VEGF- and VEGFR2-Targeted Liposomes for Cisplatin Delivery to

Sep 21, 2016 - The purpose of the present study was to develop the liposomal drug delivery system combining enhanced loading capacity of cisplatin and...
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VEGF- and VEGFR2-Targeted Liposomes for Cisplatin Delivery to Glioma Cells Sergey A. Shein,†,‡,◆ Ilya I. Kuznetsov,§,◆ Tatiana O. Abakumova,† Pavel S. Chelushkin,∥,⊥ Pavel A. Melnikov,# Anna A. Korchagina,† Dmitry A. Bychkov,§ Irina F. Seregina,§ Mikhail A. Bolshov,§,¶ Alexander V. Kabanov,§,∇ Vladimir P. Chekhonin,†,# and Natalia V. Nukolova*,†,○ †

Department of Fundamental and Applied Neurobiology, Serbsky Medical Research Center of Psychiatry and Narcology, Moscow, Russia ‡ Department of Molecular and Cellular Biology, The International Biotechnology Center Generium, Volginsky Village, Russia § Chemistry Department, Lomonosov Moscow State University, Moscow, Russia ∥ Institute of Macromolecular Compounds, Russian Academy of Sciences, St. Petersburg, Russia ⊥ Institute of Chemistry, St. Petersburg State University, St. Petersburg, Russia # Department of Medical Nanobiotechnology, Russian National Research Medical University, named after N.I. Pirogov, Moscow, Russia ¶ Institute for Spectroscopy, Russian Academy of Sciences, Troitsk, Russia ∇ Center for Nanotechnology in Drug Delivery, Molecular Pharmaceutics Division, UNC Eshelman School of Pharmacy, Chapel Hill, North Carolina 27599, United States ○ Koch Institute for Integrative Cancer Research, Massachusetts Institute of Technology, Cambridge, Massachusetts 02139, United States S Supporting Information *

ABSTRACT: Targeted delivery of anticancer drugs to brain tumors, especially glioblastoma multiforme, which is the most frequent and aggressive type, is one of the important objectives in nanomedicine. Vascular endothelial growth factor (VEGF) and its receptor type II (VEGFR2) are promising targets because they are overexpressed by not only core tumor cells but also by migrated glioma cells, which are responsible for resistance and rapid progression of brain tumors. The purpose of the present study was to develop the liposomal drug delivery system combining enhanced loading capacity of cisplatin and high binding affinity to glioma cells. This was achieved by using of highly soluble cisplatin analogue, cisdiamminedinitratoplatinum(II), and antibodies against the native form of VEGF or VEGFR2 conjugated to liposome surface. The developed drug delivery system revealed sustained drug release profile, high affinity to antigens, and increased uptake by glioma C6 and U-87 MG cells. Pharmacokinetic study on glioma C6-bearing rats revealed prolonged blood circulation time of the liposomal formulation. The above features enabled the present drug delivery system to overcome both poor pharmacokinetics typical for platinum formulations and low loading capacity typical for conventional liposomal cisplatin formulations. KEYWORDS: PEGylated liposomes, cisplatin, anti-VEGF and anti-VEGFR2 monoclonal antibody, glioma, targeted drug delivery

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intended to modulate CDDP toxicity were unsuccessful.3 Efforts to create novel platinates with decreased toxicity had only partial success: out of 3000 developed compounds only 35 entered clinical trials,2 with only five approved for clinical use either worldwide (carboplatin4 and oxaliplatin5) or regionally

INTRODUCTION

Cisplatin (CDDP) is one of the most potent and widely used antitumor agents.1 It displays activity against a wide variety of solid tumors and is approved for the clinical treatment of various tumors, including (but not limited to) metastatic testicular, breast, advanced bladder, and nonsmall-cell lung cancers.1,2 However, severe side effects (such as nephrotoxicity, neurotoxicity, and ototoxicity, as well as nausea and vomiting)1,2 associated with off-target CDDP distribution significantly limit its clinical use. Biochemical approaches © 2016 American Chemical Society

Received: Revised: Accepted: Published: 3712

June 9, 2016 August 30, 2016 September 21, 2016 September 21, 2016 DOI: 10.1021/acs.molpharmaceut.6b00519 Mol. Pharmaceutics 2016, 13, 3712−3723

Article

Molecular Pharmaceutics (nedaplatin,6 heptaplatin,7 and lobaplatin8). Most of the above analogues reveal improved toxicity profiles compared to CDDP, but this advantage is acquired at the expense of decreased efficacy. Directed CDDP delivery using nanoparticles seems to be a more promising strategy since it can potentially improve the toxicity profile and retain high antitumor efficacy. To date, a wide variety of different nanocontainers have been offered as CDDP carriers.9 Being of nanoscale size, they all exploit the enhanced permeability and retention (EPR) effect to accumulate passively in tumors.10 Among various nanocarriers, liposomes (i.e., spherical lipid bilayers of 50−500 nm with an aqueous interior) are the most promising. Several liposomebased drugs have been approved for clinical use.11 They are intrinsically biocompatible since naturally occurring lipids are used in liposomal compositions and their outer surface can be easily modified thus imparting various properties. For example, variation of surface charge and PEGylaton provides “stealth” properties to liposomes, while conjugation with targeting groups improves their site-directed delivery. As a result, several liposomal CDDP formulations (lipoplatin12 and SPI-7713) entered advanced stages of clinical evaluation.9,11 The main obstacle impeding further advancement of liposomal CDDP formulations is low cisplatin solubility in water, leading to insufficient loading capacity of these nanocarriers. Recently, this drawback was addressed by encapsulating cis-diamminedinitratoplatinum(II) (CDDP3) into liposomes.14,15 CDDP3 is a more soluble CDDP analogue, which readily converts into CDDP in saline media. As a result, loading capacity of CDDP3loaded liposomes can be drastically increased, thus providing ability to use lower doses of liposomal formulations. Despite the wide use of liposomal formulations for CDDP delivery, there is a lack of reports on the treatment of intracranial models of brain tumors.16 This is caused either by ineffective liposome penetration into intracranial neoplasms at intravenous administration17 or by a variety of neuropathologic side effects at intracerebral administration.16 Functionalization of liposomes by targeting groups could be the first step toward an improvement of drug-loaded liposomes biodistribution and accumulation in the brain tumor. Vascular endothelial growth factor (VEGF) and its receptor type II (VEGFR2) are promising targets for brain tumors since they are reported to be overexpressed by not only core tumor cells, but also migrated glioma cells that are responsible for resistance and rapid progression of brain tumors.18,19 Though VEGF and VEGFR2 are also expressed by healthy cells, the efficient targeted delivery of nanoparticles to VEGF- and VEGFR2positive tumors can be achieved because of (i) overexpression of these targets in malignant cells (VEGF upregulation appears during the development of some pathological processes like a hypoxia, injury, inflammation, tumor, etc.); (ii) low level of endocytosis in VEGF-positive healthy cells compared to malignant cells; (iii) differences in the structure of vessels: vessels in the healthy tissues limit interaction of liposomes with healthy VEGF-positive cells, whereas leaky vasculature in tumors does not hinder such interaction.18,20 Recently, we have developed two novel hybridomas producing monoclonal antibodies to both the above targets (anti-VEGF and antiVEGFR2 mAbs) that reveal a high affinity to glioma C6 and U87 MG cell lines.21,22 Based on this finding we have designed the liposomal drug delivery system combining enhanced loading capacity and high binding affinity to VEGF and VEGFR2. The present study

describes our work on preparation and physicochemical characterization, in vitro analysis, and pharmacokinetics of obtained targeted liposomes.

