, and Fol-PEG-CaZol. Authors observed important differences in the tumor uptake: 82% of the administrated Fol-PEG-CaZol nanoMOFs were accumulated in the tumor compared to only 52% after treatment with the nontargeted nanoMOFs. This was attributed to the active targeting mediated by the folate molecules at the surface of the nanoMOFs. Once again, a fair conclusion about the role of PEGylation versus folate decoration is difficult to draw since a control group treated with PEG-CaZol, without folate molecules, was not included in this study. Zhang et al. also incorporated PEG and folate to coat the zinc imidazolate ZIF-8 NPs to develop an antitumoral formulation. [127] Two different drugs were encapsulated, the p-glycoprotein inhibitor verapamil and the antitumoral Dox. After i.v. administration in a xenograft subcutaneous model of melanome, the best tumor inhibition was observed when ZIF-8 nanoMOFs encapsulating both drugs were PEG coated and folate decorated. The authors assumed that the superiority of the treatment was due to the enhanced permeability and retention effect (EPR) and folate-mediated active targeting.

A very few studies dealing with nanoMOFs' distribution and therapeutic effect after oral administration have been performed to date. Lucena et al. reported the anti-inflammatory effect of orally administered 5-FU encapsulated in Cu-BTC, [96] using a peritonitis mice model induced by carrageenan. Authors concluded that the cytotoxicity against tumor cells observed in vitro could be related to the route of leukocyte activation or suppression of the inflammatory process. Future experiments should include a more adequate animal model to evaluate the mechanisms behind 5-FU-loaded Cu-BTC antitumor activity. Another example was proposed by Hartlieb et al. [128] where ibuprofen was cocrystallized with γ-cyclodextrins (γ-CD), the carboxylic group of the drug being coordinated to alkali metal cations such as K+. The main target of this porous framework, built up from (γ-CD)6 cubes, was to reduce the time required for the maximum uptake of ibuprofen (C max ) and to increase the halflife of the drug within the body. The benefit in terms of drug bioavailability after oral administration was demonstrated, using a mice model, with a 100% longer half-life in blood samples.

Recent Advances in Theranostics

Many of the recent advances in the field of MOFs for drug delivery have already been discussed in the precedent sections. In the past few years, another target in the field was to develop more complex systems to integrate therapeutics and diagnostics in a unique theranostic tool. [45,66,[129][130][131] The potential of iron carboxylates as theranostic agents was first proposed by some of us in 2010, when MRI measurements were performed on Wistar rats after i.v. injection of a suspension of MIL-88A NPs. [24] In this report, it was shown that the iron-based core was responsible for favorable relaxivities and imaging properties.

The surface of the iron carboxylate MIL-100(Fe) NPs was recently decorated by Sene et al. with γ-Fe 2 O 3 -cit ultrasmall NPs of superparamagnetic of iron oxide (USPIO) leading to novel nanoobjects (MIL/USPIO-cit). The system demonstrated good stability in aqueous solution and physiological media. Moreover, the presence of maghemite NPs conferred to the nanoMOFs a very high r 2 relaxivity, comparable to the best commercial available systems and these imaging properties were further confirmed in vivo by T 2 -weighted MRI. This potential theranostic tool improved imaging contrast properties of the nanoMOFs in vivo while conserving the high drug loading/release capacities. [130] Another theranostic approach based on ZIF-8 NPs was recently reported by Yang et al. in a sophisticated system to combine MRI, multidrug chemotherapy, and photothermal synergistic therapy (Figure 7). [129] Here the nanoMOFs were used as a shell within the "sandwich" nanocomposite made from a core of CoFe 2 O 4 mesoporous NPs, a polydopamine layer and a shell of ZIF-8. The core acted as an MRI probe, as a photothermal agent, and as a loading platform for Dox. The polydopamine layer prevented from the leakage of Dox while the nanoMOFs' shell allowed encapsulation of the hydrophobic anticancer drug camptothecin and as the switch for the pH and NIR stimulation-responsive release of the two drugs. After in vivo i.v. administration in a xenograft tumor model in mice, an efficient photothermal integration was observed together with high drug concentration at the tumor site by quick release of encapsulated drugs, negligible toxicity, and a synergic antitumor effect of the hybrid nanocomposites. [129] A UiO-type Zr carboxylate MOF has also been proposed for photodynamic therapy by Zhang et al. [45] In this study, CT images demonstrated the accumulation of the nanoMOF into the tumor, after intravenous administration in rats bearing an hepatoma. One can nevertheless raise concerns about this study as part of a lack of control experiments in Adv. Mater. 2018, 1707365 healthy rats. Indeed, as previously mentioned, nanoparticles tend to spontaneously accumulate into the liver healthy macrophages due to the opsonization process. It is therefore difficult to conclude if the tumor accumulation resulted from a specific targeting or corresponded only to unspecific capture by the Kupffer cells of the liver, common to most nanoparticles.

Conclusion and Perspectives

MOFs' materials have emerged initially due to their promises for numerous industrial applications such as gas separation/storage or catalysis, as a consequence of their tunable porosity and large chemical and structural diversity. In parallel, the field of biomaterials has been under continuous development with numerous natural and synthetic materials being considered for controlled drug delivery and theranostic applications. The temporal convergence of MOF discovery and biomaterials science expansion explains probably why these new hybrid materials, prepared by chemists far from the clinical environment, were soon identified as potential candidates in the search of new and more safe drug nanocarriers. [138,24,139] Since then, in the last decade, studies dealing with applications of MOFs in the pharmaceutical sciences have grown steadily (see Figure 8). Initial reports in this domain have focused on fundamentals that govern the drug loading and release of a large array of therapeutic and theranostic systems, while later reports incorporated basic in vitro characterization including first toxicity studies. Noteworthily, during the past 3-4 years, an increasing number of studies have highlighted the first in vivo benefits of nanoMOFs to treat different diseases in animal models. One can thus expect in the upcoming years exciting new developments concerning the preclinical in vivo evaluation of MOFs. [59,107,125,127,[129][130][131]133,128,140] A key condition of success will nevertheless require furthering strengthening the exchanges between chemists, pharmacists, and clinicians, for the creation of multidisciplinary projects dealing with the translation of nanoMOFs into more realistic biomedical solutions. When analyzing the evolution of the field in the last 5 years (Figure 8), one realizes that now is time to move from "simple" collaborations and teaming to integrate additional data, methodologies, perspectives, and concepts from the various disciplines implied, which would lead to important advances in fundamental understanding and to solve real biomedical problems. Individual researchers also need to acquire a deeper understanding in these disciplines and be fluent in their languages and methodologies. Numerous university programs already offer multidisciplinary training in biomaterials science, especially for postgraduate students and through Ph.D. programs. Some of these programs have already started to deliver researchers with the knowledge and skills to work at the interfaces of biomaterials disciplines that will hopefully make a precious contribution for the progression of MOFs' bioapplications.

The remaining challenges are, however, still numerous if one targets the use of nanoMOFs as third-generation drug delivery systems as summarized by Yun et al. [141] According to Park's group classification, third-generation formulations are modulated delivery systems that must be able to cross both physicochemical and biological barriers (Table 3). [141] A particular effort will be required to get a rational design approach covering all the development stages, from the choice of the nanoMOF composition, the development of safe synthesis, and the surface modification and formulation conditions, the best route of administration and dose, bearing in mind the current treatment for a specific disease [129] Copyright 2017, American Chemical Society. models; ii) stop assuming effects, such as EPR or active targeting without performing the experiments that scientifically demonstrate these phenomena, by including the adequate control groups or using labeling strategies that ensure quantitative analysis of drug biodistribution; and, for instance, iii) avoid the use of model drugs or fluorescent probes that do not possess the same physicochemical properties as those of the real drug. The risk also exists to limit the in vivo studies to the reproduction of those experiments already performed with other nanocarriers (i.e., liposomes, lipid/polymer NPs or polymer micelles), without showing any ground-breaking advantages for nanoMOFs. Moreover, until now, cancer treatment is almost the unique application of nanoMOFs when several other important biomedical applications may be considered, such as improved oral bioavailability for traditional drugs (see the work by Hartlieb et al. [128] ), oral controlled release, nonviral gene delivery, vaccine adjuvant delivery systems, or delivery across the blood-brain barrier, among others.

Even if the number of works rapidly increases every year (Figure 8), MOFs as delivery systems are still at their infancy. Indeed, they have appeared more than 30 years after the first drug delivery systems, such as liposomes in the 1960s, [144] polymeric particles in the 1970s, [145] and polymeric micelles in the 1980s. [146] Thus, results with MOFs have not reached yet the maturity of other nanovectors in terms of pharmaceutical development, and it is difficult to make a rigorous comparison. However, MOFs present two important advantages that could help to overcome limitations to cross biological barriers, such as the gastrointestinal or the blood-brain barrier, in contrast to other materials used in drug delivery. First, the great flexibility of MOFs' chemistry offers unprecedented possibilities to easily modify the surface of the NPs to modulate biological interactions. Second, MOFs' drug loading is in many cases far superior to most of the other studied drug delivery systems, meaning that even with a limited passage, the pharmacological effects would be significantly improved; one shall also consider that despite biodegradation, the kinetics of release of the drug can be controlled through a careful tuning of the host-guest interactions. Finally, when developing new MOF-based delivery systems, one shall also pay attention to the ability to produce nanoMOFs under conditions compatible with the pharmaceutical industry from their scale-up to their integration into adapted pharmaceutical formulations.

In conclusion, MOFs are promising materials for biomedical applications but before to be able to reach the clinical arena, MOF scientists should take advantage of the knowledge acquired during decades by others and go further in solving the relevant limitations of other already developed nanomedicines.

Figure 1 .

 1 Figure 1. The stealth effect of functionalized nanoparticles. (A) Uncoated NPs are covered easily with different proteins from the blood and captured by macrophages; (B) Dense coating of NPs shield NPs against protein adsorption and thus NPs are not recognized by macrophages 6,10 .

Figure 3 .

