Docteur

Biotechnologie

Jungyoon Choi

» Thèse

Jean-Luc Coll
email: jean-luc.coll@univ-grenoble-alpes.fr

DUPERRAY M Alain

LADAVIERE Mme Catherine

GAGGIOLI M Cédric

BOTURYN M Didier

GARANGER Mme Elisabeth

Mme Isabelle

COLL M Jean-Luc

Emilie Rustique

Maxime Henry

Mélanie Guidetti

Véronique Josserand

Lucie Sancey

Jérôme Boutet

THÈSE Pour obtenir le grade de

Keywords: integrin alpha v beta 3, Neuropilin-1, lipid nanoparticles, silica nanoparticles, prostate cancer, nanomedicine, cancer imaging, dual targeting integrin alpha v beta 3, Neuropilin-1, lipid nanoparticles, silica nanoparticles, prostate cancer, nanomedicine, cancer imaging, dual targeting Cancer imaging Theranostic Integrins RGD Nanoemulsion NIR imaging

Development of theranostic molecules targeted to prostate cancer

List of Figures

Project 1, cRGD-LNP

We initially tested two different synthesis methods to produce the targeted LNPs:

(1) using maleimide-labeled peptides on the LNPs or [START_REF] Matsumura | A new concept for macromolecular therapeutics in cancer chemotherapy: mechanism of tumoritropic accumulation of proteins and the antitumor agent smancs[END_REF] starting with maleimidelabeled LNPs and the peptides. All synthesis were performed by Emilie Rustique in CEA-LETI and I verified the LNPs' functionality at first.

Characteristics of LNPs

Lipid nanoparticle (LNP), initially called LipImage™ 815, was synthesized following the previously described method (Jacquart et al., 2013). LNPs are composed of an oily core mixed with oleyl-IR780 near-infrared (NIR) dyes and a PEGylated phospholipid monolayer which encapsulates the core (Figure 15). To conjugate the peptides on the surface of LNPs, PEG surfactants and the peptides were linked by the maleimide-thiol coupling reaction: R-maleimide + R'-SH  R-(thiol ether bond)-R'. The conjugation was performed in two directions in order to select the better one. In the first method (Fuctionalization1, F1), maleimide-labeled peptides were reacted with SH-labeled LNPs, while in the second method (Functionalization2, F2) maleimide-LNPs were reacted with SH-peptides. cRGD-LNP (αvβ3-targeting), cRAD-LNP (no affinity for αvβ3 due to a nonsense peptide sequence with alanine instead of glycine), and standard/OH LNP (negative control without peptide) were generated (Figure 15C).

The F1-LNPs were slightly larger than the F2-LNPs; the hydro-dynamic diameter of F1-LNP was 50-55 nm, whereas the F2-LNP was approximately 47 nm, measured by DLS. Zeta potentials were similar amongst all the LNPs being in a range of -4 to -9 mV. Based on the number of dye molecules and of particles, we calculated that each LNP was loaded with 104 to 112 molecules of dye/particle (Table 7). Unfortunately, exact numbers of ligands that are present on the LNPs could Targeting tumors with cyclic RGD-conjugated lipid nanoparticles loaded with an IR780 NIR dye: In vitro and in vivo evaluation

Introduction

There has been a lot of effort to generate nanovectors suitable for the delivery of contrast agents and drugs. Several drug delivery systems are approved by the FDA, but only a few targeted formulations are currently tested in humans (Shi et al., 2016). Among the different formulations, Lipid Nanoparticles (LNPs) are promising in particular because they are showing: 1) excellent biocompatibility, 2) long circulation time in blood, 3) passive tumor accumulation via the enhanced permeability and retention (EPR) effect, 4) capacity of loading large amount of poorly soluble drugs as well as imaging probes, 5) possible chemical surface modifications that may help for active specific receptor targeting (Goutayer et al., 2010;Gravier et al., 2011;Hirsjarvi et al., 2013;Jacquart et al., 2013;Merian et al., 2015). LNPs are made of a lipophilic core encapsulated by a PEGylated mono-layer of phospholipids.

LNPs loaded with a lipophilic-derivative of IR780 dye (addition of a C18 chain to IR780 iodide dye) are stable for 6 months in phosphate buffer saline at 4 C in the dark. They present a longlasting EPR-mediated retention in subcutaneous tumors (Jacquart et al., 2013) due to the presence of a PEG-coating. Unfortunately, PEGylation is also lowering their interaction with cell membranes and this is finally ending-up with a low efficiency of intracellular delivery of the cargo [START_REF] Bozzuto | Liposomes as nanomedical devices[END_REF]. We expect that the presence of a targeting ligand should improve this final step.

The cyclic Arg-Gly-Asp (RGD) peptide known to target preferentially the a v b 3 integrin (Dechantsreiter et al., 1999;[START_REF] Haubner | Structural and functional aspects of RGD-Containing cyclic pentapeptides as highly potent and selective integrin aVb3 antagonists[END_REF] has been largely used for this purpose as recently reviewed in Arosio et al. [START_REF] Arosio | Advancement in integrin facilitated drug delivery[END_REF]. RGD was used in particular for applications such as cancer imaging, radiotherapy, phototherapy and nanomedicine [START_REF] Abdollahi | Inhibition of alpha(v)beta3 integrin survival signaling enhances antiangiogenic and antitumor effects of radiotherapy[END_REF][START_REF] Chen | Cyclic RGD peptide-modified liposomal drug delivery system: enhanced cellular uptake in vitro and improved pharmacokinetics in rats[END_REF][START_REF] Kunjachan | Nanoparticle mediated tumor vascular disruption: a novel strategy in radiation therapy[END_REF]Liu and Wang, 2013;Ray et al., 2011;Shuhendler et al., 2012;Wang et al., 2014;Yan et al., 2016;Yuan et al., 2015). Our group also described cyclic-RGDfKdecorated LNPs loaded with DiD, a dye adapted to Near-infrared in vivo imaging (Goutayer et al., 2010). This cRGD-LNP showed a good tumor accumulation and longer retention in tumors. We then generated Indocyanine Green (ICG)-containing LNP because ICG is approved for human injection, but ICG was leaking out rapidly from the LNP in vivo (Gravier et al., 2011;Merian et al., 2015). We thus finally generated stable and bright LNPs containing IR780-Oleyl modified dye, which turned out to be suitable for in vivo imaging (Jacquart et al., 2013).

In the present study, we developed the IR780-Oleyl-LNPs targeted by cyclic RGD. We demonstrate that cRGD-LNP are functional because they bind to integrin a v b 3 and are internalized.

However, despite these satisfying results in vitro, the presence of cRGD targeting does not augment the global accumulation of targeted versus non-targeted NP.

Materials and methods

Materials

Suppocire NB TM (semi-synthetics C12-C18 saturated triglyceride) was purchased from Gattefossé (Saint-Priest, France); Lipoid S75 (soybean lecithin at >75% phosphatidylcholine) from Lipoid (Ludwigshafen, Germany); Myrj TM S40 (polyethylene glycol 40 stearate); Super refined soybean oil from Croda Uniqema (Chocques, France); Cyclic RGD and cyclic RAD from Pepscan Presto (Lelystad, Netherland). IR780 iodide dye ( Cell culture media and buffers were purchased from Gibco (Life technologies, Paisley, UK); Bovine serum albumin (BSA) from Interchim (Montluçon, France); Puromycin dihydrochloride from Sigma-Aldrich (USA); Recombinant human vitronectin (rhVTN-N, truncated) and Geneticin (G418 sulfate) from Life Technologies. Unless otherwise specified, phosphate buffer saline (PBS) and Hanks' Balanced Salt Solution (HBSS) are always supplemented with 1 mM CaCl 2 and 1 mM MgCl 2 .

dimethyl-1-propyl-2H-indol-2-ylidene)ethyli- dene]-1-cyclohexen-1-yl]ethenyl]-3,3-dimethyl-1-propylindo- lium iodide)

Targeted and non-targeted LNP preparation

The formulation and composition of LNPs (previously called Lipimage815) have been already described in details (Jacquart et al., 2013). Briefly, the IR780-Oleyl dye was synthesized starting from the commercial IR780 iodide by adding a C 18 lipid chain. A mixture of oil, Suppocire NB TM and lecithin was prepared, then, the dye was introduced in it. An aqueous solution with a mix of Myrj TM S40 (PEG-40 Stearate) and SA COHN -PEG 100 -SPDP (3% of the totality of PEG) in 154 mM NaCl buffer was added to the oil phase. Fifty nmlarge lipid particles were formed through the emulsification process during a 5 min-sonication using a VCX750 Ultrasonic processor (power output 190 W, 3 mm probe diameter, Sonics). A solution of DTT was added to the LNPs and agitated for 2 h to obtain LNPs-SH. The LNP dispersions were dialyzed overnight at RT against 1000 times their volume in 154 mM NaCl buffer [START_REF] Eliasof | Correlating preclinical animal studies and human clinical trials of a multifunctional, polymeric nanoparticle[END_REF][START_REF] Phillips | Clinical translation of an ultrasmall inorganic optical-PET imaging nanoparticle probe[END_REF](14)ZelluTrans,Carl Roth,France).

Following the first step of LNP-SH synthesis, the peptide conjugation was carried out by the thiol-maleimide coupling reaction. 2 mL of LNP-SH (15% lipids) solution were mixed with 75 mL of each peptide linked with maleimide (cyclic RGD and cyclic RAD À 750 nmol in DMSO). After 4 h of a coupling reaction at RT in the dark, the remaining SH groups were neutralized using a solution of 1-(2-hydroxyethyl)-1H-pyrrole-2,5-dione for 30 min. Another dialysis was completed overnight to purify the peptide-conjugated LNPs. The final LNP dispersions were filtered through a 0.22 mm Millipore membrane for sterilization and stored at 4 C in the dark.

Characterization of LNPs

LNPs' hydrodynamic diameters, polydispersity indexes (PDI) and zeta potentials were determined in sterile 0.1X PBS (no CaCl 2 / MgCl 2 ) using a dynamic light scattering (DLS) instrument (Zeta Sizer NanoZS, Malvern, UK), performed in triplicate. The absorbance and fluorescence spectra were measured in 1X PBS using a spectrophotometer Thermo Scientific Evolution 201 and a fluorescence spectrometer Perkin Elmer LS55, respectively. The emission spectrum was measured with excitation at 790 nm. Each measurement was performed in triplicate. Concentrations mentioned for LNPs represent the concentrations of particles.

Cell lines and culture conditions

HEK293(b 3 ) cells, a human embryonic kidney cell line HEK293 that was stably transfected with the human b 3 integrin gene (kindly provided by J-F. Gourvest, Aventis, France), were cultured in DMEM-GlutaMAX TM supplemented with 10% FBS and 700 mg/mL Geneticin. HEK293(b 3 )-a v RFP cells, a variant cell line of HEK293 (b 3 ) transfected with the human a v integrin gene tagged by RFP at its extracellular domain (N-terminus part), were cultured as the HEK293(b 3 ) cells but in the presence of 1 mg/mL Puromycin (See Supplementary information). DU145 cells, a human prostate carcinoma cell line (purchased from ATCC), and M21 cells, a human melanoma cell line, were cultured in DMEM-GlutaMAX TM with 10% FBS. PC3 cells, a human prostate adenocarcinoma cell line (ATCC), were cultured in DMEM-F12-GlutaMAX TM with 10% FBS. Cells were maintained at 37 C, 5% CO 2 in a humidified incubator.

Evaluation of a v b 3 expression by flow cytometry

Cells were grown and harvested at 60% confluency. The cells were briefly washed with 1X PBS and resuspended with trypsin. After being centrifuged at 1200 rpm for 5 min, the cells were suspended in 1X PBS. 5 Â 10 5 cells per tube were suspended in 100 mL cold 1X PBS and mixed with 20 mL of anti-a v b 3 monoclonal antibody (PE-conjugated mouse anti-human CD51/CD61, clone 23C6, BD Pharmingen) during 1 h on ice in the dark. After rinsing twice they were finally suspended in 200 mL cold PBS. Their fluorescence intensity was measured using a flow cytometer (BD Accuri C6; BD Biosciences).

Cell attachment assay with vitronectin

A 96-well plate (Nunc Maxisorp ELISA plates) was coated with 5 mg/mL vitronectin (50 mL per well) overnight at 4 C. Then, the wells were drained and incubated with 3% (w/v) BSA (50 mL per well) for an hour at 37 C. The negative control wells were coated with 3% BSA only. The plate was rinsed once with HBSS. 50 mL of 25 000 cells and 50 mL of the LNP per well were added in the wells and incubated at 37 C for 15 min (HEK293(b 3 ) and DU145) or 30 min (PC3). When needed, the cells and LNPs were diluted in HBSS. The plates were drained and gently washed twice with 200 mL of 1X HBSS per well. The surface-bound cells were fixed with ethanol for 15 min at RT and dried at RT or 37 C. The fixed cells were then stained with 1% methylene blue (50 mL per well) for 15 min at RT, rinsed in running water and dried. The blue dye was finally eluted in 0.1 N HCl (50 mL per well) and the plates were scanned at a 620 nm absorbance wavelength using a Beckman Coulter AD340 plate reader. All the measurements were performed in triplicate (or quadruplicate) and repeated twice. and grown overnight in DMEM with 10% FBS and 700 mg/mL Geneticin at 37 C. The cells were washed once with pre-warmed 1X PBS and the chambered coverglass was moved to a microscope (Zeiss AxioObserver Z1) equipped with a humidified culture chamber. After being placed on the microscope, the cells were treated with the LNPs diluted to 45 or 4.5 nM in HBSS. Then, 3% (v/ v) FBS were shortly added to each well. The images were taken every 5 min starting from the treatment time point (T0).

