Self-Regulated Multifunctional Collaboration of Targeted Nanocarriers

Aug 22, 2014 - Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Institute .... Herein, to achieve multifunctional collab...
1 downloads 0 Views 6MB Size
Download PDF
Article pubs.acs.org/Biomac

Self-Regulated Multifunctional Collaboration of Targeted Nanocarriers for Enhanced Tumor Therapy Hongjun Gao,†,§ Tangjian Cheng,† Jianfeng Liu,∥ Jinjian Liu,∥ Cuihong Yang,∥ Liping Chu,∥ Yumin Zhang,∥ Rujiang Ma,† and Linqi Shi*,† †

State Key Laboratory of Medicinal Chemical Biology, Key Laboratory of Functional Polymer Materials, Ministry of Education, Collaborative Innovation Center of Chemical Science and Engineering (Tianjin), Institute of Polymer Chemistry, Nankai University, Tianjin 300071, People’s Republic of China § Kingfa Science & Technology Company, Ltd., Guangzhou 510520, People’s Republic of China ∥ Tianjin Key Laboratory of Radiation Medicine and Molecular Nuclear Medicine, Institute of Radiation Medicine, Chinese Academy of Medical Science & Peking Union Medical College, Tianjin 300192, People’s Republic of China S Supporting Information *

ABSTRACT: Exploring ideal nanocarriers for drug delivery systems has encountered unavoidable hurdles, especially the conflict between enhanced cellular uptake and prolonged blood circulation, which have determined the final efficacy of cancer therapy. Here, based on controlled self-assembly, surface structure variation in response to external environment was constructed toward overcoming the conflict. A novel micelle with mixed shell of hydrophilic poly(ethylene glycol) PEG and pH responsive hydrophobic poly(β-amino ester) (PAE) was designed through the self-assembly of diblock amphiphilic copolymers. To avoid the accelerated clearance from blood circulation caused by the surface exposed targeting group c(RGDfK), here c(RGDfK) was conjugated to the hydrophobic PAE and hidden in the shell of PEG at pH 7.4. At tumor pH, charge conversion occurred, and c(RGDfK) stretched out of the shell, leading to facilitated cellular internalization according to the HepG2 cell uptake experiments. Meanwhile, the heterogeneous surface structure endowed the micelle with prolonged blood circulation. With the self-regulated multifunctional collaborated properties of enhanced cellular uptake and prolonged blood circulation, successful inhibition of tumor growth was achieved from the demonstration in a tumor-bearing mice model. This novel nanocarrier could be a promising candidate in future clinical experiments.

■

immune responses,8−11 resulting in unsatisfactory efficacy. To improve the efficacy of tumor therapy and advance the development of drug delivery, new strategies are urgently needed to overcome the conflict of prolonged circulation and enhanced cellular uptake. The in vivo barriers for systemic administered nanoparticles (NPs) include the transport via blood to the tumor extracellular matrix, binding to the cell membrane, cellular internalization, and intracellular delivery.12,13 NPs were generally designed with stealthlike features, protective layers, targeting moieties, membrane-penetration moieties, sensors to trigger drug release, and other functionalities. Nanotechnology-based nanomedicines rely largely on the enhanced permeability and retention (EPR) effect, due to the defective lymphatic drainage in solid tumor and leaky tumor blood vessels.14,15 Benefiting from this passive targeting effect, well-defined NPs can reach the tumor

INTRODUCTION Taking into account the multiple physiological and biological barriers,1 ideal nanocarriers following systemic administration should possess the properties of (i) prolonged circulation in plasma, (ii) increased tumor accumulation, (iii) enhanced cellular uptake, and (iv) sufficient intracellular drug release. However, each new function would elevate complexity, and new barriers emerge.2 Often conflicts arise from different parts of specific functionality, e.g., stealth drug delivery systems that are designed to avoid the fast clearance by the reticuloendothelial system (RES), possess the properties of prolonged circulation time in blood and increased accumulation in tumor,3,4 but these systems prevent the internalization by tumor cells, resulting in low antitumor efficacy in vivo.5−7 Meanwhile, promising nanocarriers with targeted properties were designed to facilitate the cellular uptake; unfortunately, prolonged circulation and improved tumor accumulation were generally hard to achieve: sometimes surfaces conjugated with targeting ligands (e.g., c(RGDfK)) even can lead to accelerated removal of nanocarriers from blood circulation and may evoke © XXXX American Chemical Society

Received: June 27, 2014 Revised: August 20, 2014

A

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

Herein, to achieve multifunctional collaboration, we developed a new self-regulated delivery system based on the controlled self-assembly of mixed shell micelle. The mixed shell micelle was fabricated through the self-assembly of poly(ethylene glycol)-block-poly(ε-caprolactone) (PEG-b-PCL) and poly(ε-caprolactone)-block-poly(β-amino ester)-c(RGDfK) (PCL-b-PAE-c(RGDfK)) in water as shown in Figure 1A. In