 3 Figure 3. Schematic illustration of influence of PEG density on serum adsorption and macrophage uptake. NPs without coating have tendency to adsorb all types of proteins and are easily recognized by macrophages. NPs with low PEG density (0.32-0.64 PEG/nm 2) have less adsorption of the proteins, but shielding effect is not enough, thus NPs are still recognized by macrophages. NPs with both higher densities (above 0.64 PEG/nm 2 ) adsorbed much less proteins, resulting in significantly reduction of macrophage uptake.18

  nanoparticles (PCSN) have size approximately 270 nm and surface charge of -5.3 mV. Protein adsorption from fetal bovine serum (FBS) was then compared for PCSN, PEG-MSN and GSH-MSN, respectively. In the case of PCSN, a significant reduction of protein adsorption from the serum was observed in comparison with other particles. Adsorbed proteins on PCSN were mostly of low molecular weight (20-60 kDa). Immune response as well complement and coagulation proteins were also significantly reduced on PCSN, thus paving the way for a significant reduction of macrophage recognition. Accordingly, the macrophage uptake for PCSNs and PEG-MSNs, in murine macrophage-like cell line, RAW246.7 and pre-incubated in 55% FBS for 1 h, showed almost no internalization of PCSN, while PEG-MSN were still captured by macrophages. Authors concluded that the shielding ability towards adsorption of proteins from serum was more than satisfactory. Furthermore, the targeting ability of PCSN was confirmed in HER2-receptor-overexpressing cancer cell line SK-BR3 and HER2-receptornegative cell line HEK293T. In addition, in vivo targeting and anticancer efficacy were further confirmed on nude mice bearing SK-BR3 cell xenografts. Moreover, accumulation of PCSN in RES organs was significantly reduced compared to PEG-MSN and was not affected by the comparably larger size of NPs. Finally, Camptothecin-loaded PCSN (PCSN(CPT)) had higher therapeutic efficacy compared to PEG-MSN(CPT).51

Figure 4 .

 4 Figure 4. Protein corona shield nanoparticles (PCSN). PCSN effectively prevent the adsorption of blood protein, avoiding recognition by macrophages and thus have better targeting abilities.51

Figure 5 .

 5 Figure 5. (a) Schematic representation of MIL-100(Fe) structure. Reaction of Fe(III) trimers and trimesic acid leads to the formation of hybrid supertetrahedra, which further assemble into the MIL-100(Fe) zeotype architecture, with consists of larger mesoporous cages (in red) accessible through hexagonal windows, and smaller mesoporous cages (in green), accessible through pentagonal windows. (b) Schematic representations of several MOFs topical structures of interest in biomedicine : ZIF-8, UIO-66, MIL-53, MIL-88A, CD-MOF-1 and MIL-101. Metal Polyhedra, carbon atoms are in color or black, respectively.

Figure 6 .

 6 Figure 6. (a) General scheme of a microfluidic oil-segmented droplet confined MOF synthesis.123

Figure 8 .

 8 Figure 8. Schematic illustration of the coating with polymers: (I) PEG and (II) Stp10-C.134

Fe 3 O 4 @

 34 Bio-MOF-FC NCs could produce a negative (darkening) contrast. In vivo T 2 -weignted MR images were performed in tumor-bearing BALB/c mice after i.v. injections. If for the Fe 3 O 4 @Bio-MOF-FC animal group, an enhancement of the negative contrast in tumor site was observed, authors did not include any comparison with a standard contrast, as well as no control group with injection of non-coated particles. Moreover, no further study was performed to demonstrate in vivo the benefits of this anticancer strategy.165

Figure 12 .

 12 Figure 12. Schematic illustration for synthesis of 5-FU loaded Fe 3 O 4 @Bio-MOF-FC NCs. 165

Figure 9 .

 9 Figure 9. Schematic illustration of the cell uptake of exosome coated nanoMOFs and release mechanism of the cargo. 157

Figure 10 .

 10 Figure 10. (a) Schematic illustration of the Fe-soc-MOF@PPy synthesis; (b) Their use for PPT application. 166

119

 119

Figure 11 .

 11 Figure 11. MIL-100(Fe) NPs reversible aggregation in physiological media. (A and B) Colloidal stability ((A)size and (B) surface charge) of MIL-100(Fe) NPs in different pH regions, measured by DLS. (C) Evolution of particles aggregation over time in rat serum; represented by DLS.119

  mechanism of degradation was attributed to the progressive replacement of trimesate ligands from the MOF structure by phosphate ions from PBS. Moreover, authors demonstrated that drug loading (Gem MP) or surface modification (CD phosphate) did not affect the crystallinity of the MOF. This interesting work helps to better understand the process behind the degradation of MOFs in physiological media, thus provided valuable insights over the fate of the nanoMOFs in vivo. However, in vivo conditions are much more complicated, including protein adsorption on the outer surface of MOFs, that are likely to alter the degradation mechanism. 171 Another thesis (Ioanna Christodoulou) carried out at ISMO (Paris Saclay) under the supervision of Dr R. Gref in collaboration with Dr C. Serre (IMAP) is in progress to shed more light on the degradation mechanisms of nanoMOFs for biomedical applications.

Figure 12 .

 12 Figure 12. Morphology of MIL-100(Fe) nano-and microsize of MOFs and size changes after incubation in water or PBS. (A) MOFs before and (B) after degradation caused by incubation of MOFs with PBS for 6 h; measured by TEM. (C) Size dependence of incubation of NPs in water (orange) or in PBS (blue); measured by DLS in duration of 2 days. (D) MicroMOFs before and (E) after degradation in PBS, observed by Raman-microscopy. (F) Size distribution of microMOFs before (green) and after (red) degradation. 171

Folate-

  NPs conjugates tightly bind the folate receptor and trigger cellular uptake by endocytosis204 . The biodistribution after single i.v. tail injection was evaluated both in healthy mice and in a xenograft subcutaneous tumor model. Unfortunately, in the study with healthy mice only one group of animals was treated by surface modified nanoMOFs and therefore authors could not compare the biodistribution results with uncoated nanoMOFs, making impossible to conclude about the benefits of PEGylation in this model, in which 60 % and 20 % of the administered dose was found in liver and kidney respectively. In the experiment using the xenograft tumor animal model 3 groups were included and were treated with PBS (control group), Ca-Zol, and Fol-PEG-CaZol. Authors observed important differences in the tumor uptake: 82% of the administrated Fol-PEG-CaZol nanoMOFs were accumulated in the tumor compared to only 52% after treatment with the non-targeted nanoMOFs. This was attributed to the active targeting mediated by the folate molecules at the surface of the nanoMOFs. Once again, a fair conclusion about the role of PEGylation versus folate decoration is difficult to draw since a control group with treated with PEG-CaZol, without folate molecules, was not included in this study. Zhang et al. also incorporated PEG and folate to coat the zinc imidazolate ZIF-8 NPs to develop an antitumoral formulation205 . Two different drugs were encapsulated, the p-glycoprotein inhibitor verapamil and the antitumoral Dox. After i.v. administration in a xenograft subcutaneous model of melanome, the best tumor inhibition was observed when ZIF-8 nanoMOFs encapsulating both drugs were PEG coated and folate decorated. The authors assumed that the superiority of the treatment was due to the enhanced permeability and retention effect (EPR) and folate mediated active targeting.Very few studies dealing with nanoMOFs distribution and therapeutic effect after oral administration have been performed to date. Lucena et al. reported the anti-inflammatory effect of orally administered 5-FU encapsulated in Cu-BTC 206 , using a peritonitis mice model induced by carrageenan. Authors concluded that the cytotoxicity against tumor cells observed in vitro could be related to the route of leukocyte activation or suppression of the inflammatory process.

  photothermal integration was observed together with high drug concentration at the tumor site by quick release of encapsulated drugs, negligible toxicity and a synergic antitumor effect of the

Figure 16 .

 16 Figure 16. (A) Synthesis of Co/DPZ/C Nanocarrier and (B) theranostic strategy for MR imaging-guided multi-drug chemotherapy and photothermal synergistic therapy. 208

Figure 17 .

 17 Figure 17. (A) Synthesis of MD@Lip and (B) Apoptosis caused by nanoparticles. 167

Figure 1 .

 1 Figure 1. Schematic representation of nanoMOFs-based particles behavior after intravenous administration.A colloidal suspension of nanoMOFs is stable in pH below 6.5 (except at pH close to 4). Once NPs are administered in the blood (pH 7.4), NPs start to aggregate, resulting in early lung filtration. Within hours, aggregates of nanoMOFs start to disassemble and progressive release of encapsulated drug is triggered (represented as white spots). Lung inflammation induced by the presence of the aggregates is rapidly reversed after the disassembling process takes place.1

Figure 2 .

 2 Figure 2. Schematic illustration of treatment. (B) At day 0, mice were intravenously injected with LLC-1 cancer cells, at day 15, 40 animals were randomly assigned to the 4 treatment groups: MIL-100(Fe)-GMP NPs (15 mg/Kg GMP; 50 mg/ Kg MIL-100(Fe)), GMP (15 mg/Kg GMP), MIL-100(Fe) NPs (50 mg/Kg MIL-100(Fe)), and the sham group (glucose 10 %). At day 21, 5 animals per group were sacrificed to evaluate metastasis development. At day 22, animals received a second treatment dose and were sacrificed at day 29 to evaluate treatment efficacy. (C) Only MIL-100(Fe)-GMP caused a reduction of the total number of lung metastasis (N metastasis) at day 29 compared to GMP, MIL-100(Fe), and the sham group. (E) Lung metastasis in the different treatment groups appear as white spots. 1

Figure 3 .C

 3 Figure 3. Schematic representation of the coating reactions: (A) grafting process of amino-PEG on the surface of MIL-100(Fe) 5 , (B) Graftfast reaction of acryl-PEG on the surface of MIL-100(Fe) 6 , (C) CD-Ph and CD-Ph-PEG coated on the surface of nanoMOFs.

  β-cyclodextrin polyphosphate and β-cyclodextrin polyphosphate with adamantine-PEG, were provided by Prof. Ruxandra Gref and Dr. Xue Li from ISMO at Paris-Saclay University. TEM and SEM pictures were collected with the assistance of Dr. Farid Nouar and Dr. Mathilde Lepoitevin from our group. All preparations of NPs for in vivo study were handled inside the radioactive room within the Faculty of Pharmacy of Paris-Saclay University at Châtenay-Malabry. In vivo experiments were performed also in the radioactive room, located in the animal facility at the Faculty of Pharmacy. To be able to handle animal experiments, I obtained the certificate entitled "Concepteur de Projects modèle rongeurs" (CNRS course, 14-22/11/2016). Therefore, I performed all in vivo experiments and radioactive analysis by myself. Prior to any experiments, Dr. Catherine Cailleau, responsible for the animal experiments in Institut Galien, kindly trained me for some specific procedures with mice, including blood extraction from a facial vein and by cardiac puncture or intraperitoneal injections. Dr. Ghozlene Mekhloufi and Assia Hessani are also responsible for radioactive rooms and the safety of the users of the rooms. Prior to any experiments, they provided me with adequate training on how to handle radioactive elements and specific behavior in the rooms.

  g. Erlenmeyer flask), trimesic acid (0.25 g, 1.19 mmol) was mixed with iron (III) nitrate nonahydrate (0.72 g, 1.78 mmol) in distilled water (90 ml) under magnetic stirring (600 rpm) for 48 h. A change of color was observed during the course of the reaction from yellow to orange (Figure4). NPs were then centrifuged in 50 ml tubes (14500 rpm, 10 min) and washed twice with distilled water, and twice with ethanol after re-dispersion (sonication bath and sonication tip, 10%, 1 min) and then followed by centrifugation (14500 rpm, 10 min). Part of NPs was dried for further characterization purposes. Note that for other experiments NPs were systematically used wet to prevent from aggregation. Before any further surface modification, MIL-100(Fe) were washed only once in distilled water.