Confocal laser scanning microscopy

HEK29(b 3 )-a v RFP cells were seeded (at 5 Â 10 4 cells/500 mL/ well) in a 4-well chambered coverglass (Nunc Lab-Tek, Thermo Fisher Scientific) and kept overnight in DMEM with 10% FBS, 700 mg/mL Geneticin and 1 mg/mL Puromycin at 37 C. The cells were washed once with pre-warmed PBS and treated with 500 mL of different LNPs diluted at 45 or 4.5 nM in HBSS. After an incubation time of 5, 10, 30 or 60 min, the cells were gently washed once with HBSS and fixed with 4% PFA containing 1 mg/mL of Hoechst 33342 for 15 min at RT in the dark. Then the cells were washed and stored in 1X PBS and observed using a confocal microscope LSM510 NLO META FLIM (Zeiss Axiovert 200M). A 63Â oil immersion objective was used. The Hoechst-stained nuclei were excited with a 720 nm biphoton laser and detected with a 390-465 nm filter. The RFPs were excited with a 543 nm laser and signals were collected with a 565-615 nm filter.

For the internalization study, the HEK293(b 3 )-a v RFP cells were seeded in the 4-well chamber coverglass and incubated overnight. The incubation, fixation, and Hoechst-staining were as described above. The cells were then observed using a confocal microscope LSM710 NLO (Zeiss AxioObserver Z1) with a 63x oil immersion objective. The LNP signal was excited using a 633 nm laser, set up at 50% of its maximum intensity and a 636-758 nm detection filter was used. The Hoechst was excited with a 405 nm laser and detected with a 415-502 nm filter. The RFP was excited with a 560 nm laser and detected with a 563-631 nm filter. The scanned images were pseudo-colored in blue for nuclei, in green for a v RFP and in red for the LNP.

Human tumor xenograft mice

All animal procedures are in compliance with the guidelines of the European Union (regulation n86/609), taken in the French law (decree 87/848) regulating animal experimentation. All efforts are made to minimize animal suffering and to reduce the number of animals used. All animal manipulations are performed with sterile techniques and are approved by the ethical committee of Grenoble for the use of animal research (France). Subcutaneous tumor mice bearing DU145 or M21 tumor were used for an in vivo evaluation of the LNPs. Eight weeks old male NMRI nude mice were purchased from Janvier, Le Genest-Saint-Isle, France. Mice were anesthetized (isoflurane/oxygen 3.5%-4% for induction and 1.5%-2% thereafter, CSP, Cournon, France) and were subcutaneously injected with 10 7 of DU145 cells or 4 Â 10 6 of M21 cells per mouse in the right trunk. When the tumors reached a volume between 200 and 500 mm 3 , the mice were used for LNP injections. Tumor growths took 70 days for the DU145 tumors and 43 days for the M21 tumors in average.

NIR in vivo imaging after injections of LNPs

Isoflurane-anesthetized mice were injected in the tail vein with 200 mL of 450 nM LNPs (equal to 10 nmol of fluorophores). NIR fluorescence images were acquired 1, 3, 5 and 24 h after injection with a Fluobeam 1 800 device from Fluoptics (Grenoble, France) equipped with a 780 nm excitation laser and an LP820 nm detection filter. The mice were sacrificed after the last acquisition and fluorescence present in the extracted organs and serum were also acquired. Major organs were cryo-conserved for ex vivo studies. Image analysis was performed using the Wasabi software (Hamamatsu, Massy, France).

Statistical analysis

Statistical analysis of the data was performed by using two-way ANOVA with Bonferroni post-tests. A difference of p value < 0.05 was considered statistically significant. All the data are shown as mean AE SD.

Results

Characteristics of LNPs

The outer layer composition and chemical structure of peptidecoated LNPs are illustrated in Fig. 1. LNPs are made of a phospholipid monolayer that delineates their oily core containing the fluorophores, and an outer shell of PEG surfactant with or without the peptide ligand for tumor targeting. The cyclic RGDfK (a v b 3 À targeting peptide) or cyclic RADfK (a negative control of cRGD in which the glycine was replaced by an alanine, Fig. 1-iv) peptides were linked on the PEGs by a maleimide-thiol coupling reaction (Fig. 1-iii).

The two peptide-conjugated LNPs showed a slightly larger diameter of 55 nm as compared to the 50 nm of the standard LNP without peptide. The polydispersity indexes (PDI) of the three LNPs were lower than 0.13, indicating that they are monodispersed and do not form aggregates. The zeta potentials were kept closed to neutrality around À4 mV for each LNPs without significant difference due to the presence of the peptides (Table 1). Such zeta-potentials are avoiding non-specific electrostatic interactions.

Due to the inclusion of IR780-Oleyl, the maximum absorbance was observed for wavelengths centered at 790-792 nm and the maximum emission band was detected at 810-812 nm (Supplementary Fig. 1). Based on the measured concentrations of IR780 and of LNPs, we calculated that each nanoparticle contained approximately 108 molecules of dye (Table 1).

CRGD-LNPs prevent cell binding on vitronectin

We investigated if the presence of cRGD could prevent the binding of three different cell lines on vitronectin, a "natural" protein ligand integrins on cells grown in a monolayer. Confluent HEK293(b 3 ) cells were treated with 4.5 nM or 45 nM cRGD, cRAD, or standard LNP in a solution containing 3% FBS. The cells were observed for more than 3 h on a microscope equipped with a culture chamber (Fig. 3). In the presence of 45 nM cRGD-LNP the cells started retracting during the first 5 min and gradually lost adhesion over time. When used at 4.5 nM, cRGD-LNP still affected cell adhesion, although less severely, and this effect was only transient and the cell-cell contacts were restored after 3 h. The cRAD and standard LNPs did not affect cell shape when tested at 45 and 4.5 nM LNP (Fig. 3 and Supplementary Fig. 2). As can be seen on the bottom panel, cell shrinkage, loss of focal and of cell-cell junctions occurred very rapidly and were visible as soon as 5 min after addition of the cRGD-LNPs (Fig. 4B). This indicated that cRGD-LNPs were functional and could bind to a v b 3 even when the integrin was already engaged.

Internalization of cRGD-LNP

Using the HEK293(b 3 )-a v RFP cells, we further investigated the internalization of cRGD-LNPs. Please note that the "red" signal of the a v RFP protein is presented in green. In contrast with the LNPs and cRAD-LNPs which did not provide a significant cell labeling, an intense labeling of the cell membranes was observed with cRGD-LNPs at 45 nM (Fig. 5). Similar results were also obtained when the different LNPs were used at 4.5 nM (Supplementary Fig. 3). Importantly, intracellular vesicles (probably endosome) were also positively stained by both the cRGD-LNP and by the integrin-fusion protein after 60 min of incubation (Fig. 5, first row), but not after 10 or 30 min (Supplementary Fig. 4), suggesting that the internalization was mediated by the a v b 3 integrin via a slow process that necessitated between 30 and 60 min. The vesicles co-stained by cRGD-LNPs/integrin a v b 3 were also found when the cells were washed 15 min after addition of the cRGD-LNPs (Fig. 5, second row). In these conditions, the plasma membranes were not strongly labeled: the integrin/cRGD-LNP complexes initially located to the membrane were internalized, while the cell medium containing the cRGD-LNP was removed, preventing any binding after the rinsing.

In vivo evaluation

In order to demonstrate the cRGD-specific targeting of tumors in vivo, we compared the biodistribution of the three LNPs in mice bearing DU145 or M21 tumors. M21 is a human melanoma cell line that strongly expresses the a v b 3 integrin ((Felding-Habermann et al., 1992), Supplementary Fig. 5). After i.v. injections of 200 mL of 450 nM LNPs, we followed the distribution of the NIR-labeled LNPs in real time using a NIR camera. Unexpectedly, no major differences in the tumor accumulation were found between the cRGD-, cRAD-and standard LNPs in both DU145 or M21 tumors (Fig. 6 and Supplementary Fig. 6). This was confirmed ex-vivo after sacrifice of the mice and excision of the tumors (Fig. 7 and Supplementary Fig. 7). As well, no significant differences of EPRmediated accumulation were detected when the mice bearing M21 tumors were followed for 4-5 days after the injections.

Thus, the 3 types LNPs accumulate passively via the EPR effect in the 2 tumor types. However, the presence of cRGD does not provide a significant augmentation of the number of LNPs that can be detected macroscopically in the tumors.

Discussion

The main objective of this study was to demonstrate a specific binding and internalization of the cRGD-LNP in order to generate tumor-targeted nanoparticles for theranostic purpose.

We found that cRGD-LNP can bind and interfere functionally with integrin a v b 3 in vitro. This indicated that the chemical coupling of cRGD on the PEGylated lipid nanocapsules was correct and resulted in an efficient recognition of the integrin receptors on the cell surface. cRGD-LNPs were able to bind to "free" integrins present on the surface of the cells in suspension as well as on those already bound to vitronectin. This later effect caused an efficient shrinking of adherent cells due to the inhibition of cell-cell and cell-matrix contacts. This effect was strong enough to detach integrin positive cells when cRGD-LNPs were used at a 45 nM concentration. At a lower concentration (4.5 nM), the detachment was transient and cell adhesion and spreading were recovered after 3 h. As expected this effect was more pronounced on DU145 cells expressing low levels of integrins than on HEK293(b 3 ) cells which are strongly a v b 3 positive. Of note, the HEK293(b 3 ) cell line is integrin b 5 negative while DU145 and PC3 are known to express low to moderate levels of Integrin internalization can occur via macro-pinocytosis, clathrin-dependent or clathrin-independent endocytosis [START_REF] Bridgewater | Integrin trafficking at a glance[END_REF]. We previously demonstrated that internalization of a tetravalent cRGD presenting scaffold is mediated via clathrin-mediated endocytosis (Sancey et al., 2009). Other integrin-targeted NPs such as trimethyl chitosan-based NPs labeled with FQSIYPpIK (an a v b 3 targeting peptide) are using clathrin-mediated endocytosis [START_REF] Xu | The transport mechanism of integrin alphavbeta3 receptor targeting nanoparticles in Caco-2 cells[END_REF]. In the present study, we also observed that only the cRGD-LNPs are actively endocytosed and can be seen trafficking within a v -RFP and IR780-Oleyl co-labeled vesicles. In addition, a weaker and diffuse IR780 fluorescent signal was also building up over time in the cytoplasm and this diffuse signal was also detected in the presence of the negative controls cRAD-and non-targeted-LNPs. This may indicate that while only cRGD-LNPs can enter specifically via a cRGDintegrin-dependent pathway(s), all 3 types LNPs can fuse nonspecifically with the lipids of the cell membranes and can then release their content directly to the cytoplasm. This could have an important impact for LNP-mediated drug delivery. In particular, we are currently investigating the possible cytotoxic effect generated under an 800 nm light-activation of the IR780 dye since this dye is known to generate toxic reactive oxygen species usable for photodynamic therapy.

In vivo, no specific accumulation was noticed in the integrin positive DU145 or M21 subcutaneous tumors after intravenous injection of the cRGD-LNPs versus cRAD-or Standard-LNPs. Similar fluorescence intensities as well as similar kinetics of accumulation in the tumors were measured with the 3 types of LNPs. This suggested that the functionalization of the LNPs by cRGD motifs did not generate a significantly augmented capture or retention of the targeted versus not-targeted LNPs in the tumor mass. This is in agreement with our former work in which we used cRGD-targeted DID-loaded nanoemulsions. In this previous work, we had a very modest augmentation of cRGD-dependent targeting in vivo when using HEK293(b 3 ) tumors but not TSa/pc's (Goutayer et al., 2010). Actually, HEK293(b 3 ) cells are very strongly expressing the integrin while TSa/pc are weakly expressing it. On the other hand, we classified HEK293(b 3 ) subcutaneous tumors as "EPR-deficient" because an intravenous injection of lipid nanoemulsions generated a Tumor/Skin (T/S) ratio of 1.5, while TSa/Pc were found EPR-positive with a T/S ratio of 3.0 (Karageorgis et al., 2016). According to this scheme, DU145 and M21 tumors can thus be classified as EPR-positive based on their T/ S ratio of AE3, as well as integrin positive. This suggests that active, cRGD-mediated, targeting is not augmenting the accumulation of the nanoemulsions when the EPR effect is present. However, several groups including ours (Dufort et al., 2011;Karageorgis et al., 2017) have shown that RGD-vectorization is efficiently inducing a therapeutic effect (for very recent examples, see: [START_REF] Fang | Targeted glioma chemotherapy by cyclic RGD peptide-functionalized reversibly core-crosslinked multifunctional poly(ethylene glycol)-b-poly(epsiloncaprolactone) micelles[END_REF]Zhu et al., 2016), and a recent review by Zhong et al. (Zhong et al., 2014). This suggests that cRGD-active targeting may be important at the very last step, i.e. to deliver the cargo more efficiently into the cytoplasm of integrin positive cells. We are now actively trying to validate this hypothesis that may be important for therapeutic applications in particular when the toxic agent cannot cross the plasma membrane on its own. and internalization in vitro were consistent with our previous work in 2012 (Goutayer et al., 2010).