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EXPERIMENTAL SECTION Materials. Cholesterol (Chol); L-α-phosphatidylcholine (egg, chicken) (PC); 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-2000] ammonium salt (DSPE-PEG2000); 1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-[maleimide (polyethylene glycol)-2000] ammonium salt (DSPE-PEG2000-maleimide); 1,2-dipalmitoyl-snglycero-3-[phospho-rac-1-glycerol] sodium salt (DPPG) were purchased from Avanti Polar Lipids (USA). Didodecyldimethylammonium bromide (DDAB) and silver nitrate (AgNO3) were purchased from Fluka (Switzerland). cis-Diammineplatinum(II) dichloride (CDDP), 2-iminothiolane, and 1,1′dioctadecyl-3,3,3′,3′-tetramethylindocarbocyanine perchlorate (DiI) were from Sigma-Aldrich (USA). All chemicals were of analytical grade and used as received. All aqueous solutions were prepared using double distilled water. Monoclonal antibodies against VEGF and VEGFR2 were obtained by hybridoma technology.21,22 Antibodies were produced in the ascites and purified on Protein A agarose (Invitrogen, USA). Nonspecific mouse immunoglobulins (IgG) purified from a Sp2/0-Ag14 ascites were used as a negative control. Cell Lines. C6 rat glioma and HEK 293 cells were grown in DMEM supplemented with 1 mM sodium pyruvate, 2 mM Lglutamine, 1% antibiotic−antimycotic mixture (10 000 units/ mL penicillin, 10 000 units/mL streptomycin, 25 μg/mL amphotericin B), and 5% FBS; U-87 MG human glioblastoma cells were grown in DMEM with 1 g/mL glucose (37 °C, 5% CO2). Cell dissociation was achieved using 0.05% trypsinEDTA. All the media and supplements for cell culture growing were purchased from Invitrogen (USA). Preparation and Analysis of CDDP3. cis-Diamminedinitratotplatinum(II), CDDP3, was synthesized using procedures modified from refs 14 and 15. Briefly, aqueous solution of cisplatin (2 mg/mL) was mixed with AgNO3 in the molar ratio of 1:2 and stirred overnight (room temperature (r.t.), dark). Then the reaction mixture was centrifuged and passed through paper and 0.2 μm syringe filters (PVDF, Fisher Scientific, USA) to remove AgCl precipitate. The filtrate was concentrated using a rotary evaporator (Heidolph Laborota 4000, Germany) and then air-dried to obtain powder of CDDP3 (solubility in water 26 ± 3 mg/mL as analyzed by XRF). Concentration of Pt(II) was analyzed by “dried drop” method using X-ray fluorescence analysis (XRF) (XRF spectrometer “REspect”, Russia) with addition of molybdenum as an inner standard. Obtained CDDP3 was analyzed by (i) Fourier transform infrared spectroscopy (FTIR) using PerkinElmer Spectrum One spectrometer (PerkinElmer, Inc., USA). Spectra were recorded from 520 to 4000 cm−1 in frustrated total internal reflection mode with Universal ATR Accessory console; (ii) thermal analysis using Diamond Pyris thermoanalytical system (PerkinElmer, Inc., USA), operating at heating rate of 10 °C/ min over the range of 40−500 °C; (iii) UV absorbance using SpectraMaxM5 microplate reader (Molecular Devices Co., USA). The UV spectra were recorded from 220 to 600 nm in PBS or water (2 mg/mL). Synthesis of Liposomes. The liposomes were prepared by emulsification of the lipid films consisting of PC, Chol, DSPEPEG2000, and either positively charged DDAB or negatively charged DPPG at different ratios (Table 1). Targeted 3713

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Molecular Pharmaceutics Table 1. Compositions of Synthesized Liposomesa

a

name

PC (%)

Chol (%)

DSPE-PEG2000 (%)

DDAB (%)

Lip(+) Lip Lip(−)

56.0 62.2 43.5

29.7 33.0 23.1

4.4 4.8 3.3

9.9

The specific binding of VEGF- and VEGFR2-targeted liposomes was analyzed by standard protocol of indirect enzyme-linked immunosorbent assay (ELISA). Briefly, VEGFor VEGFR2-coated plates were incubated with the samples (targeted and nontargeted liposomes, free IgG and free specific mAbs) for 2 h at 37 °C; washed with PBS and additionally incubated with secondary goat antimouse IgG (Fab specific)peroxidase antibody (1 h, 37 °C). The reaction was completed by incubation with 3,3′,5,5′-tetramethylbenzidine (TMB, BD Biosciences, USA) and stopped by 3 N HCl. Moreover, the specific binding of VEGF- and VEGFR2-targeted liposomes was also confirmed by inhibition ELISA. In the case of VEGFtargeted liposomes, VEGF-coated plates were initially incubated with 5-fold excess of polyclonal antibodies to VEGF (1 h, 37 °C) prior to incubation with the samples. In the case of VEGFR2-targeted liposomes, samples were preincubated with VEGFR2 antigen. After that, the standard protocol of indirect ELISA was followed. Drug Release from Liposomes. The rate of drug release from liposomes was studied by dialysis method using semipermeable membranes (MWCO 3.5 kDa, Spectra/Por Float-A-Lyzer G2, Spectrum Laboratories Inc., USA). Purified drug-loaded liposomes (1 mL) were placed into capsules and dialyzed against 15 mL of the corresponding buffer: PBS (pH 7.4) or acetate buffer (ABS, pH 5.5 with 0.9% NaCl) under continuous shaking at 25 and 37 °C in dark. The concentration of released drug in samples was determined by XRF. The amount of drug released from liposomes was expressed as a percentage of the total drug and plotted as a function of time. Stability of Liposomes. The dispersion stability of empty and drug-loaded liposomes was studied by measurement of the size and PDI of samples during incubation at 4 and 37 °C in PBS for at least 3 weeks. Sedimentation stability of liposomes was visually evaluated in blood serum (10% v/v, 37 °C, 10 days). Storage stability was evaluated by DLS (particle size), XRF (drug concentration), and ELISA (immunochemical stability) after storage in PBS at 4 °C for 1 month. In Vitro Cytotoxicity Assay. Cytotoxicity of free drugs, empty, and drug-loaded liposomes was studied on glioma C6 or HEK293 cells using MTS assay (Promega, USA). Samples of CDDP and CDDP3 were dissolved (immediately, 1 day or 1 week before treating cells at concentrations up to 150 μg Pt/ mL) in water, PBS, or NaCl (150 mM). Drug-loaded liposomes were dissolved in water or NaCl one day before treating cells (500 μg Pt/mL). Cells (5000 cells/well) were seeded in 96-well plates in DMEM for 48 h before experiments (37 °C, 5% CO2). Cells were treated with various samples for 24 h, and then cultured for additional 24 h in fresh media at 37 °C. Cytotoxicity was determined by colorimetric MTS assay, and absorbance at 490 nm was detected by microplate reader Victor X-3 (PerkinElmer, USA). Untreated cells were taken as a positive control (100% viability) and medium without cells was used as a blank. The cell viability (%) was calculated as (Asample/ Acontrol) × 100. All experiments were repeated four times. Flow Cytometry. Flow cytometry analysis of VEGF- and VEGFR2-targeted liposomes was performed using two protocols. Nonspecific IgG-liposomes and liposomes without targeting moiety were used as controls. First, glioma cells (C6 or U-87 MG) were grown in 24-well plates (105 cells/well, DMEM with 5% FBS) for 24 h. Then, DiI-labeled liposomes were incubated with cells (3 h, 37 °C) at concentration 50 μg of liposomes/mL. Afterward, cells were washed, dissociated by 0.05% trypsin-EDTA, centrifuged (1500 rpm, 5 min), and