Figure 4 .

 4 Figure 4. The reaction of MIL-100(Fe) NPs. The left picture shows a clear solution, after mixing all reagents. The right picture presents the same solution after mixing for 48 h. A solution became cloudy, which indicates the formation of MIL-100(Fe) NPs.

Figure 5 .

 5 Figure 5. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) collected at the range 2° < 2θ < 30°

Figure 6 .

 6 Figure 6. FTIR of MIL-100(Fe) collected at the range of 4000-500 cm -1 .

Figure 7 .

 7 Figure 7. N 2 sorption isotherm (77K, P 0 =1 bar) of MIL-100(Fe), with BET surface area 1570 m 2 /g. Activation condition: 150°C under vacuum for 6 h.

Figure 8 .

 8 Figure 8. TGA of MIL-100(Fe), measured at oxygen flow (20 ml/min) and heating speed 3°C/min at temperature range 25-600°C.

Figure 9 .

 9 Figure 9. TEM images of MIL-100(Fe) NPs before and after surface modification. (A) bare MIL-100(Fe) with a size of 38±8 nm. (B) MIL-100(Fe)@polyethylene glycol Graftfast with a size of 44±5 nm. (C) MIL-100(Fe)@polyethylene glycol NHS with a size of 45±7 nm. (D) MIL-100(Fe)@β-cyclodextrin-phosphate with a size of 52±7 nm. (E) MIL-100(Fe)@β-cyclodextrinphosphate-polyethylene glycol with a size of 46±6 nm.

Figure 10 .MIL- 100 (Figure 11 .

 1010011 Figure10. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) NPs: (red) MIL-100(Fe), (blue) MIL-100(Fe)@PEG Graftfast, (yellow) MIL-100(Fe)@PEG NHS, (green) MIL-100(Fe)@βCD-Ph, (orange) MIL-100(Fe)@βCD-Ph-PEG.

7

 7

4 .

 4 Evaluation of size dependence in various media. For a better understanding of the MIL-100(Fe) NPs behavior in various media (milliQ water, PBS and rat serum), different colloidal stability tests were performed. Both particle size distribution (% Number) and ζ-potential for uncoated and coated MIL-100(Fe) NPs were Name Quantity of the coating (wt %) MIL-100@PEG Graftfast ~20 MIL-100@PEG NHS ~14 MIL-100@CD-Ph ~10 MIL-100@CD-Ph-PEG ~26

  7.4), MIL-100(Fe) are highly aggregated (size >1000 nm), while ζ-potential slightly increases (-30±2 mV). Compared to uncoated MIL-100(Fe), at pH 5.4 NPs coated with PEG Graftfast, PEG NHS, β-CD-Ph and β-CD-Ph-PEG have a size of 77±5, 87±4, 96±1 and 104±1 nm, respectively and ζ-potential -45±1, -22±1, -40±2 and -32±2 mV, respectively. Compared to uncoated MIL-100(Fe), only NPs coated with PEG NHS possess a shifted point of zero charge, which occurs at pH 4.5 (size of 1500±130 nm and surface charge of -3±2 mV), as expected due to the reaction between the -COOH groups and the NHS molecules. These coated NPs, however, do not show any aggregation at pH 7.4. All these coated NPs look at first sight efficient to prevent from aggregation at physiological blood pH (ca. 7.4) with only a minimal aggregation. At this pH, NPs coated with PEG Graftfast, PEG NHS, β-CD-Ph and β-CD-Ph-PEG have a particle size of 198±8, 75±4, 60±2 and 106±3 nm, respectively and ζ-potential -42±1, -30±2, -31±3 and -35±3 mV, respectively. The most stable coating is finally the β-CD-Ph-PEG coating. These NPs kept the same size throughout all the investigated pH range (~105 nm), while a ζ-potential remains strongly negative.

Figure 12 .Figure 13 )

 1213 Figure 12. Hydrodynamic diameter of MIL-100(Fe) NPs uncoated and coated with PEG Graftfast, PEG NHS, β-CD-Ph and β-CD-Ph-PEG, at different pH conditions. Values show the average hydrodynamic diameter. In case of two particles population, the first peak is represented in a bar, while the second peak is written.

Figure 13 .

 13 Figure 13. Hydrodynamic diameter dependence of MIL-100(Fe) uncoated or coated in PBS (c= 1 mg/ml).

  First, a fast increase of pH, changing from acidic solution (3.6-4.4) to pH 7.4 (blood pH) which usually leads to a high aggregation of NPs. Secondly, there is albumin in serum that triggers the formation of a protein corona around these exogenous NPs.1,4 Note that all values of ζ-potential of various NPs in serum were the same (-7±2 mV), which is probably related to the serum itself.Initial surface charge of NPs did not have an effect on the final ζ-potential of NPs in serum.

Figure 14 .

 14 Figure 14. Comparison of hydrodynamic diameter (nm) of NPs uncoated and coated with PEG Graftfast, PEG NHS, CD and CD-PEG in water (1 mg/ml, blue) and in solution mimicking in vivo conditions (0.2 ml NPs in glucose, 5 mg/ml mixed with 1.5 ml filtrated rat serum, orange).

  Figure16. Lungs NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injection of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue)3 H-Gem MP free, (red)14 C-MIL-100(Fe)@ 3 H-Gem MP, (green)14 CMIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet)14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top figure represents 14 C-BTC accumulations, which indicates the accumulation of NPs (ligand in NPs and detached ligand). The bottom figure represents 3 H-Gem MP accumulations, which indicated an accumulation of the drug (Gemcitabine monophosphate encapsulated or released).

121 liver

 121 was not reduced upon coating. Indeed, liver concentration was 19.4±2.5, 25.5±5.4, 25.9±5.1

  Figure19. Kidneys and heart NPs accumulation profile at 0.5, 1, 6 and 24 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue)3 H-Gem MP free, (red)14 C-MIL-100(Fe)@ 3 H-Gem MP, (green)14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet)14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top left figure represents14 C-BTC accumulation in the kidneys, coming from NPs (ligand in NPs and detached ligand), while the right top figure represents 3 H-Gem MP accumulation in the kidneys, coming from the drug (Gemcitabine monophosphate encapsulated or released). The bottom left figure represents14 C-BTC accumulation in the heart, while the right bottom figure represents 3 H-Gem MP accumulations in the heart.

  Figure20. Fat and muscles NPs accumulation profile at 6 h after intravenous injections of 0.2 ml of NPs formulations (5 mg/ml, ~50 mg/kg: (dark blue)3 H-Gem MP free, (red)14 C-MIL-100(Fe)@ 3 H-Gem MP, (green)14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast, (violet)14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS, (light blue)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph, (orange)14 C-MIL-100(Fe)@ 3 H-Gem MP@CD-Ph-PEG). The top left figure represents14 C-BTC accumulation in fat, coming from NPs (ligand in NPs and detached ligand). The top right figure represents 3 H-Gem MP accumulation in fat, coming from the drug (Gemcitabine monophosphate encapsulated or released). The bottom left figure represents 14 C-BTC accumulation in muscles, and bottom right figure represents 3 H-Gem MP accumulations in muscles.

Figure 22 .

 22 Figure 22. Diagrams representing the percentage distribution of NPs or drug in different organs, blood, urine, and feces at 24 h after intravenous injection (0.2 ml, 5 mg/ml per mouse) in mice. Results were calculated from the total injected dose. The top left figure represents 14 C-BTC accumulations of NP-Gem MP, while the top right figure represents 3 H-Gem MP accumulations of NP-Gem MP. The bottom left figure represents 14 C-BTC accumulations of NP-Gem MP-CD-PEG, while the bottom right figure represents 3 H-Gem MP accumulations of NP-Gem MP-CD-PEG.

30: 70 ,

 70 50:50, 70:30 and 90:10 were tried, all other parameters from the original synthesis being kept unchanged.

Figure 23 .

 23 Figure23. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) NPs obtained following different synthesis routes: (black) room temperature synthesis (c1, patent concentration), (red) microwave synthesis, 60°C, 1 h reaction, patent concentration, (bright blue) microwave synthesis, 60°C, 1 h reaction, patent concentration/2, (pink) microwave synthesis, 60°C, 1 h reaction, patent concentration*2, (green) room temperature synthesis plus 0.5 mol NaOH/ligand, (dark blue) room temperature synthesis plus acetic acid/Fe=2.5.

Figure 24 .

 24 Figure 24. SEM or TEM images of MIL-100(Fe) obtained by different synthesis routes. (A) RT synthesis+ 0.5 mol NaOH/linker, 34±5 nm, (B) MW c1, 1 h, 34±6 nm, (C) MW c3, 1 h, 22±2 nm, images not very clear, (D) MW c2, 1 h, 25±3 nm.

  However, if one aims at using these nanoMOFs to treat other diseases than lung cancer, it is mandatory to change such biodistribution. The biggest limitation of these MOF nanoparticles is their poor colloidal stability, particularly in serum conditions. Therefore, we utilized a recently developed new type of smaller MIL-100(Fe) NPs, made at room temperature by our team, and further surface modified these NPs with various coatings following known routes, based on non covalent or covalent interactions: (a) PEG Graftfast, (b) PEG NHS, (c) β-CD-Ph and (d) β-CD-Ph-PEG. Their physico-chemical characterization was performed (IR, PXRD, TGA, N 2 sorption, DLS (size and ζ-potential), SEM and/or TEM) and their colloidal stability assessed in different media: milliQ water, PBS and rat serum mimicking in vivo conditions. If these nanoparticles are

Figure 1 .

 1 Figure 1. Methotrexate (MTX) molecular structure.

Figure 2 .

 2 Figure 2. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) before MTX encapsulation (red) and MIL-100(Fe) after MTX encapsulation (HCl 10 mmol/l) in weight ratio of MIL-100(Fe)/MTX 1:1 (blue) and 1:3 (yellow).

Figure 3 .

 3 Figure 3. FTIR spectra of MIL-100(Fe) (black), MTX@MIL-100(Fe) encapsulated (in HCl 10 mmol/l) at MIL-100(Fe)/MTX 1:1 (red), and 1:3 (blue) weight ratio.

1 Figure 4 .

 14 Figure 4. TGAs of MIL-100(Fe) before MTX encapsulation (red) and after MTX encapsulation in HCl 10 mmol/l, MIL-100(Fe)/MTX 1:1 (yellow) and 1:3 (blue) weight ratio; measured at oxygen flow (20 ml/min) and heating speed 3°C/min at temperature range 25-600°C.

Figure 6 .

 6 Figure 6. The release profile of MTX in PBS (37°C) from MTX@MIL-100(Fe), calculated from UV-Vis (PBS, 37°C).

Figure 7 .