Tumor accumulation and biodistribution of LNPs in different tumors

Evaluation of SiNP binding to cells by flow cytometry

Evaluation of SiNP binding to cells by confocal microscopy

In vivo study with cRGD/ATWLPPR-PEG-SiNPs

However, the in vivo results are diverging.

The improved tumor accumulation by cRGD-targeting in vivo was apparent in 2012 using cRGD-lipidots (DiD loaded, 35 nm; cRGD peptides linked on DSPE-PEGMW5000 residues).

However, in the current study using cRGD-LipImage (IR780-oleyl dye loaded, 50 nm, cRGD conjugates on PEG-100 stearateMW5000), we mainly observed the EPR-derived accumulation in vivo.

Unfortunately, we were not able to verify the cRGD-targeting at a microscopic level using ex vivo samples as conventional microscopes equipped in our institute were not optimized for IR780 NIR dye imaging.

There are several factors that are probably related to our recent targeting results. First, the size of the NP impacts strongly on the EPR-mediated accumulation and the NP's penetration within tumors. Using the iRGD peptide, it has been shown that its high capacity of tumor penetration was coming from its mechanism of internalization similar to that of nutrient uptake: a cell-to-cell transfer mechanism helped this peptide to diffuse deeper in the tumor core [START_REF] Pang | An endocytosis pathway initiated through neuropilin-1 and regulated by nutrient availability[END_REF].

But, when they tried labeling micelle NPs with the iRGD peptide, the penetration turned out to be lower than that of the free iRGD peptide (see Figure 13B) [START_REF] Sugahara | Tissue-penetrating delivery of compounds and nanoparticles into tumors[END_REF]. Actually, the poor vascularisation, the predominant presence of non-functional blood vessels, and high interstitial pressure in solid tumor mass [START_REF] Chang | Mosaic blood vessels in tumors: frequency of cancer cells in contact with flowing blood[END_REF] are likely to prevent the diffusion of large NPs. Accordingly, it has been shown that NPs' diffusion limit was a function of NP's size.

This was established using 12, 60, and 125 nm NPs without targeting agent. Researchers found that only the 12 nm-large NPs were able to penetrate into the tumor's core while the two others stayed in the peri-vascular regions (Figure below) [START_REF] Popovic | A nanoparticle size series for in vivo fluorescence imaging[END_REF]. We are not able to define a precise cut-off of size where the NPs significantly lose their diffusion capacity, but, we can assume that this cut-off should be somewhere between 10 and 60 nm depending on the nature of the NP (surface coating, charge, deformability, etc.). Based on our work with 25 and 50 nm LNPs (Hirsjarvi et al., 2013), it was shown that the 25 nm LNP had more accumulation in tumors than the 50 nm-large one, thus smaller LNPs might be the best candidates for RGD targeting but a systematic comparison will be needed to verify this hypothesis.

The effect of the protein corona (following protein opsonization ) is another important consideration in vivo [START_REF] Monopoli | Biomolecular coronas provide the biological identity of nanosized materials[END_REF]. Once the NPs are injected in the vascular system, they are immediately exposed to a vast number of proteins, which activate the protein corona formation on the NPs. It has been shown that the protein corona can form very quickly (in half a minute) interacting with approximately 400 different proteins, and also that the protein corona has a dynamic exchange of proteins with the plasma [START_REF] Tenzer | Rapid formation of plasma protein corona critically affects nanoparticle pathophysiology[END_REF]. Bakewell, S.J., Nestor, P., Prasad, S., Tomasson, M.H., Dowland, N., Mehrotra, M., Scarborough, R., Kanter, J., Abe, K., Phillips, D., et al. (2003). Platelet and osteoclast beta3 integrins are critical for bone metastasis. Proceedings of the National Academy of Sciences of the United States of America 100, 14205-14210. Banerjee, S.K., Zoubine, M.N., Tran, T.M., Weston, A.P., and Campbell, D.R. (2000). Overexpression of vascular endothelial growth factor164 and its co-receptor neuropilin-1 in estrogen-induced rat pituitary tumors and GH3 rat pituitary tumor cells. International journal of oncology 16, 253-260. Barabino, G., Porcheron, J., Cottier, M., Cuilleron, M., Coutard, J.G., Berger, M., Molliex, S., Beauchesne, B., Phelip, J.M., Grichine, A., et al. (2016). Improving Surgical Resection of Metastatic Liver Tumors With Near-Infrared Optical-Guided Fluorescence Imaging. Surg Innov 23, 354-359. Bauer, K., Mierke, C., and Behrens, J. (2007). Expression profiling reveals genes associated with transendothelial migration of tumor cells: a functional role for alphavbeta3 integrin. International journal of cancer Journal international du cancer 121, 1910-1918. Benezra, M., Phillips, E., Overholtzer, M., Zanzonico, P.B., Tuominen, E., Wiesner, U., and Bradbury, M.S. (2015). Ultrasmall integrin-targeted silica nanoparticles modulate signaling events and cellular processes in a concentrationdependent manner. Small (Weinheim an der Bergstrasse, Germany) 11, 1721-1732.

Berge, M., Bonnin, P., Sulpice, E., Vilar, J., Allanic, D., Silvestre, J.S., Levy, B.I., Tucker, G.C., Tobelem, G., and Merkulova-Rainon, T. (2010). Small interfering RNAs induce target-independent inhibition of tumor growth and vasculature remodeling in a mouse model of hepatocellular carcinoma. The American journal of pathology 177, 3192-3201.

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Introduction

Delivery of drugs or contrast agents to tissues, cells or intracellular molecular targets is a major challenge, particularly for poorly soluble drug candidates. In all cases, it is important to obtain 1) a favorable tumor/normal-tissue ratio of accumulation of the delivered molecules and 2) the lowest non-specific accumulation to reduce toxicity and contaminating signal.

Targeted molecules can be schematically grouped into two categories: small molecules that diffuse freely until they reach their target and large macromolecules that exhibit difficulty in reaching the target but that carry a large cargo and are often multifunctional.

Because small targeted molecules enter easily into cells, their biological function and toxicity will rely on their specific activity and on rapid wash-out. Accordingly, during the last decade, targeted therapies such as epidermal growth factor receptor tyrosine kinase inhibitors (EGFR-TKIs) represent a crucial innovation that has significantly altered the quality of life and overall survival rate of cancer patients. Unfortunately, they still present some side effects and frequently result in resistant tumors and relapse.

In contrast, several efficient small molecules do not function properly after systemic administration because of their poor solubility, fast metabolization and/or rapid clearance from the bloodstream (1) or because of efficient detoxification "pumps" on tumor cells, such as P-glycoprotein, conferring a multidrug resistance (MDR) phenotype. Although the use of nanocarriers constitutes an interesting alternative for alleviating these problems, nanovectors tend to diffuse very poorly within tissues. Fortunately, nanoparticles (NPs) can accumulate passively in tumors due to the so-called "enhanced permeability and retention" (EPR) effect initially described by Yasuhiro Matsumura and Hiroshi Maeda in 1986 (2, 3), which allows macromolecules with a MW >40 kDa (>5 nm) to accumulate in tumors. To understand how the EPR effect operates, we need to separate the parameters linked to the patient from those linked to the NP itself.

The enhanced permeability and retention effect: biological point of view

The EPR effect is due primarily to an elevated systolic blood pressure that pushes blood into tumor tissue but also to tumor angiogenesis, which dictates the blood vessel's density and properties. This pathologic angiogenesis produces a tumor endothelium lacking pericytes, presenting a large fenestration (±100 nm) and thus leaky blood vessels. In addition, because lymphangiogenesis is not very active in tumors, lymphatic drainage is inefficient. Although the impact of the EPR effect is unquestionable, albeit variable in mice, its importance or existence in humans remains under discussion. Indeed, only human sarcoma tumors present tumor accumulation of liposomes, whereas breast, lung, ovarian, head and neck, and glioblastoma tumors showed lower accumulation of liposomes than in the surrounding tissues (4). Tumors grow rapidly in a confined environment (and this is particularly true in animal models as compared to human tumors), inducing pressure on the surrounding stroma and extracellular matrix that leads to an increase in elastic stress and to a 10-100-fold increase in interstitial fluid pressure (IFP) in the tumor as compared to healthy surrounding tissues.

Accordingly, the convection flow, which can be represented by the difference in pressure between the therapeutic solution and tumor, will push the fluid (and cancer cells) away from the tumor and into the surrounding tissue, reducing the transport of molecules away from vessels [START_REF] Torosean | Nanoparticle uptake in tumors is mediated by the interplay of vascular and collagen density with interstitial pressure[END_REF]. It should be noted that the transport of small molecules is driven by diffusion rather than convection. Since diffusion is a function of the difference in concentration between the therapeutic solution and tumor rather than pressure, the diffusion of small drugs is much less affected by flow, and these molecules can often diffuse toward areas of increased fluid flow ( 6) while NPs cannot.

The enhanced permeability and retention effect: biophysical perspective

The EPR effect also depends on the physicochemical properties of nanovectors (e.g., size, electric charge) that will allow NPs to move relatively easily. The first rule is that the NP must remain in the blood circulation for a long time (>6 h). This is the reason why chemists have made great efforts to adapt their size and to create a surface as invisible as possible to avoid capture by the reticulo-endothelial system (RES); additionally, a large diameter will prevent rapid elimination through the kidneys. This stealth quality also prevents electrostatic interaction with extravascular matrices once in the tumor microenvironment but will also reduce the capacity of the NPs to interact with the target cell; the NPs will thus remain entrapped in the interstitium until they are degraded or until they are endocytosed.

It is thus tempting to introduce a targeting molecule onto nanovectors at this step because the ligand could interact with the target receptor and allow intracellular delivery.

The targeting dilemma

There has long been an important debate concerning the use of active versus passive targeting of nanovectors (4,[START_REF] Bazak | Cancer active targeting by nanoparticles: a comprehensive review of literature[END_REF][START_REF] Poh | Comparison of nanoparticle penetration into solid tumors and sites of inflammation: studies using targeted and nontargeted liposomes[END_REF]. Although most studies present very positive and clear results in vitro, the final gain achieved in vivo is modest [START_REF] Goutayer | Tumor targeting of functionalized lipid nanoparticles: assessment by in vivo fluorescence imaging[END_REF]. It should be noted that the presence of a targeting ligand is usually not expected to modify the global biodistribution of an NP and its overall accumulation in the tumor because this is mainly EPR dependent. In fact, the presence of a targeting ligand can sometimes reduce NP bioavailability because it can augment the hydrodynamic radius, reactivity toward the microenvironment, and opsonization (capacity to bind blood proteins) and thus reduce the circulation time in blood.

There is a consensus that receptor-mediated internalization can improve therapeutic efficacy once an NP is inside the tumor as a result of the EPR effect. This is particularly true for cancer drugs such DNA and siRNA molecules that cannot be internalized alone by cells and that are active only when delivered intact to the appropriate intracellular compartment [START_REF] Choi | Mechanism of active targeting in solid tumors with transferrin-containing gold nanoparticles[END_REF][START_REF] Huang | A reexamination of active and passive tumor targeting by using rod-shaped gold nanocrystals and covalently conjugated peptide ligands[END_REF] but also for drugs such as doxorubicin that are sensitive to the MDR phenotype.

The important additional cost in the production of targeted nanovectors versus passive, EPRdriven, NPs and versus standard treatments should proportional to the benefits in order to be acceptable. Because it is almost impossible to perform a side-by-side comparison in humans, preclinical data are extremely important because they are proven to be reliable, as established for CRLX101 [START_REF] Eliasof | Correlating preclinical animal studies and human clinical trials of a multifunctional, polymeric nanoparticle[END_REF] or cRGDY-PEG-C dot particles [START_REF] Phillips | Clinical translation of an ultrasmall inorganic optical-PET imaging nanoparticle probe[END_REF]. Nonetheless, these data should be sustained by a precise knowledge of the biomarkers of the pathology being addressed in patients (i.e., the level and accessibility of the targeted receptors) and by a deep understanding of the patient's pathological characteristics that will influence access of the nanovectors to its target [START_REF] Maeda | Toward a full understanding of the EPR effect in primary and metastatic tumors as well as issues related to its heterogeneity[END_REF]. In a recent review, Hiroshi Maeda concludes: "given the highly heterogeneous and continuously evolving nature of the tumor microenvironment, the optimal design of NP is likely to be disease specific. This is a formidable task, especially considering the difference from one tumor to the next, from primary tumor to its metastasis, from one day to the next in the same tumor and the changes induced by treatment" ( 14).