DPPG (%)

30

The amounts of lipids are in a molar ratio.

liposomes were prepared using DSPE-PEG2000-maleimid (0.8% molar). Lipids were dissolved in chloroform/methanol (3:1) solution followed by evaporation under reduced pressure, and the traces of solvents were removed by freeze-drying of lipid films. The films were hydrated in an aqueous solution of CDDP (2 mg/mL) or CDDP3 (15 mg/mL) and sonicated for 10 min (G112SP1T, Laboratory Supplies Company, USA). Then the liposome suspensions were extruded through polycarbonate membranes with reducing pore size of 400, 200, and 100 nm (Avanti Polar Lipids Inc., USA). Extruded liposome suspensions were purified using NAP-10 desalting columns (Sephadex G-25, GE Healthcare, USA) to remove free drug. The obtained liposomes were concentrated by ultrafiltration using Amicon filters (100 kDa), sterilized using 0.22 μm sterile filters (Fisherbrand, Fisher Scientific, USA), and stored at 4 °C until further use. Synthesis of Liposomes Conjugated with mAbs. Liposomes that contained DSPE-PEG2000-maleimide lipids were conjugated with monoclonal antibodies (mAb) bearing sulfhydryl groups. We used murine anti-VEGF and antiVEGFR2 mAbs as targeting groups and nonspecific immunoglobulin G (IgG) as a negative control. The mAbs (4 mg/mL, 0.1 M borate buffer, 5 mM EDTA) were incubated with 10-fold molar excess of 2-iminothiolane at r.t. for 45 min. Then activated mAbs were separated from byproducts using NAP-10 desalting columns pre-equilibrated with PBS, 2 mM EDTA. The thiolated mAbs were immediately conjugated with freshly prepared liposomes at ratio 300 μg of mAb per 10 μmol of PC (PBS, pH 7.4, 4 °C, overnight). The resulting mAb-conjugated liposomes were separated from unbound mAbs and byproducts by gel filtration chromatography (Sepharose CL-6B, PBS, column 50 × 2.5 cm, 0.5 mL/min). The obtained liposomes were concentrated and sterilized using 0.22 μm sterile filters. Physicochemical Characterization of Liposomes. The particle sizes, polydispersity indexes (PDI), and zeta-potentials (ζ) of obtained liposomes (with and without drugs) were determined by dynamic light scattering (DLS) using Zetasizer Nano ZS ZEN 3500 instrument (Malvern Instruments Ltd., UK) at 25 °C in water or PBS. Experimental values were the average of at least three different formulations. The liposomes were examined by transmission electron microscopy (TEM) to characterize their morphology and dimensions. A drop of liposomes was applied to a carbon filmcovered copper grid to form a thin film specimen, which was subsequently stained with 1% phosphotungstic acid. The samples were then examined with a Jeol JEM-1230 transmission electron microscope (TEM, JEOL Ltd., Japan) at an accelerating voltage of 200 kV. Drug loading capacity (LC, %) was calculated as percent ratio of mass of the incorporated drug to total mass of drugloaded liposomes. Encapsulation efficiency (EE, %) was determined as percent ratio of the incorporated drug to total drug added upon loading. 3714

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Molecular Pharmaceutics Table 2. Physicochemical Characteristics of Empty and Drug-Loaded Liposomesa PBS buffer, pH 7.4b

H2O, pH 7.4 sample without drug Lip(+) Lip Lip(−) with CDDP Lip(+) Lip Lip(−) with CDDP3 Lip(+) Lip Lip(−)

particle size (nm)

PDI

ζ (mV)

particle size (nm)

PDI

ζ (mV)

145 ± 10 135 ± 11 154 ± 13

0.11 ± 0.02 0.10 ± 0.02 0.09 ± 0.03

+38 ± 4 −23 ± 4 −40 ± 5

134 ± 9 125 ± 11 136 ± 12

0.06 ± 0.04 0.05 ± 0.04 0.07 ± 0.03

+5.7 ± 0.3 −1.7 ± 0.3 −4.6 ± 0.7

151 ± 9 133 ± 10 157 ± 11

0.09 ± 0.02 0.08 ± 0.02 0.07 ± 0.03

+40 ± 4 −29 ± 4 −40 ± 4

129 ± 9 116 ± 9 128 ± 10

0.06 ± 0.04 0.06 ± 0.03 0.06 ± 0.03

+4.3 ± 0.3 −1.8 ± 0.2 −9.4 ± 1.3

2.5 ± 0.3 3.6 ± 0.4 8.0 ± 0.7

146 ± 10 143 ± 12 158 ± 13

0.06 ± 0.02 0.05 ± 0.03 0.05 ± 0.03

+43 ± 4 −26 ± 4 −39 ± 5

133 ± 8 126 ± 10 136 ± 11

0.06 ± 0.04 0.06 ± 0.04 0.08 ± 0.03

+7.3 ± 0.3 −1.6 ± 0.3 −7.6 ± 1.1

4.9 ± 0.8 16.5 ± 1.2 27.6 ± 3

LC (w/w %)

Experiments were performed in triplicate, and data are expressed as means ± SD of at least three independent experiments. Abbreviations: polydispersity index (PDI), zeta-potential (ζ), and loading capacity (LC, w/w %). b10 mM phosphate buffer, 0.15 M NaCl, pH 7.4.