 7 Figure 7. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) before MTX encapsulation (red), MIL-100(Fe) after MTX encapsulation in HCl 10 mmol/l in weight ration of nanoMOFs to MTX 1:3 (blue), and MIL-100(Fe) after 24 h of incubation in PBS (yellow).

  Steunou (Université de Versailles St Quentin). The first synthesis and characterization of these MIL-100(Fe)@γFe 2 O 3 composites was done by Dr. Sene, a former post-doctoral fellow. My role did consist mainly in the synthesis and characterization of MIL-100(Fe) NPs and the reproduction of the MIL-100(Fe)@γFe 2 O 3 NPs following previously established protocols. I also carried out the MTX encapsulation and release study. Some of the characterizations of MIL-MIL-100(Fe)@γFe 2 O 3 NPs were done by Dr. Sene (e.g., TEM images). This study is now being pursued by a new PhD candidate, Heng Zhao.

Figure 8 .

 8 Figure 8. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) NPs, MIL-100(Fe)@γFe 2 O 3 -10wt%, and MIL-100(Fe)@γFe 2 O 3 -20wt% (intensities are in arbitrary units).

Figure 9 .

 9 Figure 9. TGA of MIL-100(Fe) NPs (black), MIL-100(Fe)@γFe 2 O 3 -5wt% (red), and MIL-MIL-100(Fe)@γFe 2 O 3 -10wt% (blue); measured at oxygen flow (20 ml/min) and heating speed 3 C/min at temperature range 25-600°C.

2 O 3 Figure 10 .Figure 11 .

 231011 Figure 10. TEM images of nanoparticles: (A) MIL-100(Fe)@γFe 2 O 3 -5wt% with particle size of 72±10 nm and (B) MIL-100(Fe)@γFe 2 O 3 -10wt% with particle size of 67±12 nm. Smaller black particles represent iron oxide nanoparticles inside MIL-100(Fe).

Figure 12 .

 12 Figure12. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) NPs, MIL-100(Fe)@γFe 2 O 3 -10wt%, and MTX-30 wt%@MIL-100(Fe)@γFe 2 O 3 -10wt% (intensities are in arbitrary units).

  and 93 h) of agitation at 37°C. Calibration curves of MTX in HCl and PBS are presented in Annex 3 (A2-4 and A2-5).

Figure 15 .

 15 Figure 15. Schematic representation of the life cycle of the C. elegansa.16

Figure 16 .

 16 Figure 16. Scale-up synthesis of MIL-100(Fe) in a 30 liter reactor. Product mixed, first minutes of reaction (left picture). A solution containing the formed nanoparticles (7-days reaction) (right picture).

Figure 17 .

 17 Figure 17. TEM image of MIL-100(Fe) obtained from scale-up RT synthesis (after gold metallization). NPs have size of 32±7 nm.

Figure 18 .

 18 Figure 18. PXRD patterns (λ Cu ≈1.5406 Å) of MIL-100(Fe) NPs, MTX@MIL-100(Fe) encapsulated in HCl 10 mmol/l and milliQ water.

Figure 20 .

 20 Figure 20. TGA of MIL-100(Fe) before MTX encapsulation (red) and after MTX encapsulation encapsulation in HCl 10 mmol/l in MIL-100(Fe)/MTX 1:3 weight ratio (blue) and in milliQ water (yellow); measured at oxygen flow (20 ml/min) and heating speed 3 C/min at temperature range 25-600° C.

Figure 21 .

 21 Figure 21. The release profile of MTX from MIL-100(Fe) (MTX encapsulated in aqueous conditions in 1:3 weight ratio nanoMOFs to MTX) in 40 ml milliQ water at RT.

First, the toxicity

  test of MIL-100(Fe) nanoparticles per se was performed on C. elegans. To determine if MIL-100 affects the general growth and development of C. elegans, MIL-100(Fe) (1 mg) was mixed with a food source of C. elegans, E. coli OP50 (1 mg in 50 , in a 1:1 ratio, and seeded on NGM (5 ml) to feed the worms (Figure 22). Synchronized L1 stage worms were transferred to either E. coli OP50 or MIL-100-OP50 and left it for two days at 25°C. No significant difference in terms of worm's growth and development after 48 hours was observed between OP50 and MIL-100-OP50 (worms grow to the same stage at the same time), indicating that MIL-100 is not toxic to the worms.

Figure 22 .

 22 Figure 22. Overall growth and development of C. elegans (C. elegans -age: week 17) in the presence of MIL-100(Fe) NPs. The left picture represents C. elegans in OP50, while the right picture represents worms in MIL-100-OP50.

Figure 23 .

 23 Figure 23. Reproduction of C. elegans in MIL-100(Fe). The number of laid eggs or hatched larvae per worm per hour was quantified for C. elegans is the presence of normal food source E. coli OP50 (OP50) alone or E. coli OP50 mixed with MIL-100(Fe) (MIL100-OP50).

Figure 24 .

 24 Figure 24. The top left picture represents C. elegans on OP50, the top right picture represents worms in 1 mg of MIL-100 and OP50. The bottom picture represents worms in 1 mg of MIL-100 with encapsulated MTX and OP50.

Figure 25 .

 25 Figure 25. Toxicity effect of MTX on reproduction of C. elegans. Number of laid eggs or hatched larvae per worm per hour, was quantified for C. elegans in the presence of: (a) OP50, (b) OP50 with MIL-100, (c) OP50 with supernatant (SUP) from MIL-100 (MIL-100 SUP), (d) OP50 with MTX@MIL-100, (e) OP50 with supernatant from MTX@MIL-100 encapsulation (MIL-100-MTX SUP).

  First, the synthesis of the newly developed (in our team) route to produce smaller nanoparticles of MIL-100(Fe) was considered (patent: Panchal, M.et al., "Low Temperature Process for the Synthesis of MOF Carboxylate Nanoparticles" European application N°17305119.4 (2017), International application N° PCT/EP2018/052115 (2018)) and further surface functionalised through different green strategies based on their previously reported excellent colloidal stability and satisfactory properties in vitro. Covalent or non covalent strategies were selected as follow: (a) polyethylene glycol though the radical polymerization "Graftfast", (b) covalent grafting of polyethylene glycol through the NHS route, (c) covalent grafting through the use of β-cyclodextrin-phosphates or (d) β-cyclodextrin-phosphate adamantane polyethylene glycol molecules. The colloidal stability of the coated nanoparticles was first evaluted in various media (milliQ water, PBS, and serum). However, if a good stability as well as satisfactory size of the NPs was observed in water and PBS, once the particles were dispersed in serum their size increased to reach the 150-200 nm range. This indicates that nanoparticles end to aggregate once in serum by almost a two-fold factor as compared with pure

(

  MTX), a small chemotherapeutic agent that induces cell apoptosis, that had never been loaded before into MIL-100(Fe) NPs. MTX is used to treat various diseases (e.g. psoriasis, rheumatiod arthritis), especially when patients have no adequate response to other types of therapy. To overcome the severe side effects of MTX, we loaded successfully MTX inside MIL-100(Fe) using "green" conditions leading to a very high encapsulation rate close to 50±5 wt%. The release profile of MTX from MIL-100(Fe) in PBS lead however to a fast drug release with most of the drug released within the first 7 hours of incubation. These results were first exploited for two different potential applications: (a) core-shell magnetic MIL-100(Fe)@γFe 2 O 3 and (b) toxicity tests with C. elegans.

Figure A1- 1 .

 1 Figure A1-1. Example of cylindrical restraint tube for intravenous injection of the mice by the tail vein (left picture). Lateral tail vein, accessible for injections (right picture). (Pictures source: https://theodora.com/rodent_laboratory/injections.html and http://www.jove.com/science-education/10198)

Annex 3 . 2 .

 32 TGA-the quantity of the coatingWe calculated the quantity of the coatings from the TGA results, based on the weight loss (between 150°C and the next weight loss), presented in the FigureA2-2.

Figure A2- 2 .Annex 3 . 3 .Figure A2- 3 .Figure A2- 4 .Figure A2- 5 .Figure A2- 6 .Figure A2- 7 .

 23334567 Figure A2-2. TGA of MIL-100(Fe) uncoated (red) and coated with PEG Graftfast (blue), PEG NHS (yellow), β-cyclodextrin-phosphate (green) and β-cyclodextrin-phosphate-PEG (orange), measured at oxygen flow (20 ml/min) and heating speed 3 C•min -1 at temperature range 25-600°C.

Figure A2- 8 .Figure A2- 10 .

 810 FigureA2-8. Comparison of release profile of MTX in PBS (37°C) from MTX@MIL-100(Fe) (red, yellow), and MTX@MIL-100(Fe)@γFe 2 O 3 -10wt% (blue, green), obtained by me (AM) or Heng Zhao (ZH), calculated from UV-Vis.

  Nevertheless, to date, Doxil represents the only FDA-approved PEGylated Teresa Simon-Yarza received her undergraduate degree in pharmacy and her master degree in pharmaceutical technology from the University of Navarra, where she completed in 2013 her Ph.D. in pharmacy with a thesis focus on biomaterials for heart repair. Then, she joined the Institut Lavoisier and the Insitut Galien (University Paris-Saclay) as a postdoc, to develop metalorganic frameworks for biomedical applications. Since 2016, she is a postdoctoral researcher in the laboratory INSERM U1148 (Paris) working on 3D hydrogels for tissue repair. With a multidisciplinary background, her main research interest focuses on the biomaterials development, from the synthesis to the in vivo evaluation. Patrick Couvreur is Full Professor of pharmacy at Paris-Saclay University, member of the Académie des Sciences, and holder of the chair of "Innovations Technologiques" (2010) at the "Collège de France." He is a Senior Member of the "Institut Universitaire de France" since 2009. Christian Serre is a CNRS research director. He was previously the group leader of Porous Solids team from Institut Lavoisier de Versailles, France. Christian then created in 2016 a new Institute of Porous Materials at the Ecole Normale Supérieure and Ecole Supérieure de Physique et de Chimie Industrielles from Paris. His main expertise lies in the discovery of new robust porous metal-organic frameworks (MOFs) and their applications in health, energy, and environment. Christian was among the pioneers in the field of biomedical applications of MOFs.

Figure 2 .

 2 Figure 2. Schematic illustration of the coating with polymers: I) PEG and II) Stp10-C. Reproduced with permission.[64] Copyright 2016, American Chemical Society.

Figure 1 .

 1 Figure 1. Surface modification of nanoMOFs directly affects colloidal stability before and after in vivo administration, drug release and burst effect, and in vivo targeting and prolonged circulation in the organism.

  Adv. Mater. 2018, 1707365

Figure 4 .

 4 Figure 4. In vitro and in vivo models employed to assess nanoMOFs' toxicity.

Figure 5 .

 5 Figure5. Iron level in different organs after 1, 7, and 30 d of i.v. administration of glucose solution of 220 mg kg -1 MIL-100(Fe) NPs in rats. Reproduced with permission.[52] Copyright 2013, Royal Society of Chemistry.