Physico--chemical factors that influence NP properties and their biodistribution

As shown in Figure 1, in addition to the presence of ligand, several parameters influence the properties of an NP, for instance, size, charge, shape and hydrophobicity. Active targeting is achieved by the addition of ligands that will provide recognition of the target; at the same time, the targeted NP must avoid all non-specific interactions. This requires that its surface is correctly masked and that only the ligand is accessible. This is not an easy task because it relies on a complex chemistry that must address many parameters (e.g., NP composition and architecture, presence of a large number of chemically reactive species, necessity of mastering the geometry and orientation of the linked compounds, quality of the coverage and density, presentation of the ligand) and because it is very often impossible to completely characterize the final products. All these problems need to be addressed very early in the design process, and all chemical steps should be perfectly mastered and adequately selected to comply with the legal recommendations on the definition of a nanomaterial. This is a major bottleneck and, in general, the targeted NP should be very sophisticated…but as simple as possible to keep the development of such nanomedicine "reasonably simple, inexpensive, and scalable" [START_REF] Zamboni | Best practices in cancer nanotechnology: perspective from NCI nanotechnology alliance[END_REF].

Influence of the size of the NP

Because the fenestration of the tumor-specific endothelium ranges between 100 and 200 nm, the NP must be smaller than that cut-off. Small macromolecules (<6 nm) accumulate more rapidly and diffuse deeper in the a tumor than larger molecules; however, this is very transitory [START_REF] Dreher | Tumor vascular permeability, accumulation, and penetration of macromolecular drug carriers[END_REF], and NPs are rapidly filtered through the kidneys [START_REF] Choi | Design considerations for tumour-targeted nanoparticles[END_REF]. This situation can be acceptable for the delivery of a contrast agent,, but less for drug delivery. Within 24 hours, the majority of small molecule imaging probes are usually filtered by the kidneys with a clearance efficiency greater than 50% of the injected dose (%ID) and a poor non-specific accumulation in the main organs. This is the case for most of the hydrophilic intravenously injected and FDA-approved contrast agents used in humans for MRI, CT, SPECT or PET imaging as well as for the near-infrared dyes used for optical imaging in preclinical studies.

NPs of very small diameter (<6 nm), such as gold nanoclusters [START_REF] Zhang | Enhanced tumor accumulation of sub-2 nm gold nanoclusters for cancer radiation therapy[END_REF][START_REF] Zhang | Ultrasmall Au(10-12)(SG)(10-12) nanomolecules for high tumor specificity and cancer radiotherapy[END_REF] or gold NPs [START_REF] Liu | PEGylation and zwitterionization: pros and cons in the renal clearance and tumor targeting of near-IRemitting gold nanoparticles[END_REF], Quantum Dots [START_REF] Choi | Design considerations for tumour-targeted nanoparticles[END_REF], silica [START_REF] Burns | Fluorescent silica nanoparticles with efficient urinary excretion for nanomedicine[END_REF] or gadolinium-polysiloxane-containing ultrasmall NPs [START_REF] Bianchi | Quantitative biodistribution and pharmacokinetics of multimodal gadolinium-based nanoparticles for lungs using ultrashort TE MRI[END_REF][START_REF] Lux | Ultrasmall Rigid Particles as Multimodal Probes for Medical Applications[END_REF], are also very rapidly cleared via the kidneys; however, despite their efficient renal clearance, ultrasmall inorganic NPs may present a positive EPR effect and function in theranostic applications (this will be discussed later). A recent and excellent review presents a broad view of our understanding of the behavior of inorganic NPs [START_REF] Yu | Clearance Pathways and Tumor Targeting of Imaging Nanoparticles[END_REF].

We investigated the EPR effect attained with NPs ranging from 25 to 100 nm. Although the smallest nanoemulsions (25 nm) disappeared from the blood circulation faster than the larger ones (50 and 100 nm) due to a rapid elimination and wider tissue distribution (25), the biodistribution profiles of all these particles were similar after 24 hours and a positive but highly variable EPR effect was generally detected in different tumor types [START_REF] Hirsjarvi | Influence of size, surface coating and fine chemical composition on the in vitro reactivity and in vivo biodistribution of lipid nanocapsules versus lipid nanoemulsions in cancer models[END_REF] (Figure 2). This was also observed using polymeric micelles with diameters of 30, 50, 70 and 100 nm (27): all the micelles penetrated highly permeable tumors in mice, but only the smallest was able to penetrate poorly permeable tumors.

Clearly, the size and surface area of an NP also impacts the number of ligands that can be attached (Figure 3) and thus NP reactivity and biological function. In addition, the surface elasticity [START_REF] Hirsjarvi | Surface modification of lipid nanocapsules with polysaccharides: From physicochemical characteristics to in vivo aspects[END_REF] or the aspect ratio (29) may play a very important role. Indeed, spherical NPs will not diffuse like rod-shaped or filamentous nano-objects, and their interaction with the microenvironment and on the cell surface will differ, providing variable active targeting properties (29).

Surface modification and opsonization

As soon as they are diluted in serum-containing medium, NPs are "opsonized" [START_REF] Liu | Recognition and clearance of liposomes containing phosphatidylserine are mediated by serum opsonin[END_REF]. This is due to the adsorption of abundant proteins such as antibodies or proteins from the complement system. These "tags" are detected by the host's immune system as "foreign" and augment the phagocytosis and rapid elimination of opsonized NPs by RES cells and particularly macrophages present in the liver (Kupffer cells) and in the spleen. This results in an unfavorable short circulation half-life and thus to a reduced EPR effect. Furthermore, opsonization also affects presentation of the ligands on the NP surface and usually reduces its accessibility and its targeting potential [START_REF] Hadjidemetriou | In Vivo Biomolecule Corona around Blood-Circulating, Clinically-Used and Antibody-Targeted Lipid Bilayer Nanoscale Vesicles[END_REF].

Opsonization can be prevented by grafting the hydrophilic polymer polyethylene glycol (PEG), or less frequently by using polysaccharides (e.g., dextran, chitosan), onto the NP surface [START_REF] Hirsjarvi | Surface modification of lipid nanocapsules with polysaccharides: From physicochemical characteristics to in vivo aspects[END_REF]. The presence of PEG polymers is efficient in reducing opsonization and generating an augmented half-life in the circulation (32,[START_REF] Allen | Liposomes containing synthetic lipid derivatives of poly(ethylene glycol) show prolonged circulation half-lives in vivo[END_REF], ultimately leading to augmented therapeutic activity [START_REF] Papahadjopoulos | Sterically stabilized liposomes: improvements in pharmacokinetics and antitumor therapeutic efficacy[END_REF]. However, several publications also describe that the PEG coating can generate an immune response, complement activation [START_REF] Hamad | Distinct polymer architecture mediates switching of complement activation pathways at the nanosphere-serum interface: implications for stealth nanoparticle engineering[END_REF][START_REF] Hamad | Poly(ethylene glycol)s generate complement activation products in human serum through increased alternative pathway turnover and a MASP-2-dependent process[END_REF] and possible appearance of anti-PEG immunoglobulins [START_REF] Schellekens | The immunogenicity of polyethylene glycol: facts and fiction[END_REF]. These immunoglobulin could be responsible for their accelerated blood clearance (ABC), although a recent study indicated that the ABC phenomenon is not due to the opsonization of NPs with anti-PEG IgM [START_REF] Hashimoto | Relationship between the concentration of anti-polyethylene glycol (PEG) immunoglobulin M (IgM) and the intensity of the accelerated blood clearance (ABC) phenomenon against PEGylated liposomes in mice[END_REF][START_REF] Mima | Anti-PEG IgM Is a Major Contributor to the Accelerated Blood Clearance of Polyethylene Glycol-Conjugated Protein[END_REF].

A recent study also investigated the quality of the protein corona formed on intravenously injected liposomes and compared it to the situation in vitro [START_REF] Hadjidemetriou | In Vivo Biomolecule Corona around Blood-Circulating, Clinically-Used and Antibody-Targeted Lipid Bilayer Nanoscale Vesicles[END_REF]. This was achieved using bare (non PEGylated), PEGylated or antibody-targeted liposomes. The authors observed that the variety of proteins in the in vivo corona was wider than that in vitro and that the presence of these proteins significantly reduced the activity of the targeting antibody, confirming previous observations [START_REF] Salvati | Transferrinfunctionalized nanoparticles lose their targeting capabilities when a biomolecule corona adsorbs on the surface[END_REF].

In previous work, we also demonstrated that the nature (length) and net electric charge of PEG polymers linked to the surface of a 50 nm large silica NP strongly impacted biodistribution after intravenous injection in mice [START_REF] Faure | Control of the in vivo biodistribution of hybrid nanoparticles with different poly(ethylene glycol) coatings[END_REF], underlining the importance of the net electric charge on the surface [START_REF] Dufort | Physico-chemical parameters that govern nanoparticles fate also dictate rules for their molecular evolution[END_REF].

Electric charge

The presence of positive or negative electric charges on the surface of an NP can enhance its reactivity and may have two opposite effects: 1) increasing its affinity for compounds or cells presenting the opposite charge [START_REF] Wang | Folate-PEG coated cationic modified chitosan--cholesterol liposomes for tumortargeted drug delivery[END_REF] and 2) reducing its mobility within extracellular matrices.

Neutral particles (zwitterionic) will lie in between these two situations.

In mice, the presence of positive charges is usually associated with a stronger tumor accumulation. Using mesoporous silica NPs, which present a net negative electric charge after PEGylation or a positive one after the addition of polyethyleneimine, (PEI), Meng et al. demonstrated that the presence of the positive charge improved EPR-mediated accumulation in tumors [START_REF] Meng | Use of size and a copolymer design feature to improve the biodistribution and the enhanced permeability and retention effect of Doxorubicin-loaded mesoporous silica nanoparticles in a murine xenograft tumor model[END_REF]. Such an effect could be related to the fact that positive charges favor binding to the tumor vasculature, but with a very poor diffusion within the tumor mass. This is now being exploited in clinical trials using EndoTAG TM , a cationic liposome for the delivery of embedded paclitaxel for the treatment of pancreatic cancer [START_REF] Lohr | Cationic liposomal paclitaxel plus gemcitabine or gemcitabine alone in patients with advanced pancreatic cancer: a randomized controlled phase II trial[END_REF], advanced cancers and liver metastasis [START_REF] Fasol | Vascular and pharmacokinetic effects of EndoTAG-1 in patients with advanced cancer and liver metastasis[END_REF] and recently in triple-negative breast cancer [START_REF] Awada | A randomized controlled phase II trial of a novel composition of paclitaxel embedded into neutral and cationic lipids targeting tumor endothelial cells in advanced triple-negative breast cancer (TNBC)[END_REF] and head and neck cancer [START_REF] Strieth | Phase I/II clinical study on safety and antivascular effects of paclitaxel encapsulated in cationic liposomes for targeted therapy in advanced head and neck cancer[END_REF]. Thus far, the results indicate that the use of these liposomes is safe, although their therapeutic impact was not obvious in phase I/II trails.

We demonstrated that the presence of negative charges due to the grafting of PEG-COOH onto silica-based 50 nm large NPs prevented their elimination by the RES [START_REF] Faure | Control of the in vivo biodistribution of hybrid nanoparticles with different poly(ethylene glycol) coatings[END_REF]. Moreover, the use of shorter PEG250-COOH instead of PEG2000-COOH also allowed RES evasion yet dramatically augmented elimination through the kidneys; this reduced half-life in the blood is associated with a lower EPR accumulation.

A very good example of a negatively charged NP largely used in the clinic is Abraxane ® , a 130 nm particle composed of human serum albumin non-specifically bound to paclitaxel [START_REF] Miele | Albumin-bound formulation of paclitaxel (Abraxane ABI-007) in the treatment of breast cancer[END_REF].

Note that although it has been described that its uptake could be mediated by albumin receptors such as gp60 and possibly by the Secreted Protein Acidic and Rich in Cysteine (SPARC), Abraxane ® is usually not considered an actively targeted NP [START_REF] Hoang | Docetaxel-carboxymethylcellulose nanoparticles target cells via a SPARC and albumin dependent mechanism[END_REF][START_REF] Yardley | nab-Paclitaxel mechanisms of action and delivery[END_REF]. The role of SPARC still remains an open debate [START_REF] Schneeweiss | Efficacy of nab-paclitaxel does not seem to be associated with SPARC expression in metastatic breast cancer[END_REF].