site during blood circulation after systemic administration. However, as pointed out by Park et al., less than 5% of the administered NPs could actually be delivered to the desired tumor site, mainly attributed to the deposition in organs of the RES.16 To achieve prolonged blood circulation and increased tumor accumulation, strategies mainly toward the optimal parameter design of the physicochemical properties of NPs were developed, including surface modification with high density of hydrophilic polymers (such as PEG, PMPC, polysaccharide, etc.), slightly negative surface charge, and other methods.17 Unfortunately, unavoidable problems existed in that inefficient cellular internalization occurred for these systems, which derived from unfavorable binding to tumor cell membrane. Designing active-targeting NPs with ligands (e.g., antibodies, peptides, proteins, small molecules, and carbohydrates) attached on the surface is a well-documented specific binding approach.18 The active-targeting to membrane receptors or antigens and nonspecific binding endow NPs efficiently endocytosed by tumor cells. Nevertheless, surface conjugation with targeting ligands (e.g., c(RGDfK)) could facilitate the opsonization and immune responses, finally resulting in reduced tumor accumulation.8−11 Binding of NPs to the cell could also be achieved by introducing positive surface charge. For instance, cationic liposomes could be much more easily internalized by tumor cells than neutral and anionic liposomes as the cell membranes contain negative phospholipid bilayer.19 However, positive NPs could unspecific interact with blood components and have short circulation half-life.20−23 Recently, PEG sheddable NPs utilizing the unique tumor microenvironment (e.g., slightly acidic extracellular pH, overexpressed specific enzymes) were explored toward the conflicts of prolonged circulation and enhanced cellular uptake. Torchilin et al. developed a smart multifunctional drug delivery system to respond to the up-regulated matrix metalloprotease 2 (MMP2) in the tumor microenvironment.24 PEG chains were used to prolong the circulation time of the surface functionalized liposomal nanocarriers. The MMP2 sensitive linker between PEG and lipid was broken under the cleavage of MMP2, exposed with a cell-penetrating peptide (TATp), which enhanced the intracellular delivery of the system. Wang et al. designed a sheddable ternary nanoparticle for tumor aciditytargeted siRNA delivery.20 The nanoparticle was fabricated by introducing a tumor acidity-responsive PEGylated anionic polymer to the surface of positively charged polycation/ siRNA complexes. The PEGylation reduced the nonspecific interactions with protein, and the nanoparticle was capable of deshielding the PEG layer in response to tumor pH to facilitate the delivery of siRNA into the tumor cell. Masson et al. reported the shedding of PEG upon the hydrolysis of the orthoester linker of PEG-orthoester lipid at slightly acidic pH.25 The hydrolysis completed within 1 h at pH 5 at room temperature. Compared to lipoplexes coated with noncleavable PEG, the most optimal PEG-orthoester lipid conjugate revealed a 25-fold increase in transfection of HeLa cells. Nevertheless, Hennink et al. pointed out that most of the shedding studies so far still remained at the cell level without in vivo demonstration.26 Meanwhile, the shedding process itself and its in vivo kinetics were obviously difficult. Therefore, it is highly necessary to design novel multifunctional collaborated nanocarriers to overcome the conflict between prolonged blood circulation and enhanced cellular uptake, and finally promote the efficacy of cancer therapy.

Figure 1. (A) Schematic design of c(RGDfK)-decorated, pHresponsive, mixed shell micelles loaded with DOX (RMSM-DOX). (B) Illustration of the antitumor process of the RMSM-DOX after intravenous injection (i.v.). RMSM-DOX possessed the properties of (i) prolonged blood circulation, (ii) increased tumor accumulation, (iii) enhanced cellular internalization attributed to the charge conversion and targeting effect of exposed c(RGDfK) group at tumor acidic microenvironment, and finally (iv) sufficient intracellular drug release.

this study, the biodegradable cationic polymer PAE was utilized as the pH-responsive shell with pKa around 6.5.22,27−29 Under the pH variation from acidic conditions (pH ≈ 4) to the blood environment (pH ≈ 7.4), PAE undergoes a hydrophilic to hydrophobic phase transition. Thus, at pH 7.4, the micelle was covered with a rational ratio of hydrophilic/hydrophobic segment on the surface, and we have already demonstrated that the micelle with proper microphase separated surface had much better properties with prolonged blood circulation and reduced accumulation in RES organs (mainly liver and spleen) than a single PEGylated micelle.30 Meanwhile, with the protection of hydrophobic domain on the surface, burst release of the drug could be inhibited, and the biological stability could be increased due to the resistance to enzyme degradation.31−33 To avoid the accelerated clearance by the RES due to the modification of the c(RGDfK) group on the surface, here, the c(RGDfK) group was conjugated to the collapsed PAE domain and hidden in the shell of PEG at pH 7.4, as shown in Figure 1A. Benefiting from the self-regulated property in response to tumor pH (pH ≈ 6.5), we demonstrated that this kind of micelle possessed enhanced cellular uptake and efficient drug release attributed to the charge conversion and specific targeting effect of c(RGDfK), which could stretch out to the surface (Figure 1B), resulting in successful inhibition of tumor growth in nude mice.

■

EXPERIMENTAL METHODS

Materials. Monomethoxy poly(ethylene glycol) (CH3O−PEG− OH; Mw = 2000; Mw/Mn = 1.05; 99%, Fluka) and tert-butoxycarbonyl B