  Adv. Mater. 2018, 1707365

Figure 7 .

 7 Figure 7. A) Synthesis of Co/DPZ/C nanocarrier andB) theranostic strategy for MR imaging-guided multidrug chemotherapy and photothermal synergistic therapy. Reproduced with permission.[129] Copyright 2017, American Chemical Society.

Figure 8 .

 8 Figure8. Representative evolution on the number of articles on metalorganic frameworks for drug delivery since 2008. In PubMed, the initial research topic of "metal-organic framework" was further refined using the research topic of "drug delivery."

Table 3 .

 3 Size

	and surface charge of untreated and KF-treated MIL-100(Fe) NPs, obtained by
	microwave method, measured by DLS in different simulated physiological media at 37°C (SIF:
	pancreatin free simulated intestinal fluid, pH 6.8). 170
	MIL-100(Fe)	Medium	Size (nm)	ζ-potential (mV)
	nanoparticles
	Untreated	Mili-Q water	141±38	+14.5±2.3
	KF-treated	Mili-Q water	141±18	-21.5±3.8
		PBS	155±61	-31.5±0.0
		PBS + BSA	162±60	-9.9±0.0
		HCl pH 1.2	173±32	+22.5±2.2
		SIF	172±52	-32.6±2.1
		Low ionic SIF	147±54	-50.6±0.0
		Low ionic SIF + pancreatin	168±55	-6.7±1.8
		Low ionic SIF + mucin	254±40	-6.1±0.0

Table 2 .

 2 Recent advances on nanoMOFs in vivo studies.

	MOF	Study	Drug	Animal Model	Route	Ref
	MIL-88A(Fe)	Toxicity	-	Rat	Intravenous	90
	MIL-
	88B_4CH 3 (Fe)
	MIIL-100(Fe)	Toxicity	-	Rat	Intravenous
	MIL-88A(Fe)	Pharmacokinetics
	MIL-	Biodistribution
	88B_4CH 3 (Fe)
	Cu-BTC	Antiinflammatory	5-FU	Carrageenan test	Oral
		effect		induced peritonitis
				in mouse
	GdIII-pDBI	Toxicology	Doxorubicin	Mice	Intravenous
		parameters
	DPB-UiO-66	Photodynamic	-	Xenograft	Intratumoral
		therapy		subcutaneous tumor
				model in mice
	ZIF-8	Antitumoral	Doxorubicin	Xenograft	Intravenous
		Theranostic	Fe2O3	subcutaneous tumor model in mice
	9 different MOFs	Toxicity	-	Zebrafish embryo	Exposition to
					nanoMOF
					solution
	MIL-101	Antitumoral	Doxorubicin	Xenograft	Subcutaneous
				subcutaneous tumor	near the tumor
				model in mice	site
	CaZol	Antitumoral	Zoledronate	Xenograft	Intravenous
				subcutaneous tumor
				model in mice
	MIL-100(Fe)	Pharmacokinetic	Busulfan	Rat	Intravenous
		Biodistribution
	CD-MOF-1	Biodistribution	Ibuprofen	Mice	Oral
	Meso-MOF	Antitumoral	Doxorubicin	Xenograft	Intratumoral
				subcutaneous tumor
				model in mice
	ZIF-8	Antitumoral	Doxorubicin	Xenograft	Intratumoral
			Verapamil	subcutaneous tumor model in mice
	ZIF-8	Antitumoral	Camptothecin	Xenograft	Intravenous
			Doxorubicin	subcutaneous tumor model in mice
			Photosensitizer:
			CoFe2O4 nanoparticles
	MIL-100(Fe)	Antitumoral	Photosensitizer:	Xenograft	Intravenous

Table 5 attempts

 5 to summarize the main steps and some relevant considerations in MOF development, from physico-chemical characterization to in vivo evaluation.All this being said, one cannot forget that the nanomedicine field itself faces currently a challenging phase in which several important questions about the medical service provided by nanotechnology are being discussed224,225 . In particular, major critics are made about the models used for pre-clinical evaluation of nanomedicines. For instance, for the treatment of cancer

diseases, the excessive reliance on the mice xenograft animal models is contested because the data obtained with those models are hardly reproduced in humans. The use of extremely high doses in mice, impossible to extrapolate to humans, and the labeling with fluorescence probes to demonstrate a stealth or targeting effect, when it is known that these markers do not provide quantitative results, are other examples of irrelevant methodologies

225

. Carefully looking at the most recent in vivo studies with MOFs, here reviewed, one could address the same comments for most of them. In the next years MOFs researchers should therefore: (i) innovate and for instance move towards more reliable animal models, such as orthotopic tumor models, syngeneic mouse tumor models, including genetically engineered mice or patient-derived xenograft, in the case of cancer disease 226 , or the use of 3D in vitro models, (ii) stop assuming effects, such as EPR or active targeting without performing the experiments that scientifically demonstrate these phenomena, by including the adequate control groups or using labeling strategies that assure quantitative analysis of drug biodistribution, for instance, (iii) avoid the use of model drugs or fluorescent probes that do not possess the same physico-chemical properties than those of the real drug.

Table 4 .

 4 Barriers to overcome by the 3 rd generation drug delivery systems (adapted from 224 ).

	3rd generation drug delivery systems: barriers to overcome
	Formulation barriers (physico-chemical)	Biological barriers
	Increasing drug solubility	Lack of toxicity in vivo
	Control of drug release kinetics	Colloidal stability in the blood
	Control of drug loading	In vitro-in vivo correlation
	Control of therapeutic period	Long-term delivery
	Control of particle size, shape, functionality and	Non-invasive delivery
	flexibility	Controlling biodistribution
	Surface modification with ligands	Navigating microenvironment of diseased
	Stimuli-sensitive delivery systems	tissues to reach target cells
	Self-regulated delivery systems	Crossing mucosal barriers

The risk also exists to limit the in vivo studies to the reproduction of those experiments already performed with other nanocarriers (i.e., liposomes, lipid/polymer NPs or polymer micelles), without showing for nanoMOFs any ground breaking advantages. Moreover, until now, cancer treatment is almost the unique application of nanoMOFs in biomedicine when several other important biomedical applications may be considered, such as improved oral bioavailability for traditional drugs (see the work by Hartlieb et al.

207

), oral controlled release, non-viral gene delivery, vaccine adjuvant delivery systems or delivery across the blood brain barrier, among others.

Table 5 .

 5 Summary of the different stages and relevant points to considered to develop nanoMOFs for biomedical applications, from the physico-chemical characterization to the in vivo tests.

		Material properties	Relevant considerations
		•	Size	•	The physiological fluids in contact with nanoMOFs
		•	Shape	depend on the administration route and biodistribution
	Physico-chemical	• physiological fluids Surface charge • Degradation in physiological fluids • Identification of • Behavior in	• species	Composition of the physiological fluids varies within
		degradation products in
		physiological fluids
		•	Nanoparticles	•	Different models can be used for toxicity evaluation:
		cytotoxicity	cell lines, primary cells, 3D in vitro models. The choice should
		•	Toxicity of	be justified
		degradation products	•	Besides toxicity tests based on metabolic activity and
		•	Cellular uptake	cell proliferation, other deleterious effects of nanoMOFs may
	In vitro	• trafficking Intracellular • Intracellular	be identified, such as oxidative stress and proinflammatory cytokine induction • Identify the intracellular target of the drug and verify
		delivery		the ability of the nanoMOF to assure the delivery of the active
		•	Pharmacological	drug
		activity in cell culture	•	Depending on the administration route the capacity to
		•	Crossing biological	cross the biological barriers (i.e. gastrointestinal, endothelial,
		barriers		blood-brain barrier) should be proved

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Table 2 .

 2 Hydrodynamic diameter and surface charge of MIL-100(Fe) (milliQ water, c= 1 mg/ml).

	pH	Hydrodynamic	Hydrodynamic	ζ-potential	PDI
		diameter [nm] (%	diameter [nm] (%	[mV]
		Number)	Intensity)
	4.5	115±3	150±8	-36±2	0.23

Table 3 .

 3 Hydrodynamic diameter, surface charge and PDI of MIL-100(Fe) NPs uncoated and coated with PEG Graftfast, PEG NHS, β-CD-Ph, and β-CD-Ph-PEG at different pH (2.5, 3.5,4.5, 5.4, 6.5 and 7.4) in milliQ water at a concentration of 1 mg/ml.

	pH Hydrodynamic	Hydrodynamic	ζ-	PDI pH Hydrodynamic	Hydrodynamic	ζ-	PDI
		diameter -	diameter-	potential		diameter -	diameter-	potential
		Peak 1[nm]	Peak 2 [nm]	[mV]		Peak 1[nm]	Peak 2 [nm]	[mV]
		(% number)	(% number)			(% number)	(% number)
			MIL-100			MIL-100@PEG Graftfast
	2.5	262±2	>1000	12±2	0.70 2.5	655±21	>1000	14±2	0.80
		(60%)	(40%)			(81%)	(19%)
	3.5	113±2	-	-21±3	0.15 3.5	106±2	-	-24±2	0.22
		(100%)				(100%)
	4.5	115±3	-	-36±2	0.23 4.5	93±4	-	-42±1	0.23
		(100%)				(100%)
	5.4	106±2	-	-42±3	0.30 5.4	77±5	-	-45±1	0.30
		(100%)				(100%)
	6.5	258±16	>1000	-36±3	0.90 6.5	67±6	-	-45±1	0.35
		(77%)	(23%)			(100%)
	7.4	1002±25	-	-30±2	0.60 7.4	198±8	-	-42±1	0.44
		(100%)				(100%)
		MIL-100@PEG NHS			MIL-100@CD-Ph
	2.5	85±5	-	42±3	0.18 2.5	50±2	177±4	16±2	0.45
		(100%)				(97%)	(3%)
	3.5	101±2	-	20±1	0.18 3.5	96±2	-	-27±1	0.24
		(100%)				(100%)
	4.5	1500±130	-	-3±2	0.5 4.5	84±3	-	-29±1	0.27
		(100%)				(100%)
	5.4	87±4	-	-22±1	0.18 5.4	96±1	-	-40±2	0.34
		(100%)				(100%)
	6.5	76±3	-	-25±2	0.20 6.5	65±2	-	-33±3	0.32
		(100%)				(100%)
	7.4	75±4	-	-30±2	0.30 7.4	60±2	236±9	-31±3	0.44
		(100%)				(99%)	(1%)
		MIL-100@CD-Ph-PEG

Table 4 .

 4 Hydrodynamic diameter of MIL-100(Fe) uncoated or coated at different time: 0, 0.5, 1, 2, 3, 5, 7 and 24 h of contact with PBS (c= 1 mg/ml).