a

mg Pt/kg (n = 5). Blood was collected at 5 min and 1, 4, 7, 24, 48 h after i.v. injection and kept at −80 °C for further analysis. Platinum Determination in Biological Samples. The platinum concentration in blood samples was determined using Varian ion coupled plasma-mass spectrometer Agilent 7500c (ICP-MS, Agilent Technologies, Germany) with internal indium (In) correction.24 Calibration solutions with concentrations of Pt(II) (1, 5, 50, 100, and 250 ng/mL) were prepared by dilution of a multielement standard solution ICP-MS-68A-C. The calibration curve for Pt(II) was linear over the range of 0− 250 ng/mL. The within and between day variability were within 5% and 10%. Blood samples were diluted twice by concentrated nitric acid (stirring at 60 °C, 12 h). An internal In standard was added prior to digestion (400 ng/mL). This assay did not differentiate liposomal and free platinum as well as proteinbound platinum. Statistical Analysis. Statistical analysis was performed using Student t test by Microsoft Office Excel 2007 program, GraphPad Prism and R1. Differences were considered statistically significant when p < 0.05.

resuspended in PBS. In the other variant, cells were grown in 6well plates, washed and dissociated by 0.05% trypsin-EDTA. Afterward, cell suspensions were treated by the same liposomal formulations and thoroughly washed with PBS. Cells prepared using both protocols were analyzed by MoFlo XDP cell sorter (Beckman Coulter, Miami, FL, USA). Dead cells and cell debris were discriminated using forward scatter/side scatter characteristics (FSC/SSC). All data were analyzed using Summit V5.2.0.7477 (Beckman Coulter, USA). Confocal Microscopy. Immunofluorescent analysis of obtained liposomes was performed on fixed and live glioma cells (C6, U-87 MG). The first setup: cells were fixed (4% paraformaldehyde, pH 7.4) and incubated with PBST buffer (PBS, 0.2% Tween 20, 0.2% Triton X-100) for 2 h at r.t. for better membrane permeabilization. Afterward, cells were washed (12 h, 4 °C, PBS) and incubated with VEGF- and VEGFR2-targeted liposomes at concentration 100 μg liposomes/mL at r.t. for 2 h. In the second setup live cells were incubated with liposomes for 3 h at 37 °C prior to fixing. For colocalization and dynamic analysis we used 35 mm confocal Petri-dishes (SPL Lifesciences, South Korea). Initially, live glioma cells (5000 cells/dish) were treated with LysoTracker Red DND-99 (Molecular Probes, USA). Then cells were incubated with targeted liposome and controls for 15, 30, 40, and 60 min at 37 °C. Then cells were thrice washed by PBS and analyzed using confocal (Nikon A1R MP+, Japan) and fluorescence microscopy (Leica DMI6000, Germany). Intracranial Glioma Rat Model. All studies on animals were approved by an Ethical Committee of the Russian National Research Medical University, named after N.I. Pirogov. Glioblastoma multiforme was modeled by stereotaxic injection of 4 × 105 glioma C6 cells into the right striatum of 12-week female Wistar rats under ketamine anesthesia as described previously.23 Twenty days after implantation of C6 cells the tumor growth and its location were verified by T2weighted MRI (ClinScan 7T, Bruker, Germany), and animals were randomized based on the tumor area. Pharmacokinetics of Liposomes in Rats with Glioma C6. Glioma C6 bearing rats received either CDDP3 or liposomal formulations of the drug (liposome/CDDP3, IgGliposome/CDDP3, and VEGF-targeted liposome/CDDP3) by single femoral vein injection. Doses were based on the individual animal body weight and were equivalent to a 4.5

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RESULTS Synthesis of Highly Soluble Cisplatin Analogue. In this study, we used two platinum-based derivatives: CDDP and its highly soluble analogue, CDDP3, prepared as described in refs 14 and 15 with slight modifications. FTIR spectroscopy and elemental analysis evinced that composition of crystalline CDDP3 was cis-[Pt(NH3)2(NO3)2] (Figure S1, Tables S1 and S2). Table S1 shows that the main vibration bands coincide with those of cis-[Pt(NH3)2(NO3)2] whose structure was independently proved by single-crystal X-ray diffraction.25 This structure was also proved by a combination of thermogravimetric analysis and differential scanning calorimetry (Figure S2) that revealed an endothermic peak at 200−290 °C accompanied by decrease in mass to 62 ± 3% of initial mass. Theoretical loss for conversion of [Pt(NH3)2(NO3)2] into PtO2 upon heating is 64.3%. Further, UV spectra of CDDP3 in water and NaCl (Figure S3) revealed that CDDP3 converts in pure water into an ionic and thus highly soluble cis[Pt(NH3)2(H2O)2](NO3)2 compound. However, it easily converts into CDDP upon addition of Cl− ions or direct dissolution in saline media. Finally, we have found that CDDP3 (as a free salt or released from liposomes) reveals a positive reaction in Kurnakow’s test,26 thus implying its conversion into 3715

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Molecular Pharmaceutics

Figure 1. Characteristics of drug-loaded liposomes. (A) TEM microphotograph of CDDP-loaded negatively charged liposomes. Scale bar: 2 μm. (B) Drug release profiles for CDDP (squares) or CDDP3 (triangles) encapsulated in liposomes (Lip, black) or negatively charged liposomes (Lip(−), blue) in PBS buffer (pH 7.4, 25 °C). (C) Dependences of particle size and polydispersity (PDI) of CDDP-loaded liposomes on time in PBS buffer at 37 °C. (D) ELISA immunoassay of antibody-conjugated liposomes. Binding of free anti-VEGF mAbs, specific anti-VEGF-Lip(−), nonspecific IgGLip(−), and nontargeted Lip(−) as well as competitive inhibition of anti-VEGF-Lip(−) by preincubation of the antigen with anti-VEGF pAbs.

cis-isomer, which is an active form of the drug. Thus, a combination of physicochemical data unambiguously shows that we have synthesized a CDDP analogue that possesses better solubility in water (14.8 mg Pt/mL compared to 1.9 mg Pt/mL for CDDP) and easily converts into CDDP in saline media. Synthesis and Characterization of Drug-Loaded Liposomes. Next, we sought to design the stable cisplatin-loaded liposomes with high loading capacity. First, we varied the lipid composition of liposomes (Table 1) to detect its influence on their stability, size, and surface charge. Liposomes were prepared using the film hydration method.27 All liposomes had PEG-corona, created via admixing of 4−5 mol % DSPEPEG2000. PEGylated liposomes of different lipid composition demonstrated similar particle size and relatively low polydispersity indexes (PDI) (Table 2) and possessed spherical morphology (Figures 1A and S4). As ionic strength increased (water vs PBS), the size of the liposomes decreased by about 10−15%, but their PDI did not change significantly. The ζpotential of the liposomes was determined by the presence of charged components (liposomes with positive DDAB or negative DPPG or with DSPE-PEG only were designated as Lip(+), Lip(−), or Lip, respectively). The ζ-potential of the liposomes drastically changed with increase of ionic strength (Table 2): net charge expectedly decreased due to enhanced counterion screening. Second, we analyzed the influence of various parameters (such as lipid composition, loaded drugs, and [lipids]/[drug] molar ratio) on loading capacity (LC), encapsulation efficiency (EE), and cytotoxicity of resulting liposomes. Though a broad range of molar ratios of charged components (e.g., 1, 5, 10, 25% for DDAB and 5, 15, 30% for DPPG) was analyzed, Table 1