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

aminoethyl poly(ethylene glycol) (BOC−NH−PEG−OH; Mw = 2000; Mw/Mn = 1.05; 99.5%, Aldrich) were used after being dried under vacuum. ε-Caprolactone (ε-CL, 99%, J&K) was dried over calcium hydride and then purified by distillation under reduced pressure before use. Toluene was dried over calcium hydride and purified by distillation in the presence of Na. CHCl3 was dried over calcium hydride and purified by distillation. Sn(Oct)2 (95%), acryloyl chloride (95%), CF3COOH, N-acryloyloxysuccinimide (NAS), fluorescein isothiocyanate (FITC), Hexane-1,6-dioldiacrylate (HDD, 99%), and 4,4′-trimethylene dipiperidine (TDP, 97%) were purchased from J&K and used as received. c(RGDfK) was purchased from GL Biochem (Shanghai) Ltd. Doxorubicin-HCl (DOX·HCl) was supplied by Jingyan Chemicals Corporation (Shanghai, China). HepG2 (human liver hepatocellular carcinoma cell line) was cultured in RPMI-1640 medium. The medium was supplemented with 10% fetal bovine serum, 100 U/mL penicillin, and 100 μg/mL streptomycin. 3-(4,5Dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT), all cell culture media and supplies were ordered from Gibco (Gibco Corporation, Grand Island, NY, USA). Neutral Milli-Q water (18 MΩ) was used for the aqueous solutions. Four to five week-old Balb/c mice (female) were purchased from Experimental Animal Center of Academy of Military Medical Sciences (Beijing, China). Characterization. 1H NMR spectra were recorded on a Varian UNITY-plus 400 M NMR spectrometer at room temperature with tetramethylsilane (TMS) as an internal standard. The number-average molecular weight (Mn) and weight-average molecular weight (Mw) were measured by gel permeation chromatography (GPC) at 25 °C with a Waters 1525 chromatograph equipped with a Waters 2414 refractive index detector. GPC measurements were carried out using tetrahydrofuran (THF) as eluent with a flow rate of 1.0 mL/min, respectively. Polystyrene standards were used for calibration. Dynamic light scattering (DLS) experiments at a 90° scatter angle were performed on a laser light scattering spectrometer (BI-200SM) equipped with a digital correlator (BI-9000AT) at 636 nm at required temperature. All samples were obtained by filtering through a 0.45 μm Millipore filter into a clean scintillation vial. The final DLS samples were all kept at about 0.05 mg/mL with different aqueous buffer solution. Transmission electron microscopy (TEM) measurements were performed using a Philips T20ST electron microscope at an acceleration voltage of 100 kV. To prepare the TEM samples, the sample solution was dropped onto a carbon-coated copper grid and dried slowly at required temperature. The zeta potential values were measured on a Brookhaven ZetaPALS (Brookhaven Instrument, USA). The instrument utilizes phase analysis light scattering at 37 °C to provide an average over multiple particles. The Smoluchowski model was used to analyze the data here. The turbidity test of the single PAE micelle was performed by the determination of transmittance under the detection of UV−vis spectrophotometer (Purkinje General, China) at different pH values (from 5.0 to 7.4). The concentration of the micelle was 0.1 mg/mL, and acidic water (pH 5.0) was used as the control group. Synthesis of Polymers. The synthesis of block copolymer poly(ethylene glycol)-block-poly(ε-caprolactone) (PEG-b-PCL) and FITC-poly(ethylene glycol)-block-poly(ε-caprolactone) (FITC-PEGb-PCL) was synthesized by ring opening polymerization (ROP) of εCL using Sn(Oct)2 as the catalyst and PEG2k−OH and BOC-NHPEG2k−OH as the initiator in toluene solution, respectively. The synthesis routes were displayed in Supporting Information Figures S1A and S1B. The block copolymer poly(ε-caprolactone)-blockpoly(β-amino ester) (PCL-b-PAE) and poly(ε-caprolactone)-blockpoly(β-amino ester)-c(RGDfK) (PCL-b-PAE-c(RGDfK)) were synthesized through the combined methods of ROP, Michael addition reaction, and relative end group modification methods, as shown in Figure S1C according to a previous report.28 The DPs of the polymers were determined by the 1H NMR results shown in Figure S2. Detailed synthesis procedures were shown in the synthesis part of the Supporting Information. Preparation and Characterization of Micelles. Solution of PEG-b-PCL, PCL-b-PAE, and PCL-b-PAE-c(RGDfK) in CHCl3 with a concentration of 1 mg/mL were first prepared, respectively. For the

preparation of a specific micelle, an original solution of polymers of different kinds was mixed together first, and then the solution was added dropwise into a given amount of slightly acidic water (pH ≈ 4) under ultrasound. Then, the mixture was vigorous stirring at 30 °C for several hours followed by rotary evaporation. The obtained solution was dialyzed in a dialysis bag (molecular weight cut off: 12−14 KD) against neutral water for 2 days. The final micelle solution was lyophilized before use. Three types of micelles were fabricated according to the formulation shown in Table S1. The single PEGylated micelle (PM) was prepared only with PEG-b-PCL, while the mixed shell micelle (MSM) composed of equal weight of PEG and PAE segment on the surface of the PCL core was prepared with the mixture of PEG-b-PCL and PCL-b-PAE solution. The PCL-b-PAE-c(RGDfK) solution was added according to the calculated micelle formulation to make a mixed shell micelle with c(RGDfK) (RMSM) on the surface. For the cellular uptake experiments, the proper amount of FITC-PEGb-PCL was added into the original polymer solution to obtain FITClabeled micelles (Table S2). The single micelle with PAE as the shell was also finally dissolved in acidic water. Loading and in Vitro Drug Release of DOX. Drug-loaded micelles, PM-DOX, MSM-DOX, and RMSM-DOX were prepared by a similar method. A specific amount of DOX (without HCl) solution and polymer solution in CHCl3 were mixed first and added dropwise into the slightly acidic water (pH ≈ 4) under ultrasound followed by the same procedure as described above. The drug release experiment was performed in phosphate buffer with different pHs (7.4 and 6.5) under stirring, respectively. Briefly, 2 mL of the DOX-loaded micelle solution (PB, pH 7.4, ionic strength: 0.15 M) was transferred into a dialysis bag (molecular weight cut off: 12−14 KD), then the bag was immersed in 20 mL of buffer solution at 37 °C. Periodically, 1 mL of the solution outside the dialysis bag was taken out for fluorescent measurements. The volume of solution was kept constant by adding 1 mL of original buffer solution after each sampling. The DOX concentration was determined using a fluorescent spectrophotometer at 592 nm according to the calibration curve of DOX. Drug loading content (DLC) and drug loading efficiency (DLE) were calculated according to the following formulas:

DLC (wt%) = (weight of loaded drug/total weight of polymer and loaded drug) × 100%

DLE (%) = (weight of loaded drug/weight of drug in feed) × 100% Cellular Uptake and Viability Assays. HepG2 cells were seeded into 96-well plates at a density of 105 cells per well in 500 μL RPMI1640 medium/PBS. After an incubation of 24 h, the culture medium of each well was replaced with 500 μL of fresh medium with different pH values (pH 7.4 and 6.5) containing 10 μg/mL of FITC-labeled micelles (PM, MSM, or RMSM). After 2 h of further incubation, the culture medium was removed, and cells were washed three times with 500 μL PBS buffer. The cellular uptake of micelles was observed with an inverted fluorescence microscope (DMI6000B, Leica, Wetzlar, Germany). The cellular uptake experiments of the free DOX and the DOX-loaded micelles were performed with similar procedure in each well containing 10 μg/mL DOX. The cytotoxicity of free DOX and the DOX-loaded micelles (PM-DOX, MSM-DOX, and RMSM-DOX) were determined against HepG2 cells at different pH values (pH 7.4 and 6.5) by MTT viability assay. In the MTT assay, HepG2 cells were seeded into 96-well plates at a density of 4000 cells per well in 500 μL RPMI-1640 medium/PBS (pH 7.4). After 24 h incubation, half of the wells were adjusted to pH 6.5. Free DOX or the DOX-loaded micelles were added to each well with different concentrations (0.001, 0.01, 0.1, 0.5, 1, 5, 10, and 50 μg/mL). The saline solution was used as control. After 2 h further coincubation, the culture medium was removed and replaced with fresh medium. Twenty-four hours later, 25 μL 5 mg/mL 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl-tetrazolium bromide (MTT) assay was added into each well and the mixture was incubated for another 4 h. Then the obtained blue formazan crystals were dissolved C