	Time	Hydrodynamic	Hydrodynamic	ζ-	PDI Time	Hydrodynamic	Hydrodynamic	ζ-	PDI
	(h)	diameter -	diameter-	potential		(h)	diameter -	diameter-	potential
		Peak 1[nm]	Peak 2 [nm]	[mV]			Peak 1[nm]	Peak 2 [nm]	[mV]
		(% number)	(% number)				(% number)	(% number)
			MIL-100				MIL-100@PEG Graftfast
	0	147±5	-	-24±2	0.22	0	140±6	-	-24±2	0.25
		(100%)					(100%)
	0.5	158±5	-	-22±3	0.22	0.5	144±6	-	-22±4	0.3
		(100%)					(100%)
	1	165±3	-	-22±2	0.22	1	123±9	-	-23±1	0.32
		(100%)					(100%)
	2	144±13	-	-21±1	0.26	2	109±36	290±8	-22±1	0.48
		(100%)					(98%)	(2%)
	3	175±20	-	-21±1	0.25	3	111±32	560±135	-22±1	0.48
		(100%)					(98%)	(2%)
	5	167±47	-	-20±2	0.28	5	91±9	425±10	-21±1	0.40
		(100%)					(97%)	(2%)
	7	141±12	320±10	-21±2	0.30	7	159±12	1000±18	-20±1	0.40
		(70%)	(30%)				(98%)	(2%)
	24	676±14	-	-19±2	0.20	24	148±19	480±80	-19±2	0.40
		(100%)					(92%)	(8%)

MIL-100@PEG NHS MIL-100@CD-Ph

	7	140±12	500±18	-20±1	0.40	7	137±3	70±10	-23±2	0.35
		(98%)	(2%)				(95%)	(5%)
	24	148±19	480±80	-19±2	0.40	24	162±5	75±10	-23±2	0.35
		(93%)	(7%)				(90%)	(5%)
		MIL-100@CD-Ph-PEG
	0	130±4	-	-25±2	0.31
		(100%)
	0.5	134±2	-	-23±2	0.21
		(100%)
	1	138±5	-	-23±2	0.20
		(100%)
	2	130±3	-	-23±2	0.31
		(100%)
	3	128±2	-	-23±2	0.31
		(100%)
	5	130±1	-	-19±2	0.30
		(100%)
	7	137±2	-	-23±2	0.30
		(100%)
	24	142±5	320±10	-23±2	0.32
		(95%)	(5%)
	0	130±8	-	-22±2	0.25	0	127±4	-	-25±2	0.30
		(100%)					(100%)
	0.5	135±6	-	-22±4	0.30	0.5	136±2	-	-23±2	0.20
		(100%)					(100%)
	1	142±9	-	-23±1	0.32	1	121±5	-	-23±2	0.20
		(100%)					(100%)
	2	121±16	-	-22±1	0.48	2	142±3	-	-23±2	0.31
		(100%)					(100%)
	3	130±18	360±56	-22±1	0.48	3	129±2	-	-23±2	0.31
		(98%)	(2%)				(100%)
	5	132±9	425±10	-21±1	0.40	5	138±1	-	-19±2	0.30
		(97%)	(2%)				(100%)
										110

Table 5 .

 5 Hydrodynamic diameter (nm) of NPs uncoated and coated with PEG Graftfast, PEG NHS, CD and CD-PEG in solution mimicking in vivo conditions (0.2 ml NPs in glucose, 5 mg/ml mixed with 1.5 ml filtrated rat serum).

	Sample name	Initial	Hydrodynamic	Hydrodynamic	ζ-potential
		concentration in	diameter (nm)	diameter (nm)	[mV]
		glucose	Peak 1	Peak 2
		(mg/ml)	(% Number)	(% Number)
	Serum alone		7	-	-7±1
			(100%)
	MIL-100	5	206±8	-	-5±2
			(100%)
		2.5	180±10	-	-6±1
			(100%)
	MIL-100 @PEG	5	175±20	-	-7±1
	Graftfast		(100%)
		2.5	200±8	-	-7±1
			(100%)
	MIL-100@PEG	5	189±10	-	-7±1
	Graftfast (washed only		(100%)
	after coating)	2.5	197±12	-	-7±1
			(100%)
	MIL-100@ PEG	5	146±11	-	-7±1
	Graftfast (washed only		(100%)
	after coating, 2x diluted
	reaction for coating)
	MIL-100@PEG NHS	5	180±20	-	-5±2
			(100%)
		2.5	171±9	-	-6±1
			(100%)
	MIL-100@CD-Ph	5	162±10	-	-7±1
			(100%)
		2.5	183±8	42±9	-7±1
			(90%)	(10%)
	MIL-100@CD-Ph	5	173±6	-	-7±1
	(washed only after		(100%)
	coating)	2.5	183±9	-	-7±1
			(100%)
	MIL-100@CD-Ph-PEG	5	(100%)	-	-7±1
		2.5	187±5	68±4	-7±1
			(80%)	(20%)

  Lung NPs accumulation of uncoated NPs was observed already at 30 minutes after injection (50.1±8.4 % of the injected dose for 14 C-BTC). This accumulation was reduced approximately by a factor of 2 (18.6±1.2 % of injected dose) after 6h. And after 24 h, the pulmonary accumulation was still reduced to reach only 13.2±6.3 % of the injected dose. The level of BTC of uncoated NPs still remained in this tissue after 24 h, which was in good agreement with previous study1 , where we found that after disaggregation of NPs, individual NPs could still be retained and internalized into the cells. Noteworthy, coating of the outer surface of NPs reduced significantly the accumulation of the NPs in the lungs 30 minutes after

injection (23.6±2.4, 29.9±5.2, 19.1±9.5 and 3.8±1.3 % of injected dose only for PEG Graftfast,

Table 6 .

 6 Summary of synthesis conditions of MIL-100(Fe) (addition of NaOH, acetic acid or mixture of ethanol : water).

	Conventional, RT synthesis with NaOH, acetic acid additives or ethanol : water mixture
	Concentration Temperature (°C) Reaction time	Solvent NaOH/linker	Acetic
			(h)			acid/Fe
	c1, c2 or c3	RT	48	H 2 O	0.5	-
					1
					1.5
					2
					2.5
	c1, c2 or c3	RT	48	H 2 O	-	2.5
						5
						10
						15
						20
	c1	RT	48	ethanol :	-	-
				water
				10:90
				30:70
				50:50
				70:30
				90:10

Table 7 .

 7 Summarized conditions of MIL-100(Fe) synthesis obtained by the microwave method.

		Microwave -assisted synthesis
	Concentration	Reaction time (minutes) Temperature (°C)
	c1, c2 or c3	6	60
		30
		60

Table 8 .

 8 Hydrodynamic diameter, ζ-potential and PDI of MIL-100(Fe) obtained by different synthesis routes, measured at pH 5.4 at concentration 1 mg/ml in milliQ water.

	Name of sample	Hydrodynamic	Hydrodynamic	ζ-potential [mV]	PDI
		diameter -Peak 1	diameter-Peak 2
		[nm]	[nm]
		(% number)	(% number)
	MIL-100(Fe)	112±8	-	-40±2	0.25
	(RT, 48 h, c1)	(100%)
	MIL-100(Fe)	96±2	-	-29±1	0.20
	(60°C, 48 h, c1)	(100%)
	MIL-100(Fe) + 0.5	77±5	-	-29±1	0.21
	NaOH/linker	(100%)
	(RT, 48 h, c2)
	MIL-100(Fe) + acetic	60±6	-	-31±2	0.22
	acid/Fe=2.5	(100%)
	(RT, 48 h, c1)
	MIL-100(Fe)	54±5	200±8	-29±1	0.35
	(MW, 60°C, 1 h, c1)	(98%)	(2%)
	MIL-100(Fe)	91±3	-	-25±2	0.23
	(MW, 60°C, 1 h, c2)	(100%)
	MIL-100(Fe)	52±2	-	-38±3	0.21
	(MW, 60°C, 1 h, c3)	(98%)

Table of Contents 1

 of1 Methotrexate (MTX) is a small molecule chemotherapeutic agent (see Figure1) that induces cell apoptosis by inhibiting dihydrofolate reductase and disrupting DNA synthesis. Many cancer types are sensitive to MTX treatment; however, the application of MTX is limited by its poor pharmacokinetics, low tolerated dose, and resistance.1 MTX has, however, many sides effect on patients, including nausea and vomiting, diarrhea, stomach pain, hair loss, tiredness, dizziness, headache, fever, lung disease, and liver disease. Moreover, because of the possibility of severe toxic reactions (which can be fatal, observed in elderly patients), it is recommended to use MTX only in neoplastic diseases, or in patients with severe, recalcitrant, disabling psoriasis, or rheumatoid arthritis that are not adequately responsive to other forms of therapy (medical information by Pfizer). Several nanocarriers have been proposed so far to overcome these severe side effects of MTX treatment to ensure a targeted distribution. Generally, loading challenging drugs in the nanoparticles can actively reduce the dosage of anticancer drugs, and when combined with great targeting abilities, it can highly increase therapeutic efficiency together with a reduction of side effects.2

	1. Introduction
	. Introduction ……………………………………………………………………………145
	2. Encapsulation and release from MIL-100(Fe) ………………………..……….146
	2.1. Experimental procedure………………………………………………………………146
	2.2. Results………………………………………………………………………................147
	3. Application to core-shell nanoMOFs -USPIO………………………………...152
	3.1. Introduction…………………………………………………………………………...152
	3.2. Contribution…………………………………………………………………………...153
	3.3. Results…………………………………………………………………………………153
	3.3.1. Synthesis and characterization…………………………………………………153
	3.3.2. MTX encapsulation and release………………………………………………..157
	3.4. Conclusion…………………………………………………………………………….161
	4. Toxicity evaluations on the C. elegans…………………………………………...162
	4.1. Introduction…………………………………………………………………………...162
	4.2. Contribution…………………………………………………………………………...163
	4.3. Results…………………………………………………………………………………163
	4.3.1. Scale-up synthesis of MIL-100(Fe) NPs………………………………………..163
	4.3.2. Encapsulation and release of MTX……………………………………………..165
	4.3.3. Evaluation of the toxicity effect of MIL-100(Fe) on C. elegans………………..170
	4.3.4. Evaluation of the toxicity effect of MTX@MIL-100(Fe) on C. elegans
	reproductive system…………………………………..…………………………...172
	4.4. Conclusions……………………………………………………………………………174

Table 1 .

 1 Summary of the encapsulation conditions of MTX into MIL-100(Fe) NPs.

	5, 1, 2,

Table 2 .

 2 Encapsulation

	Name of sample	Encapsulation (wt%)	Encapsulation (wt%)	Encapsulation (wt%)
		UV-Vis	TGA theoretical	TGA experimental
	MTX@MIL-100, MIL-	38±4	41±4	35±4
	100(Fe)/MTX 1:1 weight ratio
	MTX@MIL-100, MIL-	50±5	50±3	44±3
	100(Fe)/MTX 1:3 weight ratio

of MTX inside MIL-100 (after incubation in HCl 10 mmol/l) in weight ratio of nanoMOFs to MTX 1:1 and 1:3 (37°C, overnight), calculated from UV-Vis and TGA analysis.

