Mitomycin C-Soybean Phosphatidylcholine Complex-Loaded Self

Article pubs.acs.org/molecularpharmaceutics

Mitomycin C‑Soybean Phosphatidylcholine Complex-Loaded SelfAssembled PEG-Lipid-PLA Hybrid Nanoparticles for Targeted Drug Delivery and Dual-Controlled Drug Release Yang Li,†,§ Hongjie Wu,‡ Xiangrui Yang,† Mengmeng Jia,† Yanxiu Li,†,‡ Yu Huang,† Jinyan Lin,†,§ Shichao Wu,†,§ and Zhenqing Hou*,† †

Research Center of Biomedical Engineering & Department of Biomaterials, College of Materials, Xiamen University, Xiamen 361005, China ‡ Department of Pharmacy, School of Pharmaceutical Science, Xiamen University, Xiamen 361005, China § Department of Chemistry, College of Chemistry & Chemical Engineering, Xiamen University, Xiamen 361005, China S Supporting Information *

ABSTRACT: Most present drug−phospholipid delivery systems were based on a water-insoluble drug−phospholipid complex but rarely water-soluble drug−phospholipid complex. Mitomycin C (MMC) is a water-soluble anticancer drug extensively used in first-line chemotherapy but is limited by its poor aqueous stability in vitro, rapid elimination from the body, and lack of target specificity. In this article, we report the MMC-soybean phosphatidylcholine complex-loaded PEGlipid-PLA hybrid nanoparticles (NPs) with Folate (FA) functionalization (FA-PEG-PE-PLA NPs@MMC-SPC) for targeted drug delivery and dual-controlled drug release. FAPEG-PE-PLA NPs@MMC-SPC comprise a hydrophobic core (PLA) loaded with MMC-SPC, an amphiphilic lipid interface layer (PE), a hydrophilic shell (PEG), and a targeting ligand (FA) on the surface, with a spherical shape, a nanoscaled particle size, and high drug encapsulation efficiency of almost 95%. The advantage of the new drug delivery systems is the early phase controlled drug release by the drug−phospholipid complex and the late-phase controlled drug release by the pH-sensitive polymer−lipid hybrid NPs. In vitro cytotoxicity and hemolysis assays demonstrated that the drug carriers were cytocompatible and hemocompatible. The pharmacokinetics study in rats showed that FA-PEG-PE-PLA NPs@MMC-SPC significantly prolonged the blood circulation time compared to that of the free MMC. More importantly, FA-PEG-PE-PLA NPs@MMC-SPC presented the enhanced cell uptake/cytotoxicity in vitro and superior tumor accumulation/therapeutic efficacy in vivo while reducing the systemic toxicity. A significant accumulation of MMC in the nuclei as the site of MMC action achieved in FA-PEGPE-PLA NPs@MMC-SPC made them ideal for MMC drug delivery. This study may provide an effective strategy for the design and development of the water-soluble drug−phospholipid complex-based targeted drug delivery and sustained/controlled drug release. KEYWORDS: mitomycin C, complex, self-assembled hybrid NPs, drug delivery, tumor targeting, pH-sensitive

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INTRODUCTION

types of P-glycoprotein-mediated multidrug resistant tumor cells.5−8 Nonetheless, the dose-limiting toxicity, subacute and cumulative myelosuppression, and nephrotoxicity limit its therapeutic efficacy. Moreover, the clinical use of MMC is significantly hindered due to the rapid elimination from the body with a low plasma level around the effect-relevant sites in vivo. Many new and innovative strategies to delivery MMC using different drug carriers including lipid carrier (liposomes8,9

In the field of cancer treatment, traditional therapeutic strategies, such as invasive surgery, radiotherapy, chemotherapy, or their combinations, are conceived as the primary treatment options. Unfortunately, it is difficult to completely remove tumor tissue in most cases. Even worse, cancer caused more than 10% of all human deaths all over the world,1 which highlighted the urgent need for more effective therapeutic strategies. Mitomycin C (MMC), a powerful water-soluble antibacterial and antitumor antibiotic agent,2,3 is the first line treatment extensively used for a wide range of cancers.4 MMC is a poor substrate for P-glycoprotein and retains activity against many © XXXX American Chemical Society

Received: April 6, 2014 Revised: May 31, 2014 Accepted: July 2, 2014

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dx.doi.org/10.1021/mp500254j | Mol. Pharmaceutics XXXX, XXX, XXX−XXX