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

Figure 2. (A) Hydrodynamic diameter distribution of different MSMs (PM, MSM, and RMSMs) in 10 mM PBS buffer (pH 7.4 with 150 mM NaCl) measured by DLS at 37 °C. (B) TEM image of RMSM at 37 °C. Scale bar, 100 nm. TEM Images of other micelles can be seen in the Supporting Information, Figure S3. in 150 μL dimethyl sulfoxide (DMSO). The absorbance was measured at a wavelength of 570 nm and the viability was expressed as the percentage of the control. The half maximal inhibitory concentration (IC50) value of each group was calculated according to the results. Antitumor Activity. The animal studies were performed in accordance with the Regulations for the Administration of Affairs Concerning Experimental Animals (Tianjin, revised in June 2004) and adhered to the Guiding Principles in the Care and Use of Animals of the American Physiological Society. The in vivo antitumor activity of drug loaded mixed shell micelles, PM-DOX, MSM-DOX, and RMSMDOX, were evaluated in 4−5 weeks old female nude mice (BALB/c mice). To establish the xenograft tumors, a cell suspension containing 4 × 105 HepG2 cells were injected subcutaneously on the flank of the mice. The tumor volume reached about 200 mm3 10 days later, and then the mice were randomly divided into five groups (10 mice in each group): saline, free DOX, PM-DOX, MSM-DOX, and RMSM-DOX, respectively. The dose of DOX was fixed at 5 mg/kg body weight. Saline, free DOX and DOX-loaded micelles were administrated for six times through a tail vein every other day. Tumor volume and weight of the mice were monitored every 2 days for 30 days. The estimated tumor volume was calculated by the formula

amphiphilic block copolymers PEG-b-PCL, PCL-b-PAE, and PCL-b-PAE-c(RGDfK), which were synthesized as shown in the Supporting Information. Figure S1 clearly shows the synthetic routes of PEG-b-PCL, FITC-PEG-b-PCL, PCL-bPAE, and PCL-b-PAE-c(RGDfK). Specifically, the block copolymer PCL-b-PAE-c(RGDfK) was successfully synthesized through a multistep synthesis as shown in Figure S1C, and the successful conjugation of c(RGDfK) to the end group of PCLb-PAE was confirmed by 1H NMR as shown in Figure S2D. The appearance of the characteristic peak at 7.3−7.4 ppm clearly indicated the phenyl group of c(RGDfK). Further evidence of successful conjugation of c(RGDfK) will be given in the cellular uptake efficiency of different micelles. The fabrication procedure of the DOX-loaded c(RGDfK) decorated mixed shell micelle (RMSM-DOX) is shown in Figure 1A. Similarly, three kinds of micelles, including single PM, MSM, and RMSM, were fabricated according to selfassembly of relative polymers. RMSM consisted of a hydrophobic PCL core, a mixed shell of PEG, PAE, and PAEc(RGDfK). According to our previous work,30 equal weight amount of hydrophobic to hydrophilic segments on the surface of NPs would prolong the blood circulation time and reduce the deposition in liver and spleen. Therefore, in the design of MSM and RMSM, the weight ratio of PEG to PAE segment (including PAE and PAE-c(RGDfK)) was kept at 1. The feeding amounts of relative polymers were shown in Table S1. The weight ratio of PEG/PAE/PAE-c(RGDfK) was 5:4:1, while for PM and MSM, the ratios were 5:5:0 and 1:0:0, respectively. From the DLS and TEM measurements, the obtained micelles possessed similar size and size distribution (80 ± 2 nm, 85 ± 1 nm, and 92 ± 1 nm for PM, MSM, and RMSM, Figure 2). The segment PAE, which is a pH-responsive biodegradable cationic polymer with a pKa ≈ 6.5, has been extensively studied in recent years. During the fabrication of micelles, at pH 4.0, PAE was hydrophilic and positively charged. When the pH changed into 7.4, at normal physiological environment, PAE became hydrophobic and collapsed onto the surface of the PCL core. Zhang et al.28 utilized MPEG-b-(PLA-co-PAE) block copolymer micelles as anticancer drug delivery nanocarriers with pH-responsive property. Compared with this method, our strategy possesses the advantage of easily tuning the ratio of PEG and PAE in the micelle shell without additional synthesis of copolymer. The

Tumor size (mm3) = Length × Width2 /2 Histomorphological and Immunology Analysis. Twenty days after the first administration, one tumor-bearing mouse in each group was anesthetized with 8% Chloral hydrate and sacrificed. Tumor, liver, and spleen were collected. Tumor samples were fixed for 24 h in 4% paraformaldehyde, embedded in paraffin, and cut into 8-μm-thick sections for hematoxylin/eosin (H&E) and in situ terminal deoxynucleotidyl transferase-mediated dUTP nick end labeling (TUNEL) assays according to the manufacturer’s instructions. The photos were taken using optical microscope (Leica DMI6000 B). Statistical Analysis. One-way analysis of variance (ANOVA) was utilized for the statistical analysis. The results were shown as mean ± standard deviation.