  Considering the size of the MTX molecule and the size of the windows in MIL-100, one expects MTX to only fill the larger cages that are accessible through larger hexagonal windows of ≈8.6 Å free aperture. As the BET of the MTX loaded sample becomes negligible, it is likely that a few MTX molecules block the access the pentagonal windows of the smaller cages (≈5.6 Å free aperture), preventing from any nitrogen adsorption.

			600
		Qantuity Adsorbed (mmoml/g)	100 200 300 400 500			MIL-100(Fe) MTX@MIL-100(Fe)
			0
			0.0	0.2	0.4	0.6	0.8
					Relative Pressure (p/p°)
	B	Name of sample		BET surface are (m 2 /g)	Encapsulation rate (wt%)
		MIL-100(Fe)		1570	-
		MTX@MIL-100(Fe)		95	~54
	Figure 5. N 2 porosimetry (77 K, P 0 =1 atm) of MIL-100(Fe) before MTX encapsulation (red) and

after MTX encapsulation (MIL-100(Fe)/MTX 1:3 weight ratio, encapsulation of 54 wt%) (blue), with BET surface area 1570 m 2 /g and 95 m 2 /g for MIL-100(Fe) and MTX@MIL-100(Fe), respectively. Activation conditions: 150°C under vacuum for 6 h.

Table 6 .

 6 Encapsulation (wt%) of MTX in MIL-100(Fe), in HCl 10 mmol/l or milliQ water, calculated from UV-Vis and TGA.

	Name of sample	Encapsulation (wt%)	Encapsulation	Encapsulation
		UV-Vis	(wt%)-theoretical	(wt%)-experimental
			TGA	TGA
	MTX@MIL-100(Fe),	50±5	47±3	38±6
	encapsulated in HCl
	10mmol/l
	MTX@MIL-100(Fe),	33±3	35±5	28±5
	encapsulated in milliQ
	water

Table of Contents Annex 1. Summary of the manuscript………………………………………………...184

	ANNEX 1:
	Annex 2. Synthesis protocols …… …………………………………………………….187
	Annex 3. Additional characterization……………………….……………..………….194
	Annex 4. "Nanoparticles of Metal-Organic Frameworks: on the road to in
	vivo efficacy in biomedicine"……………………… ………………………….………..208

SUMMARY OF THE MANUSCRIPT Summary-English Title: Surface engineered hybrid nanocarriers for cancer treatment

  Cancer is one of the biggest issues of the 21 st century, with high death rate in Europe. Scientists need to improve current available cancer therapies, to overcome side effects due to high dosage of anticancer drugs during chemotherapy. One way is to use nanoparticles (NPs) as drug nanocarriers. Nanoparticles need to fulfill criteria, such as: low-toxicity, biodegradability, possible high drug entrapment and controlled release, size <150 nm, long circulation in the blood and reduced recognition by macrophages. One candidate, Metal-Organic Frameworks (MOF), porous crystalline materials, has gained increased interest in the last decade for biomedical applications. A particularly attractive MOF is the mesoporous iron (III) carboxylate MIL-100(Fe)

Titre : Nanovecteurs hybrides modifiés en surface pour le traitement du cancer

  'accumuler dans les organes après administration intraveineuse. La capture pulmonaire des NP après administration intraveineuse a pu être utilisée précédemment pour traiter le cancer du poumon chez la souris. Cependant, pour le traitement de cancers situés au niveau d'autres organes, ce comportement constitue un inconvénient. Dans cette thèse, nous nous sommes concentrés sur les NP MIL-100(Fe). Nous avons étudié l'impact de différentes voies de modification de surface (covalente et non covalente) sur la stabilité des NP en milieu physiologique (pour imiter les conditions in vivo) et sur la capacité des enrobages à ajuster la biodistribution in vivo et la pharmacocinétique des NP après administration intraveineuse chez la souris. Nous avons sélectionné différentes stratégies de modification de surface : PEG (voie NHS), Graftfast PEG, 𝛽-Cyclodextrin-Phosphate et 𝛽-Cyclodextrin-Phosphate-PEG, qui ont déjà été signalées et ont montré des résultats positifs in vitro, mais n'ont jamais été étudiées in vivo. En particulier, les NP fonctionnalisées par 𝛽-Cyclodextrine-Phosphate-PEG ont montré une réduction de l'accumulation des NP dans les poumons. Toutefois, l'accumulation dans le foie et la rate s'est avérée être plus élevée que celle des NP non enrobées. Ces NP ont une forte charge de surface négative et ont donc tendance à se repousser les unes les autres et forment moins d'agrégats. Nous pouvons donc déduire que les NP modifiées par 𝛽-Cyclodextrine-Phosphate-PEG dans le sang sont suffisamment petites pour échapper à processus d'accumulation dans les poumons. Mais, cela ne permet pas d'échapper à la , 1 minutes). MIL-100(Fe)@γFe 2 O 3 was stored in absolute ethanol. Before further experiments, MIL-100(Fe)@γFe 2 O 3 was washed with water. Immediately 14 C-trimesic acid (87.84 µl, 8.78 µCi) was added to the mixture. The solution was then left under magnetic stirring at room temperature at speed 600 rpm for 48 h. Product was then transferred into Eppendorf tubes (1.5 ml), then centrifuged (14500 rpm, 10 min) and washed once with distilled water, and twice with ethanol by re-dispersion (sonication tip, 20%, 20 seconds) and centrifugation (14500 rpm, 10 min). The resulting NPs were kept in absolute ethanol in the fridge. NPs were prepared several times.

	14 C-MIL-100(Fe) NPs preparation
	develop smaller MIL-100(Fe) NPs, more suitable to cross cell membranes with possible In a 20 ml glass vial, trimesic acid (29,24 mg, 0.14 mmol) and iron (III) nitrate nonahydrate
	reduction in macrophage recognition. The last part of the thesis was devoted to the encapsulation (84.3 mg, 0.21 mmol) were mixed with distilled water (10.54 ml) under magnetic stirring during
	and release of an anticancer drug, Methotrexate (MTX), into MIL-100(Fe) NPs for two types of potential applications. The first use of MTX was established within core-shell nanoMOFs-USPIO (Ultra-Small Particles of Iron Oxides) in a view of theranostics. A second potential application was focused on the toxicity evaluation of MTX-encapsulated or bare nanoMOFs, based on C.elegans worms. 2 minutes. Encapsulation
	Résumé -Francais
	Le cancer est l'un des grands problèmes du 21 ème siècle, provoquant des taux de mortalité élevés
	en Europe. Les scientifiques doivent améliorer les thérapies anticancéreuses actuellement disponibles, afin de surmonter les effets secondaires dus aux fortes doses de médicaments anticancéreux pendant la chimiothérapie. Une façon d'y parvenir est d'utiliser des nanoparticules (NP) comme nanovecteurs de médicaments. Les nanoparticules doivent répondre à des critères tels que : faible toxicité, biodégradabilité, risque élevé de piégeage et de libération contrôlée, taille <150 nm, longue circulation dans le sang et reconnaissance réduite par les macrophages. L'un des candidats, les Metal-Organic Frameworks (nanoMOF), un matériau cristallin poreux, a suscité un intérêt croissant au cours de la dernière décennie en biomédecine. En particulier, les nanoparticules du carboxylate de fer mésoporeux MIL-100(Fe) (MIL : Materiaux de l'Institut Lavoisier), sont particulièrement intéressantes. Le MIL-100(Fe) peut, en effet, encapsuler une grande quantité de médicaments et permettre une libération contrôlée dans des conditions physiologiques. Cependant, les NP de MIL-100(Fe) ont tendance à s'agréger dans les milieux organes du système réticulo-endothélial (RES), comme le foie ou la rate, probablement en raison de leur taille encore trop grande et/ou d'un recouvrement non-homogène par les groupements PEG. Donc, nous avons entrepris de développer des NP MIL-100(Fe) plus petites, plus aptes à traverser les membranes cellulaires et à échapper à la reconnaissance par les macrophages. La dernière partie de la thèse a été consacrée à l'encapsulation et à la libération du méthotrexate (MTX), un médicament anticancéreux, dans des NP de MIL-100(Fe) pour deux types d'applications potentielles. La première a été établie dans le cadre de particules coeur-coquille magnétiques nanoMOFs-USPIO (Ultra Small Particles of Iron Oxide). Une deuxième application potentielle a concerné l'évaluation de la toxicité des nanoMOFs nus ou encapsulant du MTX, physiologiques et à sreconnaissance des macrophages et à leur élimination rapide de la circulation sanguine vers les pour des vers C.elegans.

amplitude

of 3 H-Gemcitabine monophosphate in 14 C-MIL-100(Fe) NPs

	In one Eppendorf, 14 C-MIL-100(Fe) NPs (8 mg) from the larger batch (NPs stored in ethanol)
	were first centrifuged (14500 rpm, 10 min) and washed with water by re-dispersion (sonication
	tip, 20%, 20 seconds) and centrifugation (14500 rpm, 10 min). These NPs were then re-dispersed
	in milliQ water (1.8 ml). In a second Eppendorf, a powder of Gemcitabine monophosphate (2.88
	mg) was prepared. Immediately after redispersion of NPs, NPs solution was added to
	Gemcitabine monophosphate, and the 3 H-Gemcitabine monophosphate (9.6 µl, 19.2 µCi)
	solution in ethanol was added to this mixture. NPs were kept under stirring overnight. Loaded
	NPs were finally recovered by centrifugation (14500 rpm, 10 min) and washed once in milliQ
	water. For injections, NPs were re-dispersed with glucose 10% (1.6 ml) (sonication 20%, 10
	seconds). Typically, encapsulated NPs were prepared one day before planned injections and
	further surface functionalization was performed on the day of injection.
	When injecting only free Gemcitabine monophosphate, Gemcitabine monophosphate (2.4 mg)

  mg/ml cholesterol in ethanol, 1 mol/l KPO 4 buffer pH 6.0 (108.3 g KH 2 PO 4 , 35.6 g K 2 HPO 4 , H 2 O to 1 liter), and 1 mol/l MgSO 4 and 1 mol/I CaCl 2 .

	Theoretical:
	650.8 g/mol	239.53 g/mol
	100%	x
	x= 36.8 % (oxides theoretical)
	y= 100% -36.8%= 63.2 % (ligand loss theoretical)
	Experimental:
	We check weight loss (%) at temperature 150°C, when water is already evacuated from MIL-
	100(Fe) (flat area), this point is a new 100%
	88% à 100%
	28% à 31.8 % (oxides practical)
	100% -31.8%= 68.2% (ligand loss practical)

Table 2 .

 2 Covalent and noncovalent strategy for surface modifications of nanoMOFs (CD-P: cyclodextrin-phosphates; Py-PGA-PEG: pyrenederived polyethylene glycol).