Density of ligands

The strength of the non-covalent binding of a ligand to its receptor is defined as its affinity. This affinity need not be augmented when the ligand is presented by an NP; in fact, it is very often decreased because of eventual chemical modifications [START_REF] Hennig | Nanoparticle multivalency counterbalances the ligand affinity loss upon PEGylation[END_REF], steric hindrance and possible inadequate presentation. It was demonstrated that PEGylation of EXP3174, a nonpeptidic ligand that binds the Angiotensin receptor with nM affinity, with PEG 5000 reduced its affinity by 580 times. However, if the ligand density on the surface is appropriate, the presentation of multiple copies of the ligand can generate a very strong cooperative effect, which represents its avidity. This cooperative effect derives from the fact that once the first ligand is bound, it increases the chances for others to find their target in the microenvironment. Indeed, when 11 copies of PEGylated EX3174 were covalently attached to 17 x 7 nm rod-shape quantum dots, the cooperative effect was capable of completely reversing the loss of affinity owing to the PEGylation [START_REF] Hennig | Nanoparticle multivalency counterbalances the ligand affinity loss upon PEGylation[END_REF]. It was also demonstrated that avidity can improve the efficiency of low-affinity ligands [START_REF] Weissleder | Cellspecific targeting of nanoparticles by multivalent attachment of small molecules[END_REF]. Multivalency also induces receptor clustering, which enhances the active internalization of nanosystems [START_REF] Sancey | Clustering and Internalization of Integrin alphavbeta3 With a Tetrameric RGD-synthetic Peptide[END_REF][START_REF] Stefanick | A systematic analysis of peptide linker length and liposomal polyethylene glycol coating on cellular uptake of peptide-targeted liposomes[END_REF]; the ligand density should thus be finely tuned [START_REF] Stefanick | A systematic analysis of peptide linker length and liposomal polyethylene glycol coating on cellular uptake of peptide-targeted liposomes[END_REF][START_REF] Hak | The effect of nanoparticle polyethylene glycol surface density on ligand-directed tumor targeting studied in vivo by dual modality imaging[END_REF] because when the density is too low or too elevated, the binding efficiency of the targeted NP is not optimal [START_REF] Elias | Effect of ligand density, receptor density, and nanoparticle size on cell targeting[END_REF]. A rough calculation and a schematic representation of the number of molecules that can be mathematically attached to a given surface area are presented in Figure 3; the actual definition of the most efficient number of ligands per NP should be carefully investigated (4).

Targeted NPs

Only a restricted number of actively targeted NPs are used in clinical trials (4). However, as observed in the non-exhaustive list of common targets and targeting ligands recovered from a rapid search in PubMed (see Table 1), numerous publications describe the potential of this approach in vitro and in preclinical models. I will not present an extensive analysis of each system, as excellent recent reviews can be found (4,[START_REF] Zamboni | Best practices in cancer nanotechnology: perspective from NCI nanotechnology alliance[END_REF][START_REF] Zhong | Ligand-directed active tumortargeting polymeric nanoparticles for cancer chemotherapy[END_REF], but I will rather attempt to present some selected works that describe the importance and complexity of using nanomedicines with targeting ligands of increasing molecular weight (small molecules, peptides and proteins).

Results of PubMed queries

Choice of target receptor

The most frequently chosen receptors are folate receptor (FR), integrins (and in particular α v ß 3 integrin), EGFR receptor, HER2, urokinase receptor (uPAR), transferrin receptor, CD44

and prostate specific membrane antigen (PSMA) because they are a) cell-surface receptors and b) usually overexpressed on tumors.

Targeting folate receptor using folic acid as an example of a small ligand Folic acid (FA, folate, also called vitamin B9) is an essential nutrient required for the biosynthesis of nucleotides and for cell proliferation. Two membrane-bound isoforms of FR have been identified in humans: α and ß. FRα (FOLR1) has a dissociation constant (Kd) for FA of 0.1 nM and is overexpressed on the surface of several cancers, including breast, kidney, lung, brain, and in particular in >90% ovarian cancers [START_REF] Parker | Folate receptor expression in carcinomas and normal tissues determined by a quantitative radioligand binding assay[END_REF]. FOLR1 shows limited expression in normal tissues; it is normally present only on the apical surface of epithelial cells [START_REF] Elnakat | Distribution, functionality and gene regulation of folate receptor isoforms: implications in targeted therapy[END_REF] and is thus poorly accessible from the bloodstream. FR + tumor cells expressing >10 6 receptors present receptors on the basal surface, and activated macrophages are known to present between 100,000 to 200,000 receptors ( 8). An anti-FOLR1 monoclonal antibody (farletuzumab) has entered clinical phases for the treatment of ovarian [START_REF] Sasaki | A phase I study of farletuzumab, a humanized anti-folate receptor alpha monoclonal antibody, in patients with solid tumors[END_REF][START_REF] Konner | Farletuzumab, a humanized monoclonal antibody against folate receptor alpha, in epithelial ovarian cancer: a phase I study[END_REF] and lung cancers [START_REF] Bronte | Farletuzumab for NSCLC: exploiting a wellknown metabolic pathway for a new therapeutic strategy[END_REF]. Folic acid has been developed as a drug delivery system under the name of vintafolide (EC245) in which it is associated to a vinca alkaloid via a short hydrophilic peptide spacer and a disulfide-containing self-immolating linker [START_REF] Vergote | Vintafolide: a novel targeted therapy for the treatment of folate receptor expressing tumors[END_REF][START_REF] Leamon | Folate-vinca alkaloid conjugates for cancer therapy: a structure-activity relationship[END_REF][START_REF] Li | Clinical pharmacokinetics and exposure-toxicity relationship of a folate-Vinca alkaloid conjugate EC145 in cancer patients[END_REF][START_REF] Reddy | Preclinical evaluation of EC145, a folate-vinca alkaloid conjugate[END_REF].

Folic acid receptor--targeted NPs for drug delivery

Folic acid was attached to poly(ethylene glycol)-poly(ε-caprolactone) to form large 70 nm micelles that also contained doxorubicin (Dox, 4.57 wt.%) covalently transported via a hydrazone linker (FA-hyd-PECL-hyd-Dox) [START_REF] Guo | pH-triggered intracellular release from actively targeting polymer micelles[END_REF]. Because of the presence of the hydrazine linker, Dox release is under pH control and should occur in the acidic environment of endosomes only after intracellular internalization. This resembles the disulfide selfimmolating bridge present in the vintafoldine compound, although it reacts to pH rather than to a redox-reaction. Indeed, 70% of the Dox content is released after 10 hours at pH=5.0. Of note, this is also associated with a swelling of the NP that may be beneficial for the improved retention of the NP and eventually the initiation of the drug release process in the acidic microenvironment of tumors, but such a phenomenon is not documented in the publication.

The process is associated with an augmented IC50 of the formulated Dox (0.51 mg/mL) that is 3 times higher than free Dox (IC50, 1.51 mg/mL). The half-life of the formulated Dox is also significantly augmented (± 14 hours versus 1.5 for free Dox), and its body distribution is largely affected. In particular, a ±10 times-augmented concentration of Dox in the tumor was found to be due to the targeted formulation, reaching a sustained concentration of the targeted particles of more than 6 µg/g in 24 hours. In the absence of folic acid, the untargeted formulation also increases tumor targeting by the EPR effect but the peak of concentration was reduced to 6 h. In terms of active targeting, the presence of folic acid reduced the accumulation of micelles in the liver, heart, spleen and lung while augmenting its accumulation in the tumor. Overall, this was associated with a reduced general toxicity after treatment with a dose of 5 mg/kg of Dox and a dramatic increase in the survival rate after 45 days with targeted micelles (60%) versus naked micelles (30%); conversely, all the mice were dead at day 33 with free Dox. These results are in the same range as those obtained with vintafolide because several SMDCs (small-molecule drug conjugates), such as vintafolide, that are currently being studied provide safe delivery of 10-to 20-fold more drug to FRexpressing tumors compared to the un-conjugated drug when the latter was dosed at MTD (C Leamon, personal communication). It is thus reasonable to assume that SMDCs that present a 1:1 ratio of FA:drug diffuse more efficiently than large 70 nm FA-hyd-PECL-hyd-Dox NPs, even though the latter will deliver a larger amount of drug/vector.

Another level of complexity in the formulation of Dox was recently published [START_REF] Zou | Programmed packaging of mesoporous silica nanocarriers for matrix metalloprotease 2-triggered tumor targeting and release[END_REF] and is based on the folic acid-mediated targeting of mesoporous silica NPs. Because mesoporous NPs contain empty cavities, they can be loaded with drugs, which however will usually leak out as easily as it can be loaded. In this study, folate-covered mesoporous NPs were covered by gelatin, which masked the holes and the surface of the NPs. PEGs were added to the gelatin to improve bio-availability and the EPR effect (folic acid is embedded and not accessible). Therefore, such a nanomedicine will passively target a tumor via the EPR effect and encounter a favorable gradient of tumor-associated matrix metalloprotease 2 (MMP-2), which digests the gelatin and thus uncovers the folic acid. Thereafter, the folic acid enhances the internalization of the NP and the specific delivery of Dox. When injected intravenously, these NPs show improved performance in terms of Dox antitumor activity, without systemic toxicity.

The importance of nanomedicine in regard to NP multi-functionality was also recently demonstrated using a sequentially active multi-targeted nanovector [START_REF] Satsangi | Synthesis of a novel, sequentially active-targeted drug delivery nanoplatform for breast cancer therapy[END_REF]. In this study, paclitaxel (PTX) was delivered in vitro and in vivo in mice using large, ± 200 nm FA-targeted dendrimers. In addition to FA-targeting, the liberation of PTX was under the control of cathepsin B, a protease overexpressed in tumors, which was achieved through a cathepsin Bcleavable tetrapeptide conjugating PTX to the poly(amidoamine) dendrimer in the core of the particle.

Targeting FR with a small ligand was a successful strategy and resulted in the rapid clinical translation of a self-assembling 100 nm polymeric NP (BIND-014) that delivers docetaxel for cancer therapy (now in phase II trials in patients with NSCLC, urothelial carcinoma, cholangiocarcinoma, cervical cancer, and squamous cell carcinoma of the head and neck).

The polymeric matrix consists of PEG-poly(D,L-lactic acid) (PEG5k-PLA16k) mixed with PEG5k-PLA16k-ACUPA in a 97.5% to 2.5% ratio. The ACUPA (S,S-2- [3-[5-amino-1carboxypentyl]-ureido]-pentanedioic acid) moiety is a PSMA (prostate-specific membrane antigen) substrate analog inhibitor. The pharmacokinetic profile of BIND-014 did not differ among mice, rats, monkeys and human [START_REF] Hrkach | Preclinical development and clinical translation of a PSMA-targeted docetaxel nanoparticle with a differentiated pharmacological profile[END_REF] and showed a blood circulation half-life of 20 hours. The plasma levels of docetaxel were 100 times more elevated when administered in humans as BIND-014 compared to its usual solvent-based formulation. In terms of therapeutic efficacy in humans, the initial results are very promising [START_REF] Hrkach | Preclinical development and clinical translation of a PSMA-targeted docetaxel nanoparticle with a differentiated pharmacological profile[END_REF].

Folic acid receptor--targeted NPs as contrast agents

For imaging purposes, the folic-acid-targeted peptide derivative EC20, similar to that used for vintafolide, was labeled with 99m Tc (73) and used for SPECT imaging [START_REF] Fisher | Exploratory study of 99mTc-EC20 imaging for identifying patients with folate receptor-positive solid tumors[END_REF]. In parallel, folic acid labeled with FITC was injected in patients for intraoperative near-infrared optical-guided surgical resection of ovarian cancer [START_REF] Van Dam | Intraoperative tumor-specific fluorescence imaging in ovarian cancer by folate receptor-alpha targeting: first in-human results[END_REF]. Based on the status of the receptors and the success of targeting, it is thus tempting to generate targeted nano-objects, as recently performed ( 8), whereby folic acid was used to generate folate-PEG3400-DSPE liposomes of ± 100 nm loaded with a near-infrared dye (DiD) or with 3 H . In this preclinical study, 2 different FR + tumors were engrafted subcutaneously in mice; the animals were also presenting either ulcerative colitis or inflamed injured muscles to investigate how the targeted NPs behaved in the inflammatory milieu. The results clearly established that folate-targeted and nontargeted NPs accumulate more efficiently at sites of inflammation than in solid tumors.

Targeting integrins with peptides

The integrin family is composed of 24 αβ heterodimeric members that mediate the attachment of cells to the extracellular matrix (ECM); only 8 recognize the RGD tripeptide sequence and in particular α V β 3 , an integrin selectively overexpressed on the surface of neoangiogenic endothelial cells and on (migrating) tumor cells. One of the best studied RGD-peptide ligands for α V β 3 integrin is c(RGDf-N(Me)-V), which is also known as EMD121974 or Cilengitide.

This RGD peptide displays affinity in the subnanomolar range. Importantly, Cilengitide displays a 1000-fold preference for α V β 3 over α IIb ß 3 , a very similar integrin expressed on platelets [START_REF] Dechantsreiter | N-Methylated cyclic RGD peptides as highly active and selective alpha(V)beta(3) integrin antagonists[END_REF]. The multivalent binding of α V β 3 leads to their clustering, which in turn activates various kinases and signaling events [START_REF] Sancey | Clustering and Internalization of Integrin alphavbeta3 With a Tetrameric RGD-synthetic Peptide[END_REF].