■

RESULTS AND DISCUSSION Formation of Multifunctional Mixed Shell Micelles. For the construction of active targeting drug delivery system in therapeutical applications, c(RGDfK) is a widely used cyclic peptide that is usually chosen to highly bind with the αvβ3 integrin receptors on the angiogenic endothelium in malignant or diseased tissues.34−38 Here, to explore multifunctional targeted drug delivery system, c(RGDfK) was induced in the fabrication of nanoparticles. The multifunctional mixed shell micelles were fabricated through the self-assembly of the D

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

RMSM was about +3 mV and +5 mV at pH 6.5, respectively, while the neutral point was about pH 6.8. Although the magnitude of the zeta potential of the micelles was low, the micelles could keep quite stable as the shell consisting hydrophilic PEG, which increased the stability of the colloids (DLS data of RMSM at pH 6.5 and 7.4 are shown in Supporting Information Figure S7). Moreover, the increasing amount of PAE in the shell of the mixed shell micelle had a concomitant increase in the positive charge of the micelles (Figure S6). Therefore, the micelles with pH responsive PAE shell had the ability of charge conversion according to the variation of the environment pH, which was attributed to the protonation/deprotonation of PAE. As we know, the pH is around 7.4 in blood, while in tumor tissue, the pHe is about 6.5 caused by the anaerobic glycolysis.39 Also, in the endosome, the pH value is about 5.0−6.0. Thus, during the journey from blood to tumor cell, accompanying the change of pH value, the charge conversion of MSM and RMSM from negative to positive would occur in a mildly acidic tumor microenvironment and endow them with higher positive charge at the endosomal acidity. It should be pointed out that this kind of charge conversion is a fast and reversible process caused by the transition of protonation state. The traditional charge conversion of NPs generally depends on the breakage of sensitive chemical linker utilizing the unique tumor microenvironment (e.g., slightly acidic extracellular pH, overexpressed specific enzymes).40,41 The biological or chemical hydrolysis of these linkers is always a slow kinetic process, which may take hours or more. Moreover, the chemical variation is usually nonreversible. After the EPR effect, when the NPs turn positive and penetrate into the tumor parts through the interstitial transport, the convective flow from the tumor core would transport the positive NPs back into the blood plasma. As the NPs possess positive charge, they would probably easily be cleared out of blood by RES.42 On the other hand, the NPs accumulated in tumor sites might go back to the blood circulation. Once they became positively charged and circulated in blood again, they could be cleared quickly by the RES. For the NPs studied here, this problem would not occur as the charge of NPs would turn back to negative in the plasma. In Vitro Drug Release of DOX-Loaded Micelles. The anticancer drug, DOX, was using as a model drug to further explore the in vitro drug release of DOX-loaded micelles. Three kinds of DOX-loaded micelles, PM-DOX, MSM-DOX, and RMSM-DOX were prepared. As shown in Table 1, the DLC

desired composition and components can be conveniently obtained from controllable formulations. Considering the probability of promoted dramatic opsonization and recognition by the RES after introducing the c(RGDfK) group to the surface of the NPs, in this work, the c(RGDfK) group that was covalently linked to the end of PAE segment would be subsequently pulled onto the surface of the PAE microdomain at pH 7.4 to reduce the interaction between c(RGDfK) and the opsonic proteins in blood. pH Sensitivity of PAE and Charge Conversion of Mixed Shell Micelles. In order to test the pH sensitive of PAE, the turbidity experiment was performed using the single PAE micelle made from block copolymer PCL-b-PAE in acidic water. The pH was adjusted from 5.0 to 7.4, and the transmittance of the solution was measured with acidic water (pH 5.0) as control as shown in Figure 3A. When the pH

Figure 3. (A) Transmittance curve of the single PAE micelle and (B) zeta potential of different micelles (PM, MSM, and RMSM) at various pH values at 37 °C. (n = 3). The buffer solutions were 10 mM NaAc/ HAc buffer for pH 5.0 and 6.0, and 10 mM PBS buffer for pH 6.5, 6.8, and 7.4 using HNO3 or NaOH aqueous solution to adjust the pH values, respectively.

increased from 5.0 to 6.5, the transmittance of the solution was gradually changed from 65.9% to 54.0%, decreased by about 12%. This can be attributed to the deprotonation and subsequently hydrophilic to hydrophobic transition of PAE block which was stretched out of the single PAE micelle at the original state. At the pH range of 6.5−7.0, dramatic reduction in solution transparency occurred and the solution became completely cloudy at pH 7.4, presumably due to the aggregation of single PAE micelle. TEM images of PAE single micelle showed that the size and morphology changed with the variation of pH values (shown in Supporting Information Figure S4). At pH 5.0, the size of PAE micelle was about 100 nm, and the morphology was sphere, while upon the increasing of pH value from pH 6.5 to 7.4, the size turned to be larger (from around 150 nm to several μm). Additionally, the morphology changed from sphere to random aggregates. DLS measurements also demonstrated a similar phenomenon (Supporting Information Figure S5). The charge conversion phenomena of the mixed shell micelles (MSM and RMSM) can be found in the zeta potential measurements (Figure 3B). The micelle of PM showed little charge variation from pH 7.4 to 5.0 (−9.37 ± 1.98 mV to −4.21 ± 0.29 mV) and no charge conversion happened, while MSM and RMSM exhibited apparent charge conversion from −12.3 ± 0.2 mV to +15.4 ± 0.4 mV and −9.5 ± 1.8 mV to +10.5 ± 1.3 mV, respectively. The surface charge of MSM and

Table 1. DLC and DLE of the Different DOX-Loaded Micelles (n = 3) DOX-loaded micelles

DLC (wt %)

DLE (%)

PM-DOX MSM-DOX RMSM-DOX

7.5 ± 1.2 8.6 ± 0.8 9.9 ± 1.3

16.2 ± 2.6 18.8 ± 1.7 21.9 ± 2.9

was 7.5 ± 1.2 wt %, 8.6 ± 0.8 wt %, and 9.9 ± 1.3 wt % for PMDOX, MSM-DOX, and RMSM-DOX respectively. The corresponding DLE was 16.2 ± 2.6%, 18.8 ± 1.7%, and 21.9 ± 2.9%. The differences between the drug loading perhaps derived from the structure of different formulations. Differences of the structure mainly existed in the shell composition: PMDOX only has PEG as the shell, while MSM-DOX and RMSMDOX have both hydrophilic and hydrophobic PAE segments as a mixed shell. It can be seen in Figure 4B that the DOX release E

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

Figure 4. DOX release profile of PM-DOX, MSM-DOX, and RMSM-DOX at different pH values (pH 5.0, 6.5, and 7.4) at 37 °C. (A) The comparison of PM-DOX and MSM-DOX. (B) The comparison of PM-DOX and RMSM-DOX. (n = 3).