	Interactions	Type of interactions	Example
	Noncovalent	van der Waals	Heparin
		Electrostatic	Chitosan
		π-π stacking	Hyaluronic acid
			Py-PGA-PEG
	Covalent	Coordination	CD-P
		Condensation	Amino-polymer (amino-PEG)
		"Click chemistry"	Oligonucleotides
			PEG-RGD-β-CD

C-liver

C-spleen 3 H-spleen

H-NP-Gem MP-CD-PEG 14 C-NP-Gem MP-CD-PEG

© 2018 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

Adv. Mater. 2018, 1707365

www.advmat.de www.advancedsciencenews.com

Acknowledgments 1998, 20, 45-52.

Abbreviations

4T1

Murine mammary carcinoma cell line 5-FU 5-Fluorouracil

CHAPTER 2

Biodistribution and pharmacokinetics study of surface engineered nanoparticles of MIL-100(Fe)

Pharmacokinetics, biodistribution and metabolic tests

For biodistribution (BD) and pharmacokinetics (PK) studies, we used radiolabeled compounds to estimate the linker content ( 14 C-trimesic acid ( 14 C-BTC)) resulting from NPs and/or the Gemcitabine monophosphate drug ( 3 H-Gemcitabine monophosphate ( 3 H-Gem MP)). Therefore, all preparations of NPs for in vivo study, as well as, the organs and blood preparations for The equipment available within the radioactive rooms was not exactly the same as in a non radioactive environment; therefore, some adjustments in comparison with the usual protocols were needed. For instance, centrifugation at high speed (14500 rpm) was only possible with Eppendorf tubes (1.5 ml). With bigger tubes, the speed of centrifugation was limited to 4000 rpm, which was not enough for NPs separation from the supernatant. Therefore, I prepared smaller batches of NPs, and used 1.5 ml Eppendorf tubes for centrifugation.

General protocol and synthesis of NPs for in vivo experiments

In total, 154 females C57BL/6 mice (~7-8 weeks old, 20±3 g) were employed. One week before the experiments, animals were adaptatively housed, with controlled temperature (22±2°), humidity, light/dark cycles and food, and water supply ad libitum. Animals were randomly assigned to 6 different groups (24 animals/group):

1) 3 H-Gemcitabine monophosphate free ( 3 H-Gem MP free)

2) 14 C-MIL-100(Fe)@ 3 H-Gem MP (radiolabeled (nanoMOFs + drug))

3) 14 C-MIL-100(Fe)@ 3 H-Gem MP@polyethylene glycol Graftfast ( 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG Graftfast) (PEG Graftfast functionalized radiolabeled (nanoMOFs + drug)) 4) 14 C-MIL-100(Fe)@ 3 H-Gem MP@polyethylene glycol NHS ( 14 C-MIL-100(Fe)@ 3 H-Gem MP@PEG NHS) (NHS PEG functionalized radiolabeled (nanoMOFs + drug))

CHAPTER 3

The case of the anticancer drug Methotrexate (MTX) The idea of CeleScreen is to fulfill the global need by developing an alternative testing platform based on a patent invented by Dr. Philippe Manivet (EP2775302 2013). This method implements nanocarriers for feeding C. elegans with xenobiotics in order to test for their toxicity. This method favors the ingestion of almost all xenobiotics, whatever the appetence of the nematode for each, and moreover, at a lower dose compared to earlier trials. This is very important as it allows for assaying molecules at a more "physiological" condition.

C. elegans has been studied as a toxicity model since 1960's because of their numerous advantages: (a) inexpensive and easy, (b) rapid lifecycle, (c) short lifespan, (d) conservation with human counterpart, about 40% of worm genes have human orthologs, which demonstrated high conservation of biological mechanisms between the worm and human. [10][11][12][13][14][15] Only information relevant to the study will be included. C. elegans can be maintained in the laboratory where it is grown on agar plates or liquid cultures (typically Nematode Growth Media, NGM) with

Escherichia coli (E. coli) OP50 as food source. C. elegans is a rapidly growing worm, and its entire life cycle, from an egg to an adult, takes only 3.5 days in the presence of the food at 20-25°C. After embryogenesis, C. elegans progresses through four larval stages (L1 to L4) before becoming an adult. The life cycle of C. elegans at 22 C is presented in Figure 15. 16,17

Synthesis of γFe 2 O 3 nanoparticles

Prior to the preparation of the core-shell MIL-100(Fe)@γFe 2 O 3 composites, iron oxides nanoparticles were first synthesized by a co-precipitation method. 20 ml of sodium hydroxide (15 mg/ml) were added under vigorous stirring to an acidic aqueous solution of iron (III) chloride hexahydrate (20 ml, 1 mg/l) and iron (II) chloride tetrachloride (5 ml, 2 mg/l in HCl 2 mol/l). The black precipitate was isolated by magnetic setting and washed with 20 ml of water.

After a magnetic separation, the precipitate was then stirred for 15 minutes in 30 ml of HNO 3 (2 mol/l). For a complete oxidation of the maghemite NPs, the precipitate obtained after a magnetic separation was then mixed with 10 ml of iron (III) nitrate nonahydrate (0.35 mol/l) at 80°C for 30 minutes. The precipitate isolated is mixed with water and colloidal solution was obtained.

Synthesis of MIL-100(Fe)@γFe 2 O 3 nanoparticles

The MIL-100(Fe)@γFe 2 O 3 composites were prepared by adding under stirring the colloidal solution containing 70 mg of iron oxides to an aqueous solution of the iron (III) nitrate nonahydrate (70 ml, 0.025 mol/l). Solution was further sonicated by ultrasonication tip (10% amplitude, 1 minutes), for the better dispersion. 250 mg of trimesic acid were then added and the reaction stirred for 24 h at room temperature (20-23°C). The brown precipitate was isolated by magnetic settling and then centrifuged at 14500 rpm for 10 minutes. The solid was then washed once in water and three times in absolute ethanol by centrifugation and redispersion (10%

Preparation and characterization of the organs

Prior to analysis organs were frozen. These were cut into smaller pieces (around 50 mg). The weight of the organ piece was carefully checked and placed in the plastic vials of 20 ml. Next, 1 ml of solution "Solvable" was added to the organ pieces and samples were placed in an oven at 50-60°C for 4 h. After this period, organs should be dissolved. Samples were taken out and when temperature was cooled down to room temperature, 0.1 ml of hydrogen peroxide 30% was added to the samples and carefully mixed. Another 0.1 ml of hydrogen peroxide 30% was then added again and mixed with the previous mixture. Samples were then placed in the oven for another 30 minutes at 50-60°C. After that period, samples were taken out and when the sample was cooled down, 10 ml of solution of "Ultima Gold" was added to the liver and spleen samples, while 10 ml of solution of "Hionic Fluor" was added to the kidneys, lungs, heart, fat or muscles samples.

All samples were finally well mixed and after 10 minutes placed into the scintillator sample holder.

Preparation and characterization of the blood

Prior to analysis, blood samples were centrifuged and stored frozen for further characterization.

0.1 ml of blood was placed in 2 ml plastic vials. Then, 1 ml of a "Solvable" solution was added to the blood and samples placed in the oven at 60°C for 1 h. After taking the sample out, when temperature was cooled down to RT, 0.1 ml of solution of EDTA-Na 2 0.6 M was added to the

Calculations:

Below are noted molecular masses, which are necessary for this calculation:

Encapsulation of MTX-theoretical

Below is presented the way the encapsulation results were calculated from TGA data.

Below are noted molecular masses, which are necessary for these calculations:

Encapsulation of MTX-experimental

Below are presented the way the practical encapsulation results were calculated from TGA data.

We compared samples of MTX@NPs with NPs. For an easier calculation, the graph with liposomes for the delivery of the anticancer drug doxorubicin (Dox). [57] There is another product consisting of PEGpolylactic-co-glycolyc acid NPs for the delivery of paclitaxel, Genexol-PM, already approved by Korean FDA and currently in clinical trials in the US. [58] NPs coating with PEG prolongs the circulation time in the blood, allowing their distribution in other organs. Besides, "active" targeting implies the incorporation of molecules on the surface of the nanoMOFs that are specifically designed to recognize particular receptors or antigens expressed onto the membrane of the diseased cells and tissues. [51,59,60] In general, surface coating of NPs should fulfill some criteria to be suitable for drug delivery, [37,61] such as a) biofriendly synthesis method, without any toxic additive, b) avoid intrusion of the targeting molecule into the nanocarrier, c) preserve drug release capacity, d) lack of interference with entrapped drugs, e) improvement of particle colloidal stability, and f) stability under physiological conditions. Furthermore, in the case of nanoMOFs, the presence of a high porosity makes the surface modification even more challenging due to possible nonspecific intrusion of the molecules into the pores, blocking them, and/or decreasing drug release capacities. [61] Not specified [137] www.advmat.de www.advancedsciencenews.com and the identification of the best animal model for the in vivo preclinical evaluation. Table 4 attempts to summarize the main steps and some relevant considerations in MOF development, from physicochemical characterization to in vivo evaluation.

All this being said, one cannot obviate that the nanomedicine field itself currently faces a challenging phase in which several important questions about the medical service provided by nanotechnology are being discussed. [141,142] In particular, major critics are made about the models used for preclinical evaluation of nanomedicines. For instance, for the treatment of cancer diseases, the excessive reliance on the mice xenograft animal models is contested because the data obtained with those models are hardly reproduced in humans. The use of extremely high doses in mice, impossible to extrapolate to humans, and the labeling with fluorescence probes to demonstrate a stealth or targeting effect, when it is known that these markers do not provide quantitative results, are other examples of irrelevant methodologies. [142] Carefully looking at the most recent in vivo studies with MOFs, here reviewed, one could address the same comments for most of them. In the next years, MOFs' researchers should therefore i) innovate and for instance move toward more reliable animal models, such as orthotopic tumor models, syngeneic mouse tumor models, including genetically engineered mice or patient-derived xenograft, in the case of cancer disease, [143] or the use of 3D in vitro Adv. Mater. 2018, 1707365 Table 3. Barriers to overcome by the third-generation drug delivery systems (adapted from ref. [141]). • Distinguish between NP and the drug (i.e., pharmacokinetic and biodistribution)

• When using a drug model justify the choice (hydrophobicity, molecule size, reactive groups, etc.). Idem when using a fluorescent probe • To describe toxicity consider different aspects such as animal behavior, hematological analysis, biochemistry, histological study • Distinguish between acute toxicity and chronic toxicity • Toxicity studies should mimic the therapeutic dosage schedule • Use the adequate control groups to prove efficacy (empty carrier, free drug). If they are missing give a scientific justification • When multiple parameters are evaluated within a unique delivery system (i.e., surface modification and drug release, co-delivery of drugs) add the corresponding controls (i.e., empty carriers with and without surface modification; free drugs separately and in combination, and loaded drugs separately and in combination).

Conflict of Interest

The authors declare no conflict of interest.