Molecular Pharmaceutics

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Figure 1. Preparation and characteristics of FA-PEG-PE-PLA@MMC-SPC and their targeted drug delivery. (A) Illustration of FA-PEG-PE-PLA@ MMC-SPC: a hydrophobic core (PLA) loaded with MMC-SPC and an amphiphilic lipid interface layer (PE) and a hydrophilic shell (PEG) with a targeting ligand (FA). (B) Particle size distribution. (C) Zeta potential distribution. (D) SEM image. Scale bars = 200 nm. (E) TEM image. Scale bars = 200 nm. (F) LCSM image of FA-PEG-PE-PLA@MMC-SPC@DiD. Scale bars = 10 μm. (G) Once intravenously administrated, (1) the NPs were accumulated at the tumor site via the enhanced permeability and retention (EPR) effect. Subsequently, (2) the NPs were internalized by the cells via FA receptor-mediated endocytosis, and (3) the low pH environment induced the rapid MMC release. Lastly, (4) MMC was delivered to the nuclei and presented the anticancer activity.

and phytosomes10) and polymeric carrier11,12 have been developed to overcome these drawbacks. However, for example, the use of liposomal formulations was reported to have the stability problems during storage.13 Therefore, there is still an unmet need for formulating a safe, stable, and efficient MMC delivery system for in vivo. Nanoparticles (NPs) have been regarded as a promising alternative to deliver anticancer drugs via the enhanced permeation and retention (EPR) effect for cancer treatment. The self-assembled polymer−lipid hybrid NPs as promising drug delivery systems, which can overcome the rapid blood clearance and severe side effects of anticancer drugs, control drug release, and achieve tumor selectivity, have attracted considerable interest for drug and gene delivery.14,15 Another major concern of MMC is the poor aqueous stability in vitro and rapid elimination from the body in vivo. In recent years, the drug−phospholipid complex was a key bridge connecting the conventional and novel drug delivery systems because of the significant improvement of its safety and efficacy.16−18 To address the issue, we prepared the MMCsoybean phosphatidylcholine complex (MMC-SPC) to increase the lipophilicity and stability of MMC in vitro. On one hand, MMC interacted with SPC via electrostatic interaction hydrogen bonds and van der Waals forces, and resulted in almost 100% complexation efficiency. On the other hand, MMC-SPC acted as the amphiphilic drug−phospholipid complex, entrapping drug within phospholipid and resembling

phospholipid in being substantially lipid-soluble and watersoluble. One may therefore construct the MMC-SPC-based drug−phospholipid complex to prepare the MMC-loaded PEGlipid-PLA hybrid NPs to protect MMC-SPC from the strong interactions with the plasma proteins and red blood cells. By using a novel reverse micelle−solvent evaporation technique combined with a self-assembly method, the polymer−lipid hybrid NPs were designed to serve as a physical barrier with a core−shell architecture with 3 moieties: (1) an inner reservoir core, which was hydrophobic PLA to be advantageous to load MMC-SPC; (2) an outer protective shell, which was hydrophilic PEG to evade the reticuloendothelial system (RES) while extending the circulation time as a prerequisite to achieve the EPR effect; (3) an interface lipid layer, which was amphiphilic PE to promote the cell uptake and control the drug release. Receptor-targeted NPs are attractive when they target a receptor overexpressed in tumors. This will minimize the systemic distribution of the NPs while facilitating their internalization in the target cells. Folate receptor (FR) is a well-known tumor-associated receptor that is overexpressed in a wide variety of tumors, including ovarian, lung, breast, kidney, and brain cancer cells, but its level is very low in normal tissues.19,20 Thus, we introduced DSPE-PEG-folate (FA) into the MMC-loaded PEG-lipid-PLA hybrid NPs to guide the NPs to the site of interest and increase their therapeutic efficiency. Herein, we reported a novel concept of MMC-based drug− phospholipid complex loaded and FA targeted PEG-lipid-PLA B