HepG2 cells, finally enhancing the endocytosis procedure which occurred subsequently. It was reported that c(RGDfK) has high affinity and specificity for the αvβ3 integrin receptor which is overexpressed on human primary tumors, tumor cell lines, and tumor neovasculature arising during angiogenesis.43−45 At pH 6.5, as the solubility of PAE increased to a certain degree, c(RGDfK) would stretch out and expose on the surface of RMSM, resulting in enhancement of the binding efficiency for RMSM with tumor cell through the specific interaction between c(RGDfK) and αvβ3 integrin receptor. Combined with the positive charge transition of PAE and the targeting effect of c(RGDfK), RMSM was endowed with the dramatic enhanced endocytosis at slightly acidic environment. The pH sensitivity of PAE endows the probability of solving the conflict of prolonged circulation and enhanced cellular uptake. It was demonstrated that the equal weight of hydrophilic to hydrophobic segments on the surface of the micelle (mixed shell micelle with microphase separation) showed prolonged blood circulation compared with a single PEGylated micelle with similar physiochemical properties.30 Here, a mixed shell micelle with surface microphase separation was easily obtained at pH 7.4. Additionally, as the c(RGDfK) group was hidden in the shielding of PEG shell at pH 7.4, the unfavorable accelerated clearance by the RES caused by the exposed c(RGDfK) group reported previously could be avoided to a certain degree. Further evidence of the pH-dependent efficacy of tumor cellular uptake is shown in Figure 6, which displays the cellular uptake of different formulations (free DOX, PM-DOX, MSMDOX, and RMSM-DOX) at pH 7.4 and 6.5. After 2 h incubation of different groups (each contained 10 μg/mL DOX) with HepG2 cells shown in Figure 6, for MSM-DOX and RMSM-DOX, much stronger fluorescence of DOX was observed in HepG2 cells at pH 6.5 than that at pH 7.4, while the fluorescence remained nearly constant for free DOX, and the PM-DOX group showed slightly decreased fluorescence. Due to the faster drug release behavior for the PM-DOX group, PM-DOX showed cellular internalization behavior similar to that of free DOX at pH 7.4 and 6.5 after short time incubation. On the other hand, for MSM-DOX and RMSM-DOX groups, at pH 7.4, with the protection of hydrophobic PAE polymers in the shell, the cellular internalization of MSM-DOX and RMSMDOX groups and the release rate of DOX could be decreased, resulting in low fluorescence intensity which could not be clearly observed. However, at pH 6.5, surface charge of MSM-

rates of MSM-DOX and RMSM-DOX were nearly at the same level, while PM-DOX showed a relatively faster release rate at different pHs (Figure 4A). This could be attributed to the shielding effect of the hydrophobic PAE layer at pH 7.4 and the electrostatic repulsion between PAE and the partly positive charged DOX at lower pHs (pH 6.5 and 5.0). Cellular Experiments of Different Micelles and DOXLoaded Micelles. To investigate the cellular uptake of different micelles, HepG2 cells were incubated with FITClabeled PM, MSM, and RMSM at pH 7.4 and pH 6.5. FITClabeled micelles showed green fluorescence, which can be seen in Figure 5. After incubation with HepG2 cells for 2 h, the

Figure 5. Cellular uptake of FITC-labeled PM, MSM, and RMSM on HepG2 cells observed by inverted fluorescent microscopy. Cells were co-incubated with micelles for 2 h at pH 7.4 and 6.5. Scale bar, 100 μm.

fluorescence was quite weak at pH 7.4, reflecting the insignificant nonspecific cellular uptake of PM, MSM, or RMSM at pH 7.4. However, at pH 6.5, significantly stronger fluorescence intensity can be clearly observed, indicating a pHdependent uptake process of MSM and RMSM. On the contrary, the PM group without PAE in the micelle shell kept relatively constant fluorescence intensity. Previous studies have shown that NPs with positive charge can more easily bind with cells and subsequently easier internalized than negatively charged NPs.20−23 Here, the cellular uptake results verified this and the negative to positive charge conversion of the MSM and RMSM assisted with the binding between micelles and F

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

Figure 6. Inverted fluorescent microscopy images of HepG2 cells after incubation with free DOX, PM-DOX, MSM-DOX, and RMSM-DOX at pH 7.4 and 6.5 for 2 h, respectively. The cell nuclei were stained with DAPI (blue). Scale bar, 100 μm.

Figure 7. Cell viability of free DOX, PM-DOX, MSM-DOX, and RMSM-DOX at (A) pH 7.4 and (B) pH 6.5 determined by MTT assay. The IC50 for each group was inserted. Cells were incubated with various formulations for 2 h at pH 7.4 and 6.5, respectively. Then all culture mediums were replaced with fresh medium (pH 7.4) followed by further incubation for 24 h (n = 6).

Figure 8. (A) In vivo antitumor inhibition of free DOX, PM-DOX, MSM-DOX, and RMSM-DOX. (B) In situ terminal deoxynucleotidyl transferasemediated dUTP nick end labeling (TUNEL) staining of the tumor sections after the treatments of (a) saline, (b) free DOX, (c) PM-DOX, (d) MSM-DOX, and (e) RMSM-DOX. TUNEL positive apoptotic cells were stained brown. Scale bar, 50 μm. (n = 10).

G

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

circulation, enhanced cellular uptake, and favorable pH dependent cytotoxicity facilitated the antitumor efficacy of DOX-loaded RMSM.