Cilengitide [START_REF] Stupp | Cilengitide combined with standard treatment for patients with newly diagnosed glioblastoma with methylated MGMT promoter (CENTRIC EORTC 26071-22072 study): a multicentre, randomised, open-label, phase 3 trial[END_REF][START_REF] Millard | Integrin targeted therapeutics[END_REF][START_REF] Juliano | Integrin targeted delivery of gene therapeutics[END_REF][START_REF] Mas-Moruno | Cilengitide: the first antiangiogenic small molecule drug candidate design, synthesis and clinical evaluation[END_REF] and several monoclonal antibodies, such as anti-α V β 3 (Vitaxin, Abegrin (etaracizumab) [START_REF] Hersey | A randomized phase 2 study of etaracizumab, a monoclonal antibody against integrin alpha(v)beta(3), + ordacarbazine in patients with stage IV metastatic melanoma[END_REF], CNTO-95 (intetumumab), c7E3), were evaluated for possible antiangiogenic activity in clinical trials. To date, most have failed to demonstrate efficacy, and some even accelerated tumor progression, as demonstrated in prostate cancer patients [START_REF] Heidenreich | A randomized, double-blind, multicenter, phase 2 study of a human monoclonal antibody to human alphanu integrins (intetumumab) in combination with docetaxel and prednisone for the first-line treatment of patients with metastatic castrationresistant prostate cancer[END_REF].

Cilengitide failed in phase III trials because of a lack of anti-angiogenic activity in glioblastoma [START_REF] Williams | Discontinued in 2013: oncology drugs[END_REF]. It was also described that RGD-peptides or mimetic could present a very modest anti-angiogenic activity at concentrations greater than one micromolar but also proangiogenic activity if delivered at too low (nanomolar) a concentration, in which case they promoted VEGF-mediated angiogenesis [START_REF] Reynolds | Stimulation of tumor growth and angiogenesis by low concentrations of RGD-mimetic integrin inhibitors[END_REF]. Indeed, intravenous injection of a dose of 200 mg kg -1 of RGD-mimetic or cilengitide in mice was associated with micro-molar plasma concentrations in a few hours, which rapidly decreased to nanomolar levels that persisted for 16-24 hours after administration. This phenomenon should be kept in mind when using targeted molecules that bind and cross-link homo-or heterodimeric receptors, thus generating different intracellular signaling cascades in a dose-dependent manner.

Nonetheless, RGD-based molecules are interesting targeted vectors for imaging, as was described for PET imaging in patients using 18 F-or 68 Ga-labeled Galacto-RGD [START_REF] Haubner | PET radiopharmaceuticals for imaging integrin expression: tracers in clinical studies and recent developments[END_REF][START_REF] Schnell | Imaging of integrin alpha(v)beta(3) expression in patients with malignant glioma by [18F] Galacto-RGD positron emission tomography[END_REF][START_REF] Kenny | Phase I trial of the positron-emitting Arg-Gly-Asp (RGD) peptide radioligand 18F-AH111585 in breast cancer patients[END_REF][START_REF] Beer | Biodistribution and pharmacokinetics of the alphavbeta3-selective tracer 18F-galacto-RGD in cancer patients[END_REF]. These studies demonstrated a high level of detection of primary tumors (80 to 100%), but with reduced detection for distant metastasis or invaded lymph nodes. However, inflammatory regions were also detected, and it is thus difficult to distinguish between benign and malignant lesions. Furthermore, atherosclerotic plagues were also detected ( 89) because activated macrophages express the integrin.

We also demonstrated the importance of RGD-based multimeric peptides for near-infrared optical imaging and PET and SPECT imaging of tumors as well as for the imaging of physioand pathological angiogenesis in mice [START_REF] Jin | In vivo noninvasive optical imaging of receptormediated RGD internalization using self-quenched Cy5-labeled RAFT-c(-RGDfK-)(4)[END_REF][START_REF] Jin | In vivo optical imaging of integrin alphaV-beta3 in mice using multivalent or monovalent cRGD targeting vectors[END_REF][START_REF] Jin | Noninvasive visualization and quantification of tumor alphaVbeta3 integrin expression using a novel positron emission tomography probe, 64Cu-cyclam-RAFT-c(-RGDfK-)4[END_REF][START_REF] Briat | Gelofusine reduces renal uptake of RAFT-RGD. Gelofusine reduced renal uptake of RAFT-RGD, a tetrameric RGD-based peptide that specifically targets the integrin alphavbeta3, without affecting its local distribution[END_REF][START_REF] Jin | Positron emission tomography imaging of tumor angiogenesis and monitoring of antiangiogenic efficacy using the novel tetrameric peptide probe 64Cu-cyclam-RAFT-c(-RGDfK-)4[END_REF][START_REF] Keramidas | Noninvasive and Quantitative Assessment of In Vivo Angiogenesis Using RGD-Based Fluorescence Imaging of Subcutaneous Sponges[END_REF][START_REF] Garanger | New multifunctional molecular conjugate vector for targeting, imaging, and therapy of tumors[END_REF]. Concerning therapeutic applications, we reported that RGD-based peptides can be useful in optical-guided tumor surgery of head and neck cancer [START_REF] Atallah | Near-infrared fluorescence imaging-guided surgery improves the recurrence-free survival rate in a novel orthotopic animal model of HNSCC[END_REF], peritoneal carcinomatosis (98) and infiltrative fibrosarcoma in cats [START_REF] Wenk | Near-infrared optical guided surgery of highly infiltrative fibrosarcomas in cats using an anti-alphav beta3 integrin molecular probe[END_REF]. We also proved the delivery of toxic peptides (100-102); however, these last studies noted that the 1:1 ratio between the targeting RGD moiety and the toxic peptide was not sufficiently strong.

Because we found that it was necessary to deliver larger amounts of the drug at the same time, we decided to generate a targeted NP. We used in parallel an NP of a very small size (<5 nm) (103, 104) or large 50 nm nanoemulsions [START_REF] Goutayer | Tumor targeting of functionalized lipid nanoparticles: assessment by in vivo fluorescence imaging[END_REF]. Several other groups are also working in this direction, and numerous publications report the synthesis and use of RGD-targeted NPs (reviewed in ( 59)). In particular, a silica-based small NP (<7 nm) was used as a multimodal NIR-optical imaging and PET imaging contrast agent (labeled with 124 I) (105) in the presence of an RGD-targeting ligand. In preclinical models, these targeted NPs provided a ± 2-fold increase in the targeting of a melanoma subcutaneous tumor, with an overall modest 1.5%

ID/g at 4 hours due to the presence of RGD. This study was further extended to a phase I clinical trial in humans [START_REF] Phillips | Clinical translation of an ultrasmall inorganic optical-PET imaging nanoparticle probe[END_REF], which essentially confirmed the pharmacodynamics data obtained in mice but provided very modest tumor detection, with short retention times in a melanoma metastasis present in the liver [START_REF] Phillips | Clinical translation of an ultrasmall inorganic optical-PET imaging nanoparticle probe[END_REF]. These preclinical data are in agreement with our own, which also indicated a ±2-fold increase in RGD-mediated targeting of U87MG glioblastoma subcutaneous tumors using gadolinium-based Small Rigid Platforms (103).

Similarly, we also obtained a tumor versus skin fluorescence ratio of 1.53 ± 0.07 at 24 hours after intravenous administration of large 35 nm lipid NPs in mice bearing subcutaneous tumors that overexpress very high levels of α V β 3 integrin. However, using larger NPs (± 120 nm), the presence of RGD peptides did not augment active targeting (106) in vivo, as evidenced using 19 F MRI.

These results are thus disappointing and raise questions concerning the added value of the presence of RGD in terms of the efficacy of tumor targeting. Regardless, this targeting may be more interesting for the delivery of theranostic agents, as demonstrated for siRNA (107).

The remainder of this chapter focuses on the work that has been performed using RGDtargeted gold NPs because they serve as an example of promising inorganic theranostic agents.

RGD--targeted gold NPs

Gold (Au) is a very interesting product because its assembly can produce NPs of very different shape in a controlled manner. AuNP shapes include nanoshell, nanorod, nanocage, nanostar, and nanopopcorn (108) and are used for drug delivery and imaging applications.

Because of their high atomic number, small size and chemical properties that make them amenable to functionalization, AuNPs have unique properties for medical purposes (109). In particular, these non-toxic inert particles can serve as prodrugs because they absorb photons as a function of their size and aspect ratio and release this accumulated energy by converting it into heat (110) used for phototherapy under near-infrared illumination (111) or because of their radiosensitizing effect under X-ray irradiation [START_REF] Zhang | Enhanced tumor accumulation of sub-2 nm gold nanoclusters for cancer radiation therapy[END_REF][START_REF] Zhang | Ultrasmall Au(10-12)(SG)(10-12) nanomolecules for high tumor specificity and cancer radiotherapy[END_REF]112,113). They can also serve as contrast agents for X-ray imaging, with a three time greater attenuation than iodine, and also for photoacoustic imaging, allowing tumor detection and imaging prior to remote activation (111).

Chen and colleagues used large 26 nm AuNPs to develop a chemical method of PEGylation that allowed the controlled addition of RGD-targeting ligands (114). They proved in vitro that the presence of an increasing amount of RGD was associated with an increased efficacy of gold internalization in a dose-dependent manner, but this effect was mainly linked to unspecific binding when the peptide/AuNP ratio exceeded 50 (118). In a recent study, the impact of gold-mediated hyperthermia on tumors was evaluated non-invasively using a dendrimeric RGD-targeted multimodal nanoprobe followed by MRI. This study demonstrated that RGDtargeting is indeed capable of inducing a high tumor/background ratio but also can serve to evaluate the therapeutic impact of a given treatment on the tumor vasculature (119).

Because it can provide photothermal destruction of malignant tissues and because of its apparent lack of toxicity in preclinical studies, gold is thus expected to be crucial in clinical applications. Nonetheless, long-term toxicity studies need to be performed (108) and a more precise standardization of the procedures needs to be implemented (120). It should be noted that the covalent labeling of antibodies with 50 nm gold NPs prevented their capacity to generate the antibody-dependent cellular cytotoxicity (ADCC) that is normally present and part of the antibody-mediated antitumor effect (121). Thus, the use of antibody-targeted gold NPs may not be feasible if we want to retain the innate and full activity of the therapeutic antibody; nonetheless, this may also be the case with the covalent labeling of many other antibodies.

Protein--targeted NPs

Among the 8 different targeted NPs used in clinical phases (4), four of them (MBP426;

CALAA-01; SGT-53 and SGT-94) were directed against transferrin receptor (TfR).

Targeting transferrin receptor

Transferrin (Tf) is a soluble protein that delivers iron to cells via its binding to TfR with an affinity ranging from 1 to 10 nM. Because cancer cells are actively proliferating, they require iron and thus over-express TfR (122), though it is normally ubiquitously expressed. After clathrin-mediated endocytosis, iron is released into early endosomes, and TfR is then recycled back to the cell surface. TfR is also abundantly expressed in the cerebral endothelium, where it can facilitate iron-loaded transferrin uptake from the bloodstream and its release into the brain parenchyma (123). Depending on the transferrin content (from low to moderate to high),

NPs can fail to engage TfR on endothelial cells (low Tf content), adhere to the endothelium without being transferred (elevated Tf content), or undergo receptor-mediated transcytosis (moderate Tf content). This "bell curve" that correlates the Tf biological efficacy to its density on the NP surface is thus similar to that observed using RGD.

MBP-426 (Mebiopharm) is a a transferrin-liposome-oxaliplatin conjugate. Transferrin is covalently attached via a pH sensitive N-glutarylphosphatidylethanolamine(NGPE) linker that allows oxaliplatin release in acidic (pH 5.5) endomes (T Fujisawa, Mebiopharm, Personal communication). After completion of a phase 1 and 1b studies on a total of 49 patients (Clinical trial ID : NCT00355888) it is currently in phase 2a on 19 patients as second line treatment for gastric, gastroesophageal and esophageal adenocarcinomas (Clinical trial ID: NCT00964080) in combination with leucovorin (folinic acid or FA) and fluorouracil .

The treatment were well tolerated (no over Gr 2 peripheral neurotoxicity reported) and a tumor reduction of more than 30% (PR) was observed in a patient resistant to platinum.

Based on an initial preclinical works in mice and non human primates (124-127), CALAA-01, a four component NP (128) that used the Rondel™ patented small RNA delivery technology.

It was made of linear, cyclodextrin-containing cationic polymers (CDPs) targeted by transferrin (Tf) and siRNA was manufactured and tested in humans (129). It was manufactured in two vials that were mixed just before injection so that the different

Targeting the Epithelial Growth Factor Receptor

Pegylated liposomal doxorubicin (Doxil ® /Caelyx ® ) loaded with doxorubicin and presenting covalently linked Fab' fragments of cetuximab (anti-EGFR ILs-dox) (139) were also used in a phase I clinical trial in patients with solid tumors (Clinical trial ID: NCT01702129) (140) with promising results (1 complete remission) that necessitate to be investigated in further trials.