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Preparation of FA-PEG-PE-PLA NPs@MMC-SPC. MMCSPC complex was formulated by a solvent evaporation method according to our recently reported procedure.10 FA-PEG-PEPLA NPs@MMC-SPC were prepared by a reverse micelle− solvent evaporation technique combined with a self-assembly method.21,28 Briefly, 1 mL of DCM containing PLA polymer was added to the weighed amount of MMC-SPC, followed by gentle agitation until a clear micellar solution was obtained. The solution was poured into 10 mL aqueous solution containing DPPE, DSPE-PEG-COOH (DSPE-PEG), and DSPE-PEG-FA (mass ratio of DPPE/DSPE-PEG/DSPE-PEG-FA = 2:1:0.6). The mixture was emulsified by sonication in an ice−water bath to form a stable o/w emulsion. The organic solvent was evaporated from the emulsion by magnetic stirring and further removed by evaporation under reduced pressure in a rotary evaporator (N-1001S-W, EYELA, Tokyo, Japan). The NPs were recovered by ultracentrifugation at 20 000 rpm for 20 min at 4 °C, washed with DI water, and lyophilized at −80 °C (3% w/w sucrose as a cryoprotectant) for 24 h. The lyophilized products were stored at 4 °C for future use. Additionally, on one hand, the PEG-PE-PLA NPs@MMC-SPC were prepared using the identical procedure with DSPE-PEG-FA replaced by DSPE-PEG at the equivalent DSPE molar for comparison to evaluate most of the experiments. On the other hand, FA-PEGPE-PLA NPs@MMC was prepared the same way except that MMC-SPC was replaced by MMC at the equivalent MMC mass for comparison to evaluate in vitro drug release. Drug Encapsulation Efficiency. FA-PEG-PE-PLA NPs@ MMC-SPC was dissolved in 1 mL of DCM. After evaporating DCM, 3 mL of the mobile phase (water/methanol (65/35, v/ v)) was added to dissolve MMC and filtrated by 0.45 μm filter membrane. The filtrate was analyzed by a HPLC method for the determination of MMC. A Symmetry C18 column (250 mm × 4.6 mm, 5 μm, Waters Associates, Milford, MA, USA) was used. UV−vis detection was performed at 365 nm. The encapsulation efficiency was calculated by the formula as previously reported.25 Particle Size, Polydispersity Index (PDI), Zeta Potential, and Morphology. The average particle size and PDI of FA-PEG-PE-PLA NPs@MMC-SPC were performed by DLS using a Malvern Zetasizer Nano-ZS (Malvern Instruments, Worcestershire, U.K.). The zeta potential of FA-PEG-PE-PLA NPs@MMC-SPC was estimated by ELS with Zetaplus (Brookhaven Instruments Corporation, Holtsville, NY, USA). Particle size was evaluated by intensity distribution. The morphology of FA-PEG-PE-PLA NPs@MMC-SPC was visualized by SEM (LEO 1530VP, Oberkochen, Germany) operating at 20 kV and TEM (JEM 1400, JEOL, Tokyo, Japan) operating at 200 kV. In Vitro Stability Tests. A short-term stability test. The lyophilized FA-PEG-PE-PLA NPs@MMC-SPC were suspended in PBS (pH 7.4 or pH 7.4 with 10% plasma/heparin) and incubated at 37 °C for 120 h. The particle size was assayed at 24 h intervals by DLS. A long-term stability test. The storage stability of the lyophilized FA-PEG-PE-PLA NPs@MMC-SPC was performed at 4 °C for 90 days according to our previously reported method.10 At 30-day intervals, the particle size, zeta potential, and encapsulation efficiency were assayed by DLS, SLS, and HPLC, respectively. In Vitro Drug Release. The release of MMC from FAPEG-PE-PLA NPs@MMC-SPC was determined by a dialysis technique using a dialysis bag (Mw = 8000−12 000 Da). The

hybrid NPs (Figure 1; more details are shown in Figure S1 in the Supporting Information). The MMC-based drug−phospholipid complex was obtained by complexing MMC with SPC, which not only protect the drug from the rapid degradation to some extent and prevent the premature drug leakage but also increase the lipophilicity, liposolubility, and stability of the drug and block the drug release. More importantly, because of the amphiphilic nature, the MMC-based drug−phospholipid complex was a promising alternative to prepare the reverse micelle-like structure via the self-assembly in the organic solvent containing PLA polymer.21−24 Additionally, with the aid of the pH-sensitive lipid, the PEG-PE-PLA hybrid NPs are stable at physicological pH but instable at endo/lysosomal pH, thereby yielding the NPs potentially suitable for intracellular drug delivery in vivo. Owing to a complex approach,25 PEGmediated passive targeting, FA-mediated active targeting, and pH-triggered MMC release, the novel FA-PEG-PE-PLA NPs@ MMC-SPC could deliver MMC into tumor cells (Figure 1G), presenting the enhanced antitumor activity and reduced side effects of MMC both in vitro and in vivo. To the best of our knowledge, this is the first study in which the water-soluble anticancer drug is exploited with the drug−phospholipid complexation technique to construct the dual-controlled release systems based on the pH-responsive FA-PEG-PE-PLA hybrid NPs for targeted drug delivery.