DOX and RMSM-DOX groups varied from negative charge to positive charge, leading to stronger interaction with the cell membrane and faster DOX release rate. Meanwhile, at pH 6.5, the c(RGDfK) group could stretch out of the micelle shell and facilitate the interaction with the cell. The MTT assay was performed to further evaluate the cytotoxic effects of the DOX-loaded micelles on HepG2 cells. After 2 h co-incubation at different concentrations at pH 7.4 and 6.5, the culture medium was replaced with fresh medium with pH 7.4, and the cells were incubated for another 24 h, respectively. The MTT results were consistent with the cell uptake of different formulations. As shown in Figure 7, all formulations exhibited a dose dependent manner. For free DOX, the half maximal inhibitory concentration (IC50) was similar at pH 7.4 and 6.5, which were 9.35 and 9.05 μg/mL respectively. At pH 7.4, the IC50 of PM-DOX was quite low (1.17 μg/mL), indicating its serious cytotoxic effect. MSMDOX and RMSM-DOX exhibited large difference with high IC50 at pH 7.4, 20.03, and 14.80 μg/mL, compared with 0.20 and 0.13 μg/mL at pH 6.5, suggesting pH-dependent enhanced cellular uptake and efficient cell killing ability in response to tumor microenvironment. These findings indicated the high efficiency of mixed shell micelles in reducing the dose of antitumor drugs to exert therapeutic effects and minimizing drug-induced side effects. Also, due to the low cellular internalization at pH 7.4, low cytotoxicity of the MSM-DOX and RMSM-DOX could be also expected in normal organs, where no charge conversion might occur in the intercellular environment. Antitumor Activity. Finally, to verify that the self-regulated functional collaborated mixed shell micelles have the potential to improve antitumor efficacy, we assessed the tumor growth following administration of different formulations, including saline, free DOX, PM-DOX, MSM-DOX, and RMSM-DOX. The nude mice were intravenously injected via tail vein with a dose of 5 mg DOX/kg mice body weight six times every 2 days. Figure 8A showed the variation of relative tumor volume after first injection in 30 days. After 30 days, administration of free DOX slightly inhibited the tumor growth, with 18.1-fold relative volume compared to 21.7-fold of saline group. PMDOX and MSM-DOX showed better inhibition with 13.1-fold and 11.0-fold, while the RMSM-DOX exhibited minimized tumor growth with only 5.2-fold relative volume, about onefourth of the saline group. Body weight changes of mice treated with different formulations can be seen in Figure S8. The anticancer efficacy of RMSM-DOX was also demonstrated by the TUNEL (terminal deoxynucleoitidyl transferase) assay to detect the apoptosis at the tumor tissues of different groups at day 20 after the first administration expressed in Figure 8B. Apoptotic cells could be rarely detectable in saline group, but were clearly detected in DOX and DOX-loaded groups. Administration of MSM-DOX and RMSM-DOX markedly increased the apoptotic cells in tumor tissue, indicating the improved efficacy in inducing apoptosis and inhibiting proliferation of tumor cells. H&E assay was also performed to evaluate the toxicology of the DOX-loaded micelles. The results depicted in Figure S9 clearly presented that MSM-DOX and RMSM-DOX increased the percentage of apoptotic or necrosis tumor cells compared to other groups and displayed no obvious histopathological lesions in liver and spleen. These in vivo data directly presented the superior therapeutic efficacy of this pH responsive drug delivery system, especially the RMSM-DOX group. In a word, the properties of prolonged

■

CONCLUSIONS In summary, to overcome the conflicts widely existed in normal drug delivery systems and promote the cancer therapeutic efficacy, we originally designed a novel kind of pH responsive mixed shell micelle as drug nanocarrier and demonstrated the improved antitumor effect. Benefiting from the rational design with pH sensitive PAE segment in the shell, the micelle was fabricated with proper ratio of hydrophilic and hydrophobic domain on the surface, endowing its prolonged circulation according to our previous report. Moreover, in response to the tumor acidic microenvironment, charge conversion and targeting group exposure occurred, facilitating the enhanced cellular uptake. Finally, significant promotion of therapeutic effect was successfully obtained according to the tumor-bearing BALB/c mice experiments. This delivery strategy provided a simple and versatile approach to overcome the obstacles in drug delivery system. We anticipate this kind of shell responsive mixed micelles can serve as anticancer drug carriers in future clinical trials.

■

ASSOCIATED CONTENT

* Supporting Information S

Detailed synthesis routes, 1H NMR, formulation of micelles, TEM, zeta potential results, mice body weight changes, haematoxylin and eosin (H&E) examination of tumor, liver, and spleen tissues, and other data are available free of charge via the Internet at http://pubs.acs.org.

■

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare no competing financial interest.

■

ACKNOWLEDGMENTS We thank the National Natural Science Foundation of China (Nos. 91127045, 51390483, 21274001, 81171371, 51203189, and 51303213), the National Basic Research Program of China (973 Program, No. 2011CB932503), and PCSIRT (IRT1257) for financial support.

■

REFERENCES

(1) Elsabahy, M.; Wooley, K. L. Chem. Soc. Rev. 2012, 41, 2545− 2561. (2) Cheng, Z.; Al Zaki, A.; Hui, J. Z.; Muzykantov, V. R.; Tsourkas. Science 2012, 338, 903−910. (3) Dufort, S.; Sancey, L.; Coll, J. Adv. Drug Deliver. Rev. 2012, 64, 179−189. (4) Gao, W.; Langer, R.; Farokhzad, O. C. Angew. Chem., Int. Ed. 2010, 49, 6567−6571. (5) Ulbrich, K.; Šubr, V. Adv. Drug Deliver. Rev. 2004, 56, 1023− 1050. (6) Huo, S.; Ma, H.; Huang, K.; Liu, J.; Wei, T.; Jin, S.; Zhang, J.; He, S.; Liang, X. Cancer Res. 2013, 73, 319−330. (7) Etrych, T.; Jelínková, M.; Ř íhová, B.; Ulbrich, K. J. Controlled Release 2001, 73, 89−102. (8) Shuhendler, A. J.; Prasad, P.; Leung, M.; Rauth, A. M.; DaCosta, R. S.; Wu, X. Y. Adv. Healthcare Mater. 2012, 1, 600−608. (9) Allen, T. M.; Brandeis, E.; Hansen, C. B.; Kao, G. Y.; Zalipsky, S. Biochim. Biophys. Acta, Biomembr. 1995, 1237, 99−108. H