presents the most suitable formulations in terms of LC and cytotoxicity. For example, Lip(−) and Lip(+) had very low LC (less than 5 and 1.5%, respectively) if they contained less than 30 mol % DPPG and 10 mol % DDAB. Moreover, Lip(+) was very cytotoxic if it contained more than 10 mol % DDAB (Figure S5A). We found that the decrease in ζ-potential of liposomes led to the expected increase in the drug content (Table 2): liposome formation is performed in aqueous saltfree media, where platinum exists in the form of cis[Pt(NH3)2(H2O)2]+2 cation, which reveals increased binding to anionic DPPG at these conditions. Also the concentration of the drug solution strongly affected the LC. CDDP is a poorly soluble salt with solubility in water of about 2 mg/mL; in contrast, CDDP3 has solubility that is 10 times higher than CDDP. For example, LC of Lip increased by 4.5 times when CDDP3 was used instead of CDDP (Table 2). We did not observe a considerable difference in LC with the change of [lipids]/[drug] molar ratio, although this parameter had a profound effect on the encapsulation efficacy of liposomes. We found that EE increased with an increasing lipid content at a constant amount of drug (e.g., for Lip/CDDP, EE was 6.3% and 30.2% at [lipids]/[drug] molar ratios equal to 1:1 and 6:1, respectively). Moreover, EE was more than 2-fold higher for Lip(−) than for Lip at the same [lipids]/[drug] molar ratio (16.8% vs 37.8% for Lip/CDDP and Lip(−)/CDDP, respectively). However, EE did not depend much on the concentration of loaded platinum salt. In addition, we found that the internal structure of the liposomes depended on the lipid composition: Lip(−) possessed unilamellar structure; in contrast, Lip was multilamellar (Figures 1A and S4). Under similar particle size the free internal volume of Lip was smaller, which correlates with a decrease in their LC. The loaded drugs 3716

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Molecular Pharmaceutics did not change significantly the physicochemical characteristics of unloaded liposomes either in water or PBS buffer (Table 2). Ultimately, the highest LC (27.6 ± 3%) and EE (37 ± 2.5%) were achieved via a combination of negative liposomes Lip(−) and CDDP3 at high [lipids]/[drug] ratio. Drug Release and Stability of Drug-Loaded Liposomes. The release of CDDP and CDDP3 from liposomes was analyzed by a dialysis method in PBS (pH 7.4) and ABS (pH 5.5) buffers. All formulations displayed a sustained drug release profile without initial burst release (Figure 1B). We found that release rate depended on the lipid composition of liposomes, pH of buffer (7.4 or 5.5), and drug type (CDDP or CDDP3). First, the release slowed down in the presence of negative lipids: cumulative drug release was 20 ± 5% less in Lip(−) than in Lip. Second, drug release was slightly faster (about 10−15%) in acidic conditions (pH 5.5) rather than at pH 7.4 during the entire studied period (Figure S6). Third, CDDP3 released slower than CDDP from all types of liposomes: 31 ± 3% of CDDP3 and 44 ± 2% of CDDP was released from negative liposomes during the first 24 h. Dispersion stability of liposomes (empty and loaded with drugs) was assessed by DLS analysis of samples, which were incubated in PBS at 4 and 37 °C for several weeks (Figures 1C and S7). Liposomes did not form either aggregates or precipitates in PBS buffer or in the blood serum (10% v/v), indicating their high dispersion stability. It should be noted that during the storage at 4 °C for 6 weeks we observed the leakage of the drug, which depended on the concentration of liposomes, their charge, and type of buffer. Thus, the drug leaked faster in the presence of salt and at low concentration of liposomes in the mixture. The leakage was higher for Lip than for Lip(−) in water (30% vs 20%), probably due to the electrostatic interaction between drugs and lipids. In the case of storage in PBS, the leakage was the same for both types of liposomes (about 50%, 2 months). Note that in the first week of storage the drug leakage was less than 6 ± 2%. These data demonstrate that synthesized liposomes are stable and displayed sustained drug release. Synthesis and Characterization of Targeted Liposomes. We conjugated liposomes with specific monoclonal antibodies to improve their delivery in vivo to the targeted tissue. We used specific mAbs to VEGF and VEGFR2 and nonspecific IgG as a control. We synthesized liposomes with active outer maleimide groups by adding DSPE-PEG2000maleimide in the composition of liposomes (Scheme 1). We thiolated mAbs using 2-iminothiolane and then conjugated them with maleimide functionalized liposomes, followed by thoughtful purification by size-exclusion chromatography.28 The amount of conjugated mAbs to liposomes was ∼20−30 μg mAb/mg total lipids. Using ELISA we detected that activity of specific antibodies retained at the high level after conjugation (about 80% of the initial activity, Figures 1D and S8). A negligible signal was detected for liposomes conjugated with IgG due to nonspecific adsorption of IgG to the antigen. Comparison of the activities of drug-loaded and empty mAb-conjugated liposomes indicated that CDDP and CDDP3 did not alter the function of mAbs (Figures S8). Moreover, competitive inhibition analysis using ELISA showed that the specific binding of targeted liposomes to its targets drastically decreased after their preincubation with antigen or preincubation of antigen with anti-VEGF pAbs (Figure 1D). The immunochemical activity of the targeted

Scheme 1. Scheme of Conjugation of Antibodies to Liposomes

liposomes remained for at least 10 days at 4 °C, as detected by ELISA. Predictably, the particle size and ζ-potential of liposomes slightly increased after conjugation with mAbs (134 vs 141 nm and −23 vs −17 mV in water). We also detected further increase in size (10 ± 2%) after drug loading of targeted liposomes (Table S3). All targeted liposomes were stable and practically did not change their size within 10 days in PBS buffer. The drug release rate from targeted liposomes was slightly (but not significantly) slower compared to nontargeted liposomes: the difference was about 10% and less than 5% at 8 and 48 h, correspondingly. This suggests that targeting groups do not influence the drug retention. For further in vitro and in vivo analysis of liposomes, fluorescent label (DiI, 1 mol %) was added to the lipid composition. We did not detect any leakage of DiI from liposomes during incubation for a week in PBS buffer or 50% FBS (Figure S9). Collectively, these data demonstrate the successful synthesis of VEGF- and VEGFR2-targeted liposomes with high loading capacity and sustained drug release. Cellular Uptake of Targeted Liposomes. To analyze the influence of targeting group on the cellular uptake of liposomes we used rat glioma C6 and human glioma U-87 MG cell lines, known for high expression of VEGF and VEGFR2.18,19 Fluorescence microscopy revealed that targeted liposomes bound to cell-associated antigen, which resulted in their enhanced accumulation in fixed glioma C6 cells (Figure S10). Fluorescence analysis of live cells confirmed the enhanced accumulation of VEGF- and VEGFR2-targeted liposomes (Figure 2C,D,G,H). We also detected the minor retention of nontargeted liposomes and IgG-liposomes in cells under the same conditions (Figure 2A,B,E,F). The accumulation of IgGliposomes exceeded nontargeted liposomes, probably due to hydrophobic and ionic intermolecular interactions between IgG and cell membrane molecules. Moreover, analysis of dynamic internalization of liposomes in glioma C6 cells revealed that cellular uptake of VEGF- and VEGFR2-targeted liposomes was 1.5 times faster than other studied formulations. Thus, considerable accumulation of targeted liposomes was detected after 30 min of incubation, while IgG-liposomes and non3717

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Figure 2. Accumulation of DiI-labeled negatively charged liposomes (red) by C6 and U-87 MG glioma cells: nontargeted liposomes (A,E), IgGliposomes (B,F), VEGFR2-targeted (C,G), and VEGF-targeted liposomes (D,H). Nuclei were strained with DAPI (blue). Scale bars: 10 μm (A−D) and 50 μm (E−H).