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EXPERIMENTAL SECTION Materials. Mitomycin C (MMC, purity grade = 99.5%) was purchased from Hisun Pharmaceutical Co., Ltd. (Zhengjiang, China). Soybean phosphatidylcholine (SPC) was provided by Lipoid GmbH (Ludwigshafen, Germany). Poly(D,L-lactide) (PLA, 10 kDa) were provided by Daigang BIO Engineer Co., Ltd. (Shandong, China). 1,2-Dipalmitoyl-sn-glycero-3-phosphoethanolamine (DPPE) was purchased from Corden Pharma Switzerland LLC (Liestal, Switzerland). 1,2-Distearoyl-snglycero-3-phosphoethanolamine-N-carboxy(polyethylene glycol)-2000 (DSPE-PEG-COOH) was purchased from Avanti Polar Lipids (Alabaster, AL, USA). Folate (FA) was purchased from Bio Basic Inc. (Markham, Ontario, Canada). 1,2Distearoyl-sn-glycero-3-phosphoethanolamine-N-folate(polyethylene glycol)-2000 (DSPE-PEG-FA) was synthesized by carbodiimide chemistry as previously reported.26,27 N,N′dicyclohexylcarbodiimide (DCC) and N-hydroxysuccinimide (NHS) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Dimsethyl sulfoxide (DMSO), triethylamine (TEA), tetrahydrofuran (THF), and dichloromethane (DCM) were obtained from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, China). Dialysis bag (Mw = 8000 to 12 000 Da) was ordered from Greenbird Inc. (Shanghai, China). DiD, DiR, and Lysotracker green were from Molecular Probes Inc. (Eugene, OR, USA). Rhodamine phalloidin was from Invitrogen (Carlsbad, CA). Cell Counting Kit-8 (CCK-8) was purchased from Dojindo, Molecular Technologies, Inc. (Rockville, MD, USA). Trypsin−EDTA (0.25%) and penicillin− streptomycin solutions were from Invitrogen. Fetal bovine serum (FBS) was purchased from Gibco Life Technologies (AG, Switzerland). Hoechst 33258, Dulbecco’s Modified Eagle’s Medium (DMEM), and RPMI 1640 medium was from Sigma-Aldrich. Other chemicals were all of analytic grade and were used without further purification. All solvents used in this study were HPLC grade. HeLa cells and A549 cells were provided by American Type Culture Collection (ATCC). C