Biodistribution studies in mice showed extended half life in blood (t1/2 = 21 hours). Here also, an efficient accumulation in tumors was observed (±15% ID/g) but was not associated with the presence of the targeting agent in EGFR-overexpressing tumors. But in this case also, the targeting ligand was augmenting the internalization of the drug carrier within tumor cells (92% of analyzed cells versus <5% for nontargeted liposomes) ( 141) and an improved antitumor activity of doxorubicin, epirubicin, and vinorelbine in preclinical models. Very interestingly also, the effect of such drug loaded immunoliposomes did not depend on the inhibition of the EGFR-receptor but on the internalization and intracellular delivery of the drug (142). It was also proved using Multidrug Resistant (MDR) positive cells that Dox molecules released intracellularly by these immunoliposomes were able to bypass membranebound efflux (143) and thus produced 19-216-fold greater cytotoxicity than free Dox in vitro and an augmented toxicity in vivo (142). This is an important point because it indicates that when Dox is delivered directly into the cytoplasm by a nanovector, it could bypass the membrane-associated P-glycoprotein (one of the main contributor to the MDR phenotype) present on some chemioresistant tumor cells, thus overcoming their resistance as demonstrated since a long time for several NP (144-147).

Conclusion

In summary, active targeting does not usually augment the number of NP that will accumulate in a tumor because this depends mainly on the EPR effect. However, active targeting may generate an augmented receptor-mediated internalization of the NP that will be associated with a specific intracellular trafficking. This may improve the drug efficacy in some cases if the drug is naturally not presenting good pharmacokinetics properties or also by preventing its efflux via the MDR system. However, the price to pay for this improvement must be carefully evaluated on a case by case basis.

Figure 14 .

 14 Figure 14. Expression profiles of αvβ3 and NRP-1 in human prostate cancer cell lines and other reference cell lines Numbers in the upper-right corners represent MFI values of antibody-treated cells over that of control cells.

Figure 15 .

 15 Figure 15. Structure of LNPs (A) Appearance of final product LNPs in white light. (B). Full structure of LNP and components (drawing credits to CEA-Leti). (C) Structural formulas of peptide ligands on the LNP surface: the functionalization 1 LNPs as representative examples. The functionalization 2 LNPs have the maleimide linker in the inversed orientation between the peptide and the PEG residue (adapted drawing).

Figure 16 .

 16 Figure 16. Normalized absorbance and emission spectra of F1-LNPs Measured in 1X PBS. The emission spectrum was measured with excitation at 790nm.

Figure 17 .Figure 18 .

 1718 Figure 17. Fluorescence intensity curve of three F1-LNPs measured under the NIR in vivo camera Fluobeam®800

  and other chemicals for the LNP preparation were purchased from Sigma-Aldrich (Saint-Quentin-Fallavier, France). Cyclic RGD stands for c[DfK(Mal)RG] and cRAD for c[DfK(Mal)RA] with f = D phenylalanine and Mal = 3 Maleimide Propionic Acid.

1 .

 1 Bright field microscopy HEK293(b 3 ) cells (5 Â 10 4 cells/well) were seeded in an 8-well chambered coverglass(Nunc Lab-Tek, Thermo Fisher Scientific)

Fig. 1 .

 1 Fig. 1. LNP composition. (i) Schematic representation of the LNPs. (ii-iv) Chemical structure of the targeted or non-targeted surfactants used in the different LNPs. The cyclic RGDfK (iii, targeting) and cRADfK (iv, control) peptides were conjugated on the PEG by the maleimide-thiol coupling reaction. The illustration was adapted from (Merian et al., 2015).

Fig. 2 .

 2 Fig. 2. Integrin-specific binding of cRGD-LNPs. (A) The presence of the a v b 3 integrin was detected by flow cytometry on HEK293(b 3 ), DU145 and PC3 cells. Grey: antibody-labeled cells; black: control cells (no antibody). (B) The capacity of cells to stick to the vitronectin coating was measured for each cell line (HEK293(b 3 ), DU145 and PC3) in the presence of different LNPs at 4.5, 9, 45 and 90 nM. The number of adherent cells was measured by methylene blue staining. Mean AE SD, n = 6-8. The values were normalized with those of the standard LNP. Asterisks: ** = p<0.01, *** = p<0.001, compared to the std LNP. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

3. 3 .

 3 cRGD-mediated detachment of a v b 3 -positive cells Having confirmed that cRGD-LNPs can bind onto the freelyaccessible a v b 3 receptor when the cells are in suspension, we next explored their capacity to reach and interfere with occupied a v b 3 -

Fig. 3 .

 3 Fig. 3. Confluent HEK293(b 3 ) cells treated with LNPs at 45 or 4.5 nM. The cells were observed for a period of 3 h in real time under the microscope equipped with a culture system for live cell imaging. Scale bar = 100 mm.

Fig. 4 .

 4 Fig. 4. Confocal laser microscopy of HEK293(b 3 )-a v RFP cells in the presence of 45 nM LNPs at 37 C. Blue: nuclei, green: a v RFP. Scale bar = 20 mm. (A) The cells were treated with different LNPs and visualized after 60 min of incubation. No LNP: treated with HBSS. (B) The cells were incubated with cRGD-LNPs for 5, 10, 30 or 60 min.

Fig. 5 .

 5 Fig. 5. Co-localization of cRGD-LNP with a v b 3 in HEK293(b3)-a v RFP cells. The cells were incubated with 45 nM LNPs for an hour at 37 C. Blue: nuclei, red: LNPs, green: a v RFP. Scale bar = 10 mm. (First row) The cells were incubated with cRGD-LNPs for 60 min continuously. (From the second to the fourth row) The cells were treated with LNPs and, 15 min later, the LNP solution was replaced by fresh HBSS. The incubation at 37 C was continued for another 45 min. (Bottom row) The cells were incubated with HBSS buffer only. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

Fig. 6 .

 6 Fig. 6. NIR in vivo imaging of DU145 subcutaneous tumor in mice after injections of the LNPs. NIR fluorescence images are shown for each mouse in grey levels with Min-Max values of: 0-46912. The images were taken 24 h after injections of LNPs. L: liver, T: tumor. Graphs below the images show NIR fluorescence levels on tumors and skin. X-axis: time after injection (hours), Left Y-axis: RLU/pixel/10 ms, Right Y-axis: tumor/skin ratio.

Fig. 7 .Figure 19 .

 719 Fig. 7. Fluorescence levels in the dissected organs and plasma from DU145 tumorbearing mice 24 h after injections of the LNPs. Mean fluorescence intensities/pixel/ 10 ms of each LNP group are shown with standard deviation.

Figure 20 .Figure 21 .

 2021 Figure 20. Fluorescence signals from HEK293-β3 cell suspensions after incubation with F1-LNPs, observed by NIR camera Fluobeam®800 The cells were incubated for 15 min on ice or for 30 to 60 min at 37°C with each LNP at 90 nM LNP concentrations. (A) The same volume of each cell suspensions was transferred in a multiwell plate and pictured. Cell numbers were counted after the fluorescence acquisitions. (B) MFI values normalized by cell numbers are represented in a graph. Mean±SD (n=6, two repeats in triplicate).

Figure 22 .

 22 Figure 22. Red channel images from Figure 21

Figure 23 .

 23 Figure 23. Confocal images of HEK293-β3-αvRFP cells at different times following 15 min of incubation with LNPs The cells were incubated with 45 nM LNPs at 37°C for 15 min and the LNPs-containing solution was removed (left column). The incubation was continued for another 15 min (mid column) or 45 min (right column). Blue: nuclei, red: LNP, green: αvRFP; the colocalized LNP with αvβ3 are seen in yellow. Scale bar=20 µm.

Figure 24 .

 24 Figure 24. Red channel images from Figure 23

Figure 26 .Figure 27 .

 2627 Figure 26. Curves of mean tube lengths during 10 hours of HUVEC tube formation Cells were incubated on matrigel in the absence or presence of different LNPs during 10 hours. Data are shown with mean±SD of triplets.

Figure 28 .

 28 Figure 28. Images of DU145 tumor mice at different time points after LNP injections (supplementary to Figure 27) Fluorescence images are shown in gray levels with Min-Max: 0-32093 (integration time 20 ms). B/W = black and white images under white light.

Figure 29 .Figure 30 .

 2930 Figure 29. Fluorescence analysis of dissected organs from DU145 tumor mice Organs were dissected 24 hours after LNP injections and pictured using the NIR camera.

Figure 31 .

 31 Figure 31. Ex vivo confocal images of 100 µm-thick sections from DU145 tumors, injected with cRGD-or cRAD-LNPs 3D reconstruction of stack images with a saturated volume rendering mode. Objective 64× was used. Blue: DAPI, red: LNP.

Figure 32 .

 32 Figure 32. Ex vivo confocal images of 100 µm-thick sections from DU145 tumors, injected with OH-LNP or vehicle solution (HBSS buffer) 3D reconstruction of stack images with a saturated volume rendering mode. Objective 64× was used. Blue: DAPI, red: LNP.

Figure 33 .

 33 Figure 33. NIR in vivo imaging of U87MG s.c. tumor mice after i.v. injections with F1-LNPs NIR images, taken 24 hours after injections, are shown for each mouse in gray levels with Min-Max values of 0-39424 with an integration time 20 ms. Graphs below the images show NIR fluorescence levels on tumors and skin. X-axis: hours after injection, left Y-axis: RLU/pixel/10ms, right Y-axis: tumor/skin signal ratio.

Figure 34 .

 34 Figure 34. Images of U87MG tumor mice at different time points after LNP injections (supplementary to Figure 33) Fluorescence images are shown in gray levels with Min-Max: 0-32416 for the cRGD, cRAD, Std LNPs (integration time 20 ms) and 0-39424 for the OH-LNP. B/W = black and white images under white light.

Figure 35 .Figure 36 .Figure 37 .Figure 38 .

 35363738 Figure 35. Fluorescence analysis of dissected organs from U87MG tumor mice Organs were dissected 24 hours after LNP injections and pictured using the NIR camera.

Figure 39 .

 39 Figure 39. Flow cytometry analysis of PC3-luc cells, after incubation with non-PEGylated SiNPs in a cell suspension, comparing PBS vs. serum-containing medium SiNPs at 2 nM were incubated with 1 million cells in suspension (volume 200 µL) for 30 min at RT. Histogram: representative result of experiments performed in triplicate. X-axis (FL1-A) = FITC intensity. Bar graphs: mean ±SD of triplicate experiments.

Figure 40 .

 40 Figure 40. Influence of a 30 min incubation with cRGD-siNP in comparison with other NPs on the cell viability (supplementary to Figure 39) SiNPs at 2 nM were incubated with 1 million cells in suspension (volume 200 µL) for 30 min at RT.

Figure 41 .

 41 Figure 41. Flow cytometry analysis of PC3-luc cells, after incubation with non-PEGylated SiNPs on adherent cells in a serum-containing medium PC3-luc cells, placed in 24-well plate (5x10 4 /well), were incubated with 2 nM SiNPs diluted in 300 µL of 10% FBS-containing medium during 30 min at 37°C. Histogram: representative replicate. X-axis (FL1-A) = FITC intensity. Bar graph: mean±SD of duplicate experiments.

Figure 42 .

 42 Figure 42. Flow cytometry analysis of PC3-luc cells, after incubation with PEGylated SiNPs on adherent cells, comparing PBS vs. serum-containing medium PC3-luc cells, in a 12-well plate (2x10 4 /well), were incubated with 1 nM PEGylated SiNPs (volume 500 µL) for 30 min at 37°C. Histogram: representative replicate. X-axis (FL1-A) = FITC intensity. Bar graphs: mean±SD of duplicate experiments.

Figure 43 .

 43 Figure 43. Confocal microscopy images of PC3-luc cells after incubation with non-PEGylated and PEGylated SiNPs PC3-luc cells were incubated with 2 nM NPs diluted in a 10% FBS-containing medium during 30 min at 37°C and washed twice. Cells were fixed with PFA and nuclei were stained by Hoechst 33342. Blue: nuclei, green: FITC (SiNP).

Figure 44 .

 44 Figure 44. Expression level of αvβ5 integrin in PC3-luc cells HUVEC cells were used as a positive control of αvβ5-expressing cells. PC3 luc cells (1x10 6 ) were incubated with a FITC-conjugated mouse mAb anti-human αvβ5 (clone P1F6) during 30 min at 4°C, in PBS containing 1% BSA and 1mM MgCl2/CaCl2. Control cells were treated with an antibody isotype IgG1.

Figure 45 .

 45 Figure 45. Influences of different PEG-SiNPs on HUVEC tube formation Cells were diluted in a serum-containing medium, placed in a 24-well plate coated with normal matrigel (at 50,000 cells/250 µL/well), and incubated during a half day in the presence of different PEG-SiNPs. (A) Images taken after 10 hours of incubation. A 5x objective was used. (B) Characteristics of tube formations at T10h analyzed by Wimass. Mean±SD from images of 5 fields of view per well.

Figure 46 .

 46 Figure 46. Cytotoxicity test using EA.hy926 cells by MTS assay Cells were incubated in a serum-free (SF) medium in the presence of different PEG-SiNPs for 3 days. (A) Cell viability was measured by MTS assay. Data are shown with mean±SD from triplicate experiments. (B) Microscopy images at D3.

Figure 47 .Figure 48 .