Figure 3. Cellular uptake of DiI-labeled liposomes by a monolayer or a suspension (inserts) of glioma U-87 MG (A) and C6 (B) cell lines. Cells were exposed to negatively charged liposomes (nontargeted, IgG-liposomes, and VEGF- or VEGFR2-targeted liposomes) for 3 h. Data are mean ± SD (n = 3). Statistical significance: n.s., not significant; *p < 0.05; for all other pairs, p < 0.01 (not shown).

targeted liposomes were detected after 40 and 60 min, respectively. Analysis of colocalization of DiI-labeled liposomes with LysoTracker Red in live C6 cells showed that the targeted and nontargeted liposomes were predominantly accumulated in the lysosomes after 60 min of incubation, but to a different extent. Fluorescence intensity of accumulated VEGF- and VEGFR2targeted liposomes increased by 2.4 and 3.2 times compared to nontargeted liposomes, respectively (Figure S11). The enhanced accumulation of targeted liposomes was also confirmed by cytometry analysis. Cellular uptake was analyzed both in a monolayer and a suspension of glioma C6 and U-87 MG cells (Figure 3). The accumulation of VEGF-targeted liposomes exceeded that of nonspecific IgG-liposomes by 6 and 10 times in suspensions of glioma C6 and U-87 MG cells, respectively (inserts in Figure 3). The difference was less in the case of the monolayer, where cells were trypsinized after incubation with liposomes. For example, for U-87 MG cells, the ratio of cellular uptake of IgG-liposomes to VEGF-targeted liposomes was 1:10 in the suspension and 1:4 in the monolayer.

It could be due to the following reasons: (i) the rate of cellular uptake could be different in suspension and adherent cells;29 (ii) part of targeted liposomes attached to the cell surface did not internalize into cells and dissociated during trypsinization.30 Also, we noticed that difference between targeted and nontargeted liposomes was higher in U-87 MG cells compared to C6. This data indicates that, despite the strong membrane binding of targeted liposomes, their internalization could be low due to the difference in level of endocytosis in C6 and U-87 MG cells. As expected, cellular uptake was time and concentration dependent: uptake of liposomes increased by 2−2.5 times upon increasing incubation time from 1 to 3 h (Figure S12). The difference between VEGF-targeted and IgGliposomes reduced by 1.4−1.7 times with increasing the concentration from 20 to 40 μg/mL. These studies confirmed the enhanced membrane binding and internalization of targeted liposomes by glioma cells that could lead to their increased cytotoxicity. Cytotoxicity of Drug-Loaded Liposomes. Despite the low content of encapsulated drug in positively charged 3718

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promotes the cytotoxicity of drug-loaded liposomes in glioma cells. Based on these data we opted for the targeted negatively charged liposomes loaded by CDDP3 for further in vivo studies. Pharmacokinetics of CDDP3-Loaded Liposomes in Rats with Glioma C6. We analyzed pharmacokinetics of CDDP3-loaded liposomes in rats with intracranial glioma C6. The free drug and liposomal formulations at the same drug dose (4.5 mg Pt/kg) and control (PBS) were injected into the femoral vein of rats (5 groups, n = 5). ICP-MS analysis of collected whole blood demonstrated that liposomes remained in the blood circulation much longer than the free drug (Figure 5). Even at the earliest evaluated time point (5 min post

liposomes, we analyzed the cytotoxicity of all obtained formulations in glioma C6 cells using MTS assay. As expected, the cytotoxicity of empty positively charged liposomes increased with the increasing positive charge of the particles (e.g., formulation with 5 mol % DDAB was relatively nontoxic compared to 10 mol %). Under the same conditions, the Lip and Lip(−) did not reveal cytotoxicity (Figure S5B). Further, we compared the cytotoxicity of CDDP and obtained CDDP3 under various conditions (salts were dissolved in water, PBS, or NaCl). The MTS test revealed that both platinum-derivatives had a similar cytotoxicity, which varied depending on the solvent used (Figure S13). Thus, IC50 increased by about 1.5 times if we used PBS instead of water or NaCl (IC50 for CDDP3 was 9.2 ± 2 vs 6.4 ± 2 or 6 ± 2 μg Pt/ mL, respectively). We observed difference in IC50, if samples were dissolved immediately or a week before the experiment. IC50 decreased by 2-fold if samples were kept in NaCl (150 mM) probably due to preventing of aquation of platinum derivatives (Figure S13). MTS test performed on HEK293 cells detected a similar cytotoxicity of both platinum salts (IC50 1.7 and 1.9 μg Pt/mL), indicating the successful conversion of CDDP3 into CDDP. Cytotoxicity of CDDP and CDDP3 decreased 3−5 times upon their loading into liposomes compared to free drugs (Figure 4). Cytotoxicity of neutrally and negatively charged

Figure 5. Pharmacokinetic profile of CDDP3-loaded liposomes in rats with intracranial glioma C6. Platinum levels were measured following a single intravenous dose of CDDP3-loaded nontargeted liposomes, IgG-liposomes, and VEGF-targeted liposomes or free CDDP3 at 4.5 mg Pt/kg body weight. Data are mean ± SD (n = 5).

dosage), we detected 2.5−4 times more platinum concentration for liposomal formulations than for the free drug (Table 3). The clearance time values of IgG- and VEGF-targeted liposomes were similar, and both of them were removed from the bloodstream faster than nontargeted liposomes. Consistent with previous reports,4,24 free cisplatin disappeared rapidly from circulation with only about 8% of the injected dose detectable in plasma 5 min after dosing. In contrast, all liposomal formulations showed prolonged blood circulation with about 46% and 25% of the injected dose retained 5 min after i.v. injection of nontargeted liposomes and IgG- or VEGFtargeted liposomes, respectively. After 24 h the platinum concentration in the blood was about 2 μg/mL for the free drug, whereas it was about 8 μg/mL for liposomes (i.e., 10% of injected dose). AUC value was similar for IgG-liposomes and VEGF-targeted liposomes, while it was doubled for nontargeted liposomes and decreased by 2-fold for the free drug (Table 3). In the anti-VEGF-liposome treatment group, both Vss and Cl decreased by 4-fold and 2-fold, respectively, compared to the free drug treatment group.