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further for 2 h. Absorbance at 450 nm was measured using a Bio-Rad Model 680 microplate reader (Richmond, CA, USA). The cells treated with the free MMC or PEG-PE-PLA NPs@ MMC-SPC were used as controls. The cytotoxicity of the PEGPE-PLA NPs or FA-PEG-PE-PLA NPs was tested as the method above, and the concentration of the NPs varied from 4 to 500 μg/mL. Hemolysis Assay. Freshly collected whole blood (0.8 mL) was then centrifuged at 10 000 rpm for 5 min at 4 °C. The resulting red blood cells were isolated and washed five times with sterile isotonic PBS. The washed red blood cells were suspended in 7.5 mL of PBS in an ice bath, and 0.5 mL was added to 0.5 mL of MMC drug-free NPs (PEG-PE-PLA NPs or FA-PEG-PE-PLA NPs) suspended in PBS at a concentration of 8, 40, 200, or 1000 μg/mL to make the final concentration of 4, 20, 100, or 500 μg/mL. The NPs and red blood cells were vortexed briefly, left at 37 °C for 60 min, then centrifuged at 10 000 rpm for 5 min. The supernatant was transferred to a plate, and the absorbance value of hemoglobin at 577 nm was determined with a reference wavelength of 655 nm. The same volume of PBS was used as the negative control, and water was the positive control. The percentage of hemolysis was calculated by the formula as previously reported.30 The results were included in Figure S5 in the Supporting Information. In Vivo Pharmacokinetics.25 All the animal procedures complied with the guidelines of the Xiamen University Institutional Animal Care and Use Committee. The experiments were performed on adult male SD rat weighing 200 ± 20 g (mean ± SD) from Shanghai Laboratory Animal Center. To evaluate the circulation half-life of FA-PEG-PE-PLA NPs@ MMC-SPC, FA-PEG-PE-PLA NPs@MMC-SPC at 4 mg/kg (MMC-eq dose) were injected into the tail vein of the rat (n = 4). Three hundred microliters of blood were collected into heparinized tubes at the predesigned time following the injection. The heparinized blood samples were centrifuged immediately at 3000 rpm for 15 min to harvest the plasma and stored at −20 °C prior to analysis. MMC were extracted from the plasma by deproteinization using ethyl acetate followed by centrifugation at 12 000 rpm for 10 min. The PEG-PE-PLA NPs@MMC-SPC and free MMC at MMC-eq dose were also tested in parallel as controls. The plasma MMC concentrations were determined by LC−MS/MS method with 4-aminoacetophenone as internal standard. The chromatographic separations were acquired on a Waters Acquity HPLC system (Waters Corporation, Milford, MA, USA), equipped with a Waters e2695 separations module, a Waters 2998 photodiode array detector, and a Hypersil GOLD C18 column. HPLC grade acetonitrile/water (25/75, v/v) was used as the mobile phase. The compounds were analyzed by multiple reaction monitoring of the transitions of m/z 335 → 242 for MMC and m/z 136 → 94 for 4-aminoacetophenone, respectively. The pharmacokinetic parameters such as elimination half-life (t1/2), area under the curve (AUC), volume of distribution (Vd), and clearance (CL) were calculated by fitting the blood−drug pharmaceutical concentrations to a two-compartment model using WinNonlin Professional Edition Version 2.1 (Pharsight Corporation, Mountain View, California). Ex Vivo Tumor Targeted Imaging.25 DiR, a near-infrared fluorescent probe, was encapsulated into the MMC-loaded FAPEG-PE-PLA NPs. FA-PEG-PE-PLA NPs@MMC-SPC@DiR (0.2 mL) was injected into mice bearing H22 tumor via tail vein. At 12 h postinjection, the mice were sacrificed. The tumor and major organs (spleen, liver, kidney, lung, and heart) were

lyophilized NPs were dialyzed against pH-adjusted PBS (pH 7.4, pH 7.4 with 10% FBS, pH 6.5, pH 6.5 with 10% FBS, pH 5.5, and pH 5.5 with 10% FBS) at 37 °C. At the predesigned time, 2 mL of the release medium was completely withdrawn and subsequently replaced with the 2 mL of fresh PBS. The release of MMC was determined by a HPLC method as described above. In addition, the release of the free MMC, PEG-PE-PLA NPs@MMC-SPC, or FA-PEG-PE-PLA NPs@ MMC was investigated for comparison. Cell Culture. HeLa cells were cultured in FA-deficient DMEM supplemented with 10% fetal bovine serum (FBS) and 1% penicillin−streptomycin. A549 cells were cultured in RPMI 1640 medium under similar conditions. The two cell lines have different levels of FA receptor expression. In particular, HeLa cells are FA receptor positive and A549 cells are FA receptor negative. All of the cells were cultivated in a humidified atmosphere containing 5% CO2 at 37 °C. In Vitro Cellular Uptake. To facilitate the observation of cellular uptake, the hydrophobic red fluorescent DiD were added to the oil phase and encapsulated into FA-PEG-PE-PLA NPs@MMC-SPC. HeLa cells or A549 cells were seeded at a density of 1 × 105 cells per well in 6-well plates with their specific cell culture medium. The cells were incubated at 37 °C and 5% CO2 for 24 h. One hundred microliters of the PEG-PEPLA NPs@MMC-SPC or FA-PEG-PE-PLA NPs@MMC-SPC at DiD-eq dose was added and incubated further for 8 h. The cells were washed with PBS, fixed with 4% paraformaldehyde, and stained with Hoechst 33258. The cells were observed using a Leica TCS SP5 LCSM (Leica Microsystems, Mannheim, Germany). Flow Cytometry Tests. HeLa cells were seeded in 6-well plates with a density of 1 × 105 cells/mL and incubated for 24 h, and then the original medium was replaced with the PEGPE-PLA NPs@MMC-SPC@DiD or FA-PEG-PE-PLA NPs@ MMC-SPC@DiD at DiD-eq dose. The cells were incubated for the predesigned time at 37 °C, and then washed with PBS and harvested by 0.25% trypsin−EDTA. The harvested cells were suspended in PBS and centrifuged at 1000 rpm for 5 min at 4 °C. After two cycles of washing and centrifugation, the cells were resuspended in PBS and performed by a Beckman Coulter EPICS XL flow cytometer (Beckman Coulter, CA, USA). The fluorescence intensity of the cells was calculated by CellQuest software (Becton-Dickinson, San Jose, CA, USA). Subcellular Localization. HeLa cells were incubated for 24 h and then cultured with the PEG-PE-PLA NPs@MMC-SPC@ DiD or FA-PEG-PE-PLA NPs@MMC-SPC@DiD for 12 h at 37 °C. The cells were imaged using a LCSM. Lysotracker green was used to stain lysosomes. Hoechst 33258 was used to stain nuclei. Additionally, to understand the intracellular distribution of the drug inside HeLa cells, the FITC-MMC was prepared for use as a fluorescence probe.29 HeLa cells were incubated for 12 h and then cultured with the PEG-PE-PLA NPs@MMC-SPC or FA-PEG-PE-PLA NPs@MMC-SPC for 12 h at 37 °C. Rhodamine phalloidin was used to stain cytoskeleton. Hoechst 33258 was used to stain nuclei. In Vitro Cell Viability. The cytotoxicity of the MMC-based drug−phospholipid complex-loaded PEG-PE-PLA NPs was evaluated by CCK-8 assay. HeLa cells and A549 cells were seeded at a density of 1 × 104 cells per well in 96-well plates, preincubated for 24 h, and then incubated with FA-PEG-PEPLA NPs@MMC-SPC for 24 h. FA-PEG-PE-PLA NPs@ MMC-SPC with MMC dose ranged from 3 to 12 μg/mL. After incubation for 24 h, 20 μL of CCK-8 was added and incubated D