 4748 Figure 47. Fluorescence microscopy images of PC3-luc subcutaneous tumors collected 3 hours after injections of PEG-SiNPs (A) Tumors were sectioned at 5 µm and mounted with a DAPI-containing solution. Slides were scanned using a mosaic function with a 63x/oil objective. The presented images were montages of mosaic images. (B) HE-stained tumor sections. Tumor slices are in triplicate.

  ., Omo, A., Elsharawy, M., and Li, J. (2014

  components could self assemble. Vial 1 contained the polymers conjugated to human transferrin. Vial 2 contained an siRNA directed against the M2 subunit of ribonucleotide reductase (RRM2)(130). Here also the preclinical pharmacokinetics data obtained in mice, rats an non human primates were correlating well with those obtained in the 24 patients of the first clinical trial ever made in humans using siRNA. It should be noticed that the NP exhibited similar biodistribution and tumor localization in the presence or absence of the transferrin targeting ligand (126) confirming that the tumor accumulation was EPR driven mainly. However, the transferrin-targeted siRNA NP were reducing of 50% the expression of their targeted gene relative to the nontargeted NP in preclinical studies, suggesting an improved cellular uptake due to the presence of active targeting. Because of adverse effects in this clinical trial (131) the development of the CALAA-01 product using the Rondel™ technology was then stopped and redirected to another delivery system that uses a Roche DPC technology product in use for the treatment of hepatitis B.SGT-53 and SGT-94 are two formulations containing[1,2-dioleoyl-trimethylammonium- propone (DOTAP) dioleoyl phosphatidylethanolamine (DOPE)] forming 60-100 nm large cationic liposomes, and able to target the TfR via a single chain antibody fragment (TfRscFv).These liposomes are loaded with a plasmid expressing the tumor suppressor gene p53 (SGT-53) or the tumor suppressor gene RB94 (SGT-94) (132). This scFv was preferred to transferrin itself because its has a reduced molecular weight (and size) and because it is a recombinant protein rather than a blood product(133, 134). In mice, SGT-53 was capable to deliver the plasmid to different tumors after a systemic injection and to provide antitumor activity when combined to docetaxel. Ti was then translated to a Phase I clinical trial in patients with advanced solid tumors (135). It is now evaluated alone and in combination with topotecan and cyclophosphamide in pediatric patients with recurrent or refractory solid tumors (Clinical trial ID: NCT02354547) or in patients with metastatic pancreatic cancer in combination with Gemcitabine/Nab-Paclitaxel (Clinical trial ID: NCT02340117)) and in Phase II study in combination with Temozolomide for treatment of recurrent Glioblastoma (Clinical trial ID: NCT02340156). This later study is in particular associated with the demonstration that because they target the TfR overexpressed on the cerebral endothelium, these immuno-liposomes can cross the blood brain barrier by TfR-mediated transcytosis and transfer the transported plasmids in the glioblastoma tumor cells but also on the tumor associated cancer stem cells (CSCs), which have been implicated in recurrence and resistance to treatments(136)(137)(138). Based also on preclinical basis (132), SGT-94 entered a Phase I study in patients with solid tumors (Clinical trial ID: NCT01517464).

Figure 1 :Figure 2 :Figure 3 :

 123 Figure 1: Factors such as size, electric charge, surface coating and the presence of targeting

•

  Coll, J.L., Choi, J. Challenges of Applying Targeted Nanostructures with Multifunctional Properties in Cancer Treatments. Nanoscience and Nanotechnology for Human Health, Voorde, B. M. a. M. V. d., Ed. John Wiley & Sons: 2017; pp 127-156. • Choi, J., Rustique, E., Henry, M., Guidetti, M., Josserand, V., Sancey, L., Boutet, J., and Coll, J. L. (2017). Targeting tumors with cyclic RGD-conjugated lipid nanoparticles loaded with an IR780 NIR dye: In vitro and in vivo evaluation. Int J Pharm. DOI: 10.1016/j.ijpharm.2017.03.007.

	•	'Targeted nanoparticles for prostate cancer imaging', Journée Sciences et
		Métiers, UJF, Grenoble, 2014
	•	'Cancer-targeted theranostic nanovectors that can help early diagnosis and
		advanced therapy of prostate cancer', Annual meeting of EDCSV, Grenoble,
		2015
	•	[Prize] 'Cancer-targeted theranostic nanovectors that can help early diagnosis
		and advanced therapy of prostate cancer', Euro-Korean Conference on science
		and technology, Strasbourg, 2015
	•	'Theranostic nanoparticles to help early diagnosis and targeted therapy for
		prostate cancer', SFNano European Nanomedicine Meeting, Grenoble, 2015
	•	'Theranostic nanoparticles for prostate cancer', European Molecular Imaging
		Meeting, Utrecht, Netherland, 2016
	•	'cRGD-Lipid nanoparticle loaded with IR780 iodide for cancer
		NIR/photoacoustic imaging', SFNano European Nanomedicine Meeting, Paris,

• Annual presentations in Ph.D. committee in the institute, Institute for Advanced Biosciences (IAB), La Tronche, 2014, 2015, 2016 • 'Theranostic nanoparticles for prostate cancer', Bio Korea in Europe (BiKiE) symposium, KIST Europe, Saarbrucken, Germany, 2016

Table 7 .

 7 Physico-chemical characteristics of each LNP lots

		LOT #1-F1 (functionalization1)	LOT #1-F2 (functionalization2)	LOT #2-F1 (functionalization1)
		cRGD	cRAD	OH	cRGD	cRAD	OH	cRGD	cRAD	Std
	Final lipid	10%,	10%,	8.5%,	9%,	9%,	7.5%,	9%,	9%,	10%,
	concentration	100 g/L	100 g/L	85 g/L	90 g/L	90 g/L	75 g/L	90 g/L	90 g/L	100 g/L
	Maleimide
	quantity	748 nmol	748 nmol	374 nmol	748 nmol	748 nmol	374 nmol	748 nmol	748 nmol	-
	(initial)
	Peptide
	quantity	750 nmol	750 nmol	-	750 nmol	750 nmol	-	750 nmol	750 nmol	-
	(theory)
	Particle
	concentration	1.5E+15	1.5E+15	1.25E+15	1.35E+15	1.35E+15	1.1E+15	1.31E+15	1.31E+15	1.5E+15
	(/mL)
	Fluorophore concentration	260 µM	260 µM	230 µM	250 µM	250 µM	200 µM	240 µM	240 µM	270 µM
	Number of
	dyes/particle	104	104	111	112	112	109	108	108	110
	(approx.)
	Size (nm)	54.28	53.67	53.51	47.46	47.25	47.73	55.35	55.52	49.49
	Polydispersity index	0.187	0.125	0.142	0.116	0.098	0.122	0.126	0.130	0.126
	Zeta potential
	(mV)

Table 1

 1 Physico-chemical characteristics of the LNPs (triplicate values AE SEM).

		cRGD-LNP	cRAD-LNP	LNP
	Size (nm)	55 AE 1	55 AE 1	49 AE 1
	Polydispersity index, PDI	0.126 AE 0.027	0.130 AE 0.008	0.126 AE 0.021
	Zeta potential (mV)	À4.53 AE 0.38	À4.87 AE 0.39	À4.13 AE 0.81

[Approximately 108 fluorophores (1.8 Â 10 À22 mol) per nanoparticle.]

Table 1

 1

	Query performed with keywords :	-	in cancer	%
	Targeted nanoparticles	6223	3377	54
	Folate receptor targeted nanoparticles	279	237	85
	Integrin targeted nanoparticles	273	173	63
	EGFR targeted nanoparticles	161	147	91
	HER2 targeted nanoparticles	113	102	90
	Transferin receptor targeted nanoparticles	85	53	62
	CD44 targeted nanoparticles	75	61	81
	PSMA targeted nanoparticles	40	37	93
	uPAR targeted nanoparticles	17	16	94
	Peptide targeted nanoparticles	1425	841	59
	Antibody targeted nanoparticles	908	550	61
	Folate targeted nanoparticles	526	417	79
	RGD targeted nanoparticles	176	128	73
	Aptamer targeted nanoparticles	146	106	73
	Transferrin targeted nanoparticles	138	92	67
	Lectin targeted nanoparticles	125	42	34
	Hyaluronic acid targeted nanoparticles	108	80	74
	Glucose targeted nanoparticles	69	31	45
	Galactose targeted nanoparticles	43	16	37

  . In optimized conditions, the presence of RGD increased internalization by 18 times. Starting with large 32 nm gold nanoclusters, another group covalently added RGD and Dox, thus forming larger Au-cRGD-DOX nanoclusters (58 nm of hydrodynamic radius) (115). These targeted NPs were interestingly able to deliver Dox to the target cells, in contrast to Au-DOX NPs, and more importantly to generate a dual toxic effect by combining a near-infrared induced thermal toxicity as well as Dox-mediated chemotherapeutic activity. This strategy was efficient in mice models after intratumoral or intravenous injection of the targeted NPs, confirming previous results that demonstrated the importance of gold nanorods in hyperthermia treatment of tumors(116, 117) and the initial work by Hirsh et al.

G Model IJP 16479 No. of Pages 9

G Model IJP 16479 No. of Pages 9 Please cite this article in press as: J. Choi, et al., Targeting tumors with cyclic RGD-conjugated lipid nanoparticles loaded with an IR780 NIR dye: In vitro and in vivo evaluation, Int J Pharmaceut (2017), http://dx.doi.org/10.1016/j.ijpharm.2017.03.007

Please cite this article in press as: J. Choi, et al., Targeting tumors with cyclic RGD-conjugated lipid nanoparticles loaded with an IR780 NIR dye: In vitro and in vivo evaluation, Int J Pharmaceut (2017), http://dx.doi.org/10.1016/j.ijpharm.2017.03.007

Acknowledgements   Acknowledgements to financial supports Publications and Communications Acknowledgements Jungyoon Choi was supported by the Ligue Nationale contre le Cancer (GB/MA/SC-12742), the Fondation ARC pour la Recherche sur le Cancer (DOC20150602724) and by Campus France (reference: 796786A). This work was supported by a public grant overseen by the French National Research Agency (ANR) as part of the "Investissements d'Avenir" program (reference: ANR-10-NANO-01), but also by the Institut National de la Santé et de la Recherche Médicale (INSERM U1209) and by the France Live Imaging program. The authors acknowledge the team of the "OPTIMAL small animal optical imaging facility" in our institute. Acknowledgments .................................................................................................................... 27 References ................................................................................................................................ 28 Acknowledgments Special thanks to Dr Sandrine Dufort (Nano-H) and N Mignet (Université Paris Descartes) for the help in the reviewing of this document. This work was supported by the INSERM grants as well as the ANR CALIF (ANR-08-NANO-006), GdLung (ANR-12-P2N-0009), NanoLuc (ANR-11-BSV5-0018), Multimage (ANR-12-RPIB-0010-08) and the INCA PLBio PoroCombo.

Abstract ...................................................................................................................................... 3 Introduction ................................................................................................................................ 4 The enhanced permeability and retention effect: biological point of view ............................ 5 The enhanced permeability and retention effect: biophysical perspective ............................. 6 The targeting dilemma ........................................................................................................... 6 Physico-chemical factors that influence NP properties and their biodistribution ...................... 8 Influence of the size of the NP ............................................................................................... 8 Surface modification and opsonization ................................................................................ 10 Electric charge ...................................................................................................................... 11 Density of ligands ................................................................................................................. 12 Targeted NPs ............................................................................................................................ 13 Choice of target receptor ...................................................................................................... 14 Targeting folate receptor using folic acid as an example of a small ligand ......................... 15 Folic acid receptor-targeted NPs for drug delivery .......................................................... 15 Folic acid receptor-targeted NPs as contrast agents ......................................................... 18 Targeting integrins with peptides ......................................................................................... 18 RGD-targeted gold NPs ................................................................................................... 21 Protein-targeted NPs ............................................................................................................ 23 Targeting transferrin receptor ........................................................................................... 23 Targeting the Epithelial Growth Factor Receptor ............................................................ 25 Conclusion ................................................................................................................................ 26

Abstract

The efficacy of conventional chemotherapy as well as that of targeted therapy using small molecules is very often hampered by the poor pharmacodynamics properties of the molecules and their important toxicity. The recent access to innovative nanostructured objects with particular functional properties allows us to re-formulate "old" drugs, and it also allow the formulation of other drugs that were not suitable for systemic administration.

Nano-vectors can target cancer cells while avoiding healthy tissues. They can be engineered in order to be rapidly cleared from the body or, on the contrary, to remain for extended period of time in circulation. In addition, nanomaterials can also combine multiple modalities of cancer diagnostic and treatment such as radiosenzitization or photo-dynamic therapy. Finally they can also deliver the drugs directly in subcellular compartments and avoid the multi-drug resistance system. The specific and passive accumulation of these nanomedicines in the tumors is due to the so-called "enhanced permeability and retention" (EPR) effect. Several attempts were made also to introduce specific ligands on their surface in order to generate an active targeting and uptake by the receptor-overexpressing tumor cells only, but the efficacy of this strategy is not yet completely established. Only a very restricted amount of targeted formulations are tested in humans so far. The interest and limitations of this strategy will be presented in this chapter, based on the description of representative examples.