Figure 4. Cytotoxicity of CDDP-loaded liposomes on glioma C6 cells after 24 h of incubation. Liposomes contained anionic lipid (Lip(−)) or only DSPE-PEG (Lip). IC50 values of nontargeted liposomes, nonspecific IgG-liposomes, and VEGFR2-targeted liposomes. Data are mean ± SD (n = 3). Statistical significance: n.s., not significant; *p < 0.05.

liposomes loaded with the same drug was similar (e.g., 21.4 ± 2 and 24.3 ± 2 μg Pt/mL for CDDP). However, drug-loaded positive liposomes were much more toxic (6 μg Pt/mL) because of the lipid composition. If we take into account the loading capacity of negatively charged liposomes loaded with CDDP3, then their cytotoxicity improved by 9 times compared to conventional neutral liposomes loaded with CDDP (173 vs 1544 μg lipids/mL). Finally, we analyzed how targeting groups influenced the cytotoxicity of drug-loaded liposomes by studying VEGFR2targeted liposomes (24 h incubation, C6 cells). We found that the cytotoxicity of VEGFR2-targeted liposomes/CDDP doubled compared to nontargeted or IgG-targeted liposomes/ CDDP (Figure 4). These results indicate that targeting group

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DISCUSSION In the present study we developed a novel VEGF- and VEGFR2-targeted liposomal cisplatin formulation with improved loading capacity and analyzed its accumulation in glioma cells and pharmacokinetics in rats bearing intracranial glioma C6. This formulation was designed to overcome poor pharmacokinetics typical for CDDP31 and low loading capacity (LC) typical for conventional CDDP-loaded liposomes.9 Insufficient LC of liposome-based CDDP formulations stems from low water solubility of cisplatin. This problem has been partially addressed via various approaches such as use of active loading instead of passive entrapment of drugs, optimization of 3719

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Table 3. Pharmacokinetic Parameter Estimates of Platinum in Plasma for Glioma C6 Bearing Rats Treated with CDDP3 or CDDP3-Loaded Liposomes: Nontargeted Lip(−), Nonspecific IgG-Lip(−), and Targeted Anti-VEGF-Lip(−)a Parameterb Cmax (μg/mL) AUC(μg·h/mL) Cl (mL/h·kg) Vss (L/kg) t1/2 (h) MRT (h)

CDDP3 7.9 98 0.017 1.96 71 102

± ± ± ± ± ±

1.4 7 0.003 0.13 11 15

Lip(−)/CDDP3 39.5 504 0.008 0.23 20 29

± ± ± ± ± ±

7.0 28 0.002 0.02 1 1

IgG-Lip(−)/CDDP3 19.1 230 0.009 0.49 19 27

± ± ± ± ± ±

3.5 31 0.003 0.06 3 4

anti-VEGF-Lip(−)/CDDP3 17.5 203 0.008 0.51 17 24

± ± ± ± ± ±

1.0 11 0.002 0.02 1 2

a The pharmacokinetic parameters were calculated using a noncompartment model. bCmax represents the maximum observed concentration at the earliest evaluated time point of 5 min postdosage. Abbreviations: Cmax, maximum platinum concentration; AUC, area under the curve; Cl, clearance; Vss, steady-state volume of distribution; t1/2, noncompartmental half-life; MRT, mean residence time.

lipid composition, etc.9,11 The most successful approach for LC to increase the cargo solubility, which can be achieved either by using a combination of different solvents or using more soluble platinum drugs (e.g., use of oxaliplatin in Lipoxal, Regulon, Inc.9,11) or highly soluble CDDP analogues. In the last case, CDDP3, which appears in water as [Pt(NH3)2(H2O)2](NO3)2 and readily converts into CDDP in saline media,14 has been successfully used for LC increase in both liposomes14 and lipid nanoparticles.15,32 Based on the above literature data, we have chosen passive CDDP3 loading into liposomes via the film hydration method. Liposomal formulations (Table 1) were designed to make lipid composition as close as possible to PEGylated SPI-77 (51 mol % of hydrogenated PC; 44 mol % of Chol, and 5 mol % DSPEPEG2000)13 but contained additional ionic component (anionic DPPG, as in Lipoplatin,12 or cationic DDAB). Prepared CDDP3 was characterized using a combination of physicochemical methods. We used CDDP-loaded liposomes as a control. Variation of lipid composition of liposomes (Table 1) revealed that CDDP3 accumulated more efficiently in negatively charged liposomes (most likely because of electrostatic interactions of [Pt(NH3)2(H2O)2]2+ cations with anionic lipids). Ultimately, the optimized composition (namely, negatively charged liposomes with 30% DPPG loaded by CDDP3) revealed LC ca. 30 wt %, which is superior compared to analogues in the literature, e.g., SPI-77 by 5-fold,11 Lipoplatin by 3-fold,12 CDDP-loaded liposomes conjugated with Sialyl Lewis by 1.4-fold,14 and PLGA-PEG nanoparticles by about 8-fold.33 Though Guo and colleagues designed lipid nanoparticles with LC of about 80%,32 these particles demonstrated quite fast drug release: 50% of Pt released in first 3 h. In our case, probably it is the combination of electrostatic interactions and high water solubility of the drug that provides such a profound increase in liposomal LC. This suggestion is supported by the comparison of our formulation to SPI-77 since its composition stems from that of SPI-77. Based on our data it can be seen that addition of anionic component (the same lipid, DPPG, as used in Lipoplatin) and use of more soluble CDDP analogue provides 5-fold increase of LC. All other parameters of obtained liposomes were determined by the preparation process: liposomes possessed spherical morphology with diameters dictated by the pore diameter of membrane filters; ζ-potential was determined by lipid composition and ionic strength in an expected manner. Drug release was sustained and strongly depended on the lipid composition. The cumulative release of both drugs from the neutral liposomes was slower by 20% than that from negatively charged liposomes. The release rate of CDDP was marginally faster than that of CDDP3, and both of them slightly