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

Article

(10) Hu, G.; Lijowski, M.; Zhang, H.; Partlow, K. C.; Caruthers, S. D.; Kiefer, G.; Gulyas, G.; Athey, P.; Scott, M. J.; Wickline, S. A.; Lanza, G. M. Int. J. Cancer 2007, 120, 1951−1957. (11) Schiffelers, R. M.; Koning, G. A.; Ten Hagen, T. L. M.; Fens, M. H. A. M.; Schraa, A. J.; Janssen, A. P. C. A.; Kok, R. J.; Molema, G.; Storm, G. J. Controlled Release 2003, 91, 115−122. (12) Li, Y.; Wang, J.; Wientjes, M. G.; Au, J. L. S. Adv. Drug Deliver. Rev. 2012, 64, 29−39. (13) Mo, R.; Sun, Q.; Xue, J.; Li, N.; Li, W.; Zhang, C.; Ping, Q. Adv. Mater. 2012, 24, 3659−3665. (14) Maeda, H.; Bharate, G. Y.; Daruwalla, J. Eur. J. Pharm. Biopharm. 2009, 71, 409−419. (15) Maeda, H.; Wu, J.; Sawa, T.; Matsumura, Y.; Hori, K. J. Controlled Release 2000, 65, 271−284. (16) Bae, Y. H.; Park, K. J. Controlled Release 2011, 153, 198−205. (17) Yoo, J.; Chambers, E.; Mitragotri, S. Curr. Pharm. Design 2010, 16, 2298−2307. (18) Byrne, J. D.; Betancourt, T.; Brannon-Peppas, L. Adv. Drug Deliver. Rev. 2008, 60, 1615−1626. (19) De Smedt, S.; Demeester, J.; Hennink, W. Pharm. Res. 2000, 17, 113−126. (20) Yang, X.; Du, J.; Dou, S.; Mao, C.; Long, H.; Wang, J. ACS Nano 2011, 6, 771−781. (21) He, C.; Hu, Y.; Yin, L.; Tang, C.; Yin, C. Biomaterials 2010, 31, 3657−3666. (22) Lynn, D. M.; Langer, R. J. Am. Chem. Soc. 2000, 122, 10761− 10768. (23) Cho, E. C.; Xie, J.; Wurm, P. A.; Xia, Y. Nano Lett. 2009, 9, 1080−1084. (24) Zhu, L.; Kate, P.; Torchilin, V. P. ACS Nano 2012, 6, 3491− 3498. (25) Masson, C.; Garinot, M.; Mignet, N.; Wetzer, B.; Mailhe, P.; Scherman, D.; Bessodes, M. J. Controlled Release 2004, 99, 423−434. (26) Romberg, B.; Hennink, W.; Storm, G. Pharm. Res. 2008, 25, 55−71. (27) Lynn, D. M.; Amiji, M. M.; Langer, R. Angew. Chem., Int. Ed. 2001, 40, 1707−1710. (28) Zhang, C. Y.; Yang, Y. Q.; Huang, T. X.; Zhao, B.; Guo, X. D.; Wang, J. F.; Zhang, L. J. Biomaterials 2012, 33, 6273−6283. (29) Potineni, A.; Lynn, D. M.; Langer, R.; Amiji, M. M. J. Controlled Release 2003, 86, 223−234. (30) Gao, H.; Xiong, J.; Cheng, T.; Liu, J.; Chu, L.; Liu, J.; Ma, R.; Shi, L. Biomacromolecules 2013, 14, 460−467. (31) Wu, C.; Ma, R.; He, H.; Zhao, L.; Gao, H.; An, Y.; Shi, L. Macromol. Biosci. 2009, 9, 1185−1193. (32) Ma, R.; Shi, L. Macromol. Biosci. 2010, 10, 1397−1405. (33) Xu, Y.; Ma, R.; Zhang, Z.; He, H.; Wang, Y.; Qu, A.; An, Y.; Zhu, X. X.; Shi, L. Polymer 2012, 53, 3559−3565. (34) Hynes, R. O. Nat. Med. 2002, 8, 918−921. (35) Murphy, E. A.; Majeti, B. K.; Barnes, L. A.; Makale, M.; Weis, S. M.; Lutu-Fuga, K.; Wrasidlo, W.; Cheresh, D. A. Proc. Natl. Acad. Sci. U. S. A. 2008, 105, 9343. (36) Brooks, P. C.; Clark, R.; Cheresh, D. A. Science 1994, 264, 569. (37) Schiffelers, R. M.; Koning, G. A.; ten Hagen, T. L. M.; Fens, M.; Schraa, A. J.; Janssen, A.; Kok, R. J.; Molema, G.; Storm, G. J. Controlled Release 2003, 91, 115−122. (38) Temming, K.; Schiffelers, R. M.; Molema, G.; Kok, R. J. Drug Resist. Updates 2005, 8, 381−402. (39) Warburg, O. Science 1956, 123, 309−314. (40) Du, J.; Sun, T.; Song, W.; Wu, J.; Wang, J. Angew. Chem., Int. Ed. 2010, 122, 3703−3708. (41) Oh, N. M.; Kwag, D. S.; Oh, K. T.; Youn, Y. S.; Lee, E. S. Biomaterials 2012, 33, 1884−1893. (42) Xiao, K.; Li, Y.; Luo, J.; Lee, J. S.; Xiao, W.; Gonik, A. M.; Agarwal, R. G.; Lam, K. S. Biomaterials 2011, 32, 3435−3446. (43) Miura, Y.; Takenaka, T.; Toh, K.; Wu, S.; Nishihara, H.; Kano, M. R.; Ino, Y.; Nomoto, T.; Matsumoto, Y.; Koyama, H.; Cabral, H.; Nishiyama, N.; Kataoka, K. ACS Nano 2013, 7, 8583−8592.

(44) Morlieras, J.; Dufort, S.; Sancey, L.; Truillet, C.; Mignot, A.; Rossetti, F.; Dentamaro, M.; Laurent, S.; Vander Elst, L.; Muller, R. N.; Antoine, R.; Dugourd, P.; Roux, S.; Perriat, P.; Lux, F.; Coll, J.; Tillement, O. Bioconjugate Chem. 2013, 24, 1584−1597. (45) Haubner, R.; Gratias, R.; Diefenbach, B.; Goodman, S. L.; Jonczyk, A.; Kessler, H. J. Am. Chem. Soc. 1996, 118, 7461−7472.

I

dx.doi.org/10.1021/bm5009348 | Biomacromolecules XXXX, XXX, XXX−XXX