increased in acidic media (pH 5.5) compared to neutral pH. A possible explanation of these data is the difference between CDDP and CDDP3: the former is a neutral molecule and should easily pass through the nonpolar inner part of a lipid membrane, while the latter is a dissociated salt and its release from the liposome would be hampered by the nonpolar part of a membrane. The dependence of drug release on pH and lipid composition can be explained by partial aquation of CDDP and/or conversion of both drugs into the same form (which is the equilibrium) at the same external conditions. The last argument also provides the reason why CDDP3 may not be converted into CDDP at the formulation stage (compare with refs 14,15). Taking into account fast interconversion rates of CDDP and CDDP3, one can anticipate fast equilibration of both formulations in the bloodstream. Another challenge addressed in this work was the improvement of the pharmacokinetics of CDDP. We sought to solve this problem by preparing long-circulating targeted liposomes. For this purpose we (i) decorated all liposomes with PEG chains capped by maleimide groups and (ii) conjugated them with monoclonal antibodies to VEGF or VEGFR2. Conjugation of liposomes with mAbs did not alter their stability, although predictably increased size of liposomes by about 10 nm and increased ζ-potential by about 8 mV as a result of positive charges of mAbs. Hence, we obtained stable targeted liposomes, which retained high specific activity of mAbs (>80%). A competitive inhibition assay corroborated the binding specificity of anti-VEGF-liposomes and anti-VEGFR2liposomes to their antigens. After physicochemical characterization of obtained liposomes, we analyzed how targeting groups influence the cellular uptake of liposomes by glioma cells. We observed enhanced accumulation of VEGF-targeted liposomes in U-87 MG and C6 glioma cells by confocal and FACS analyses, leading to improved cytotoxicity compared to nontargeted and IgG-liposomes. This result is in line with previously reported improvement of cellular uptake and cytotoxicity of VEGFtargeted cationic liposomes on different VEGF-positive cell lines34 and VEGFR2-targeted liposomes on leukemic cell culture.20 In contrast, if transferrin ligand was used, then the difference between targeted and nontargeted liposomes in the cellular uptake by C6 cells was significantly lower.35 In accordance with our data, Xiang et al., using chlorotoxincoupled liposome, demonstrated that the difference between targeted and nontargeted liposomes was significantly lower in rat glioma C6 compared to human glioma U-87 and U-251 cells.36 This effect could be due to different level of receptors expressed in particular cell lines and variations in activity of endocytosis from line to line. We found that cytotoxicity of the 3720

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result of differences in neovasculature of normal and tumor tissue.10 Probably, the accumulation and distribution of nanoparticles in the tumor is determined primarily by passive transport (EPR effect), while the targeting plays a significant role within the tumor.41 Thus, many therapeutic agents (RNA, cisplatin, doxorubicin) require cellular internalization for their action. Their high accumulation within the tumor tissue (not accompanied by cellular uptake) may not induce the antitumor activity.42 However, the targeting moieties can facilitate the cellular uptake of drug-loaded nanoparticles by the tumor cells and thus improve the therapeutic efficacy. We believe that the synergistic effect of nanoparticle distribution in the body and targeting approach can provide the enhanced delivery of liposomal cisplatin to VEGF- and VEGFR2-positive tumors. Taking into account high loading capacity of the developed liposomal formulation and its high efficacy in vitro, we believe that there are many reasons to anticipate improved accumulation of this formulation in in vivo glioma models.

drugs was significantly decreased after their encapsulation into obtained liposomes. A similar effect has been detected for other anticancer drugs loaded in nanoparticles.37 However, some publications stated the increase of CDDP cytotoxicity after its encapsulation into liposomes16 or lipid nanoparticles.15,32 We did not detect this behavior, probably due to slower release of the drug from our formulations compared to the reported nanoparticles, difference in ζ-potential (−7.6 mV compared to +15 mV15,32), or different cell lines used. Taking into account the high LC of liposomes loaded with CDDP3, their use might lead to lowering of the administered dose. Summarizing in vitro data, we can conclude that VEGF- and VEGFR2-targeted anionic liposomes loaded with CDDP3 represent a reliable platform for targeted drug delivery. Encapsulation of antitumor agents into liposomes changes not only their toxicity, but also their pharmacokinetics. Pharmacokinetics is mainly determined by the structural parameters of liposomes. The major impact to prolonged circulation is provided by PEG corona, its thickness and density; lower but considerable elongation of circulation time can be achieved via variation of liposome size27 or lipid composition.11 Our data confirmed previous findings that PEGylated liposomes considerably extend the blood circulation time of free drugs as a result of formation of protective layer that slows down recognition of liposome by opsonins.13 The liposomes had a more than 5-fold increase of AUC and 8-fold decrease of steady-state volume of distribution (Vss) compared to cisplatin, indicating that liposomes/CDDP3 are better retained in the blood. Since the loaded drug was released gradually with time, we did not observe any spikes in platinum concentrations in the blood over a 48 h period. This might help to avoid the initial peak platinum concentration in the kidneys that would produce a severe toxicological effect.9 In addition to the structural parameters of liposomes, conjugation with targeting group, such as monoclonal antibodies, could significantly alter liposomal concentration in the blood.38,39 Indeed, in our work we observed that the decoration of the PEGylated liposome surface by antibodies led to a substantial decrease of the liposome blood level. The IgG- and VEGF-targeted liposomes were removed from the bloodstream faster than nontargeted ones. This can be due to the enhanced removal of mAb-targeted liposomes by the mononuclear phagocyte system (MPS) because of the interaction of the Fc region of the whole mAb with the Fcγ receptors on immune cells.38 It was shown that tumors are able to sensitize the immune system that could lead to increase the liposome elimination by MPS.39 Moreover, the number of mAbs per liposome, their density, and their nature (whole mAb vs Fab) may accelerate the clearance.38,40 Thus, coating of liposomes with antibodies may enhance the tumor targeting, but could compromise the stealth effect of PEG shielding. Nevertheless, all studied liposomes prolonged the blood circulation time of cisplatin, although the circulation time highly depended on the liposomal surface. Summarizing the data discussed above, we can conclude that the developed liposomal formulation features several advantages. First, our in vitro data confirm that conjugation of specific antibodies to liposomes increases the intracellular concentration of the liposomes and improves cytotoxicity of the retained drug compared to nontargeted liposomes. Second, liposomes are stable in bloodstream and prolong the circulation time of the drug. This can lead to the improved drug accumulation in the tumor due to the EPR effect, which is a

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CONCLUSIONS In conclusion, we have developed the drug delivery system based on liposomes that combines enhanced drug loading capacity and high binding affinity to glioma cells. To achieve high loading capacity we used highly soluble cisplatin analogue. For targeted delivery of cisplatin to intracranial gliomas we used liposomes conjugated with anti-VEGF- and anti-VEGFR2 antibodies. The above features enable overcoming of typical drawbacks of conventional cisplatin formulations, including low loading capacity, poor pharmacokinetics, and insufficient accumulation in the targeted tumor tissue. The developed drug delivery system revealed sustained drug release profile, high uptake by glioma C6 and U-87 MG cells, and prolonged blood circulation time. Further studies will focus on the therapeutic outcomes of cisplatin-loaded liposomes since targeting can increase the internalization by the tumor cells and improve the therapeutic efficacy.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.molpharmaceut.6b00519. Detailed characterization of CDDP and CDDP3 as well as in vitro evaluation of liposomes (PDF)

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AUTHOR INFORMATION

Corresponding Author

*Tel: 617-335-9001. E-mail: [email protected]. Present Address

Koch Institute for Integrative Cancer Research, MIT, 500 Main Street, Cambridge, Massachusetts 02139, United States. Author Contributions ◆

S.A.S. and I.I.K. contributed equally to this work. The manuscript was written through contributions of all authors. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by grant 14-15-00698 (all studies except TEM) from Russian Science Foundation (to N.V.N.) and grant 11.G34.31.0004 (TEM imaging) from Russian 3721

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Federation Ministry of Science and Education (to A.V.K.). We thank N. F. Grinenko for her help with cytotoxicity assay and S. Cherepanov for acquisition of cytometry data.

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