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Figure 2. Particle size and zeta potential of FA-PEG-PE-PLA NPs@MMC-SPC at different weight ratios of MMC-SPC to PLA polymer.

harvested from all 4 groups were fixed in 10% buffered formalin, embedded in paraffin, sectioned (4 μm), stained with hematoxylin and eosin (H&E), and examined using a digital microscopy system. Statistical Analysis. All experiments were repeated at least three times. All data were expressed as mean ± SD. Statistical tests were performed by the Student’s t test and one-way ANOVA. The statistical difference was considered to be significant when the P value was less than 0.05.

excised, followed by washing the surface with 0.9% NaCl for the ex vivo imaging of DiR fluorescence using a Maestro in vivo imaging system (Cambridge Research & Instrumentation, Woburn, MA, USA). The mice treated with 0.2 mL of the PEG-PE-PLA NPs@MMC-SPC@DiR at DiR-eq dose were used for comparison. The resulting data can be used to identify, separate, and remove the contribution of autofluorescence in analyzed images by the Carestream Molecular Imaging Software. In Vivo Biodistribution.31,32 To assess the biodistribution of MMC formulations, the H22 tumor-bearing mice were intravenously injected via tail vein with the PEG-PE-PLA NPs@MMC-SPC or FA-PEG-PE-PLA NPs@MMC-SPC at MMC-eq dose. The mice were sacrificed at the predesigned time after administration, and the spleen, liver, kidney, lung, heart, and tumor were excised and washed with cold 0.9% NaCl to remove surface blood. Tissue samples were then homogenized and mixed with the mixture of acetonitrile/ water (25/75, v/v). MMC was extracted using ethyl acetate followed by centrifugation at 12 000 rpm at 4 °C for 10 min. The content of MMC in supernatant was then measured by a HPLC method and expressed as percentage of the injected dose pergram of tissue (%ID/g tissue). In Vivo Antitumor Effect. Kunming mice aged 4−5 weeks (clean class, 18−22 g) were supplied by Shanghai Laboratory Animal Center and used in this study. Subcutaneous tumors were established in the mice by subcutaneous inoculation of 1 × 106 H22 cells in the right axillary region of mice before the treatment. The H22 tumor bearing mice were randomly divided into 4 groups (10 mice per group): group 1 for 0.9% NaCl, group 2 for MMC injection, group 3 for PEG-PE-PLA NPs@ MMC-SPC, and group 4 for FA-PEG-PE-PLA NPs@MMCSPC. The mice were intravenously administrated at 4 mg/kg (MMC-eq dose) every 2 days for 4 times. Each mouse was earmarked and followed individually throughout the whole experiment. The width and length of the tumors and the body weight of mice were measured every day until the animals were terminated. Tumor volume (V) was calculated by the formula as previously reported.33 The mice were terminated on day 9. The tumors were excised and then weighed. The H22 tumor

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RESULTS AND DISCUSSION

Preparation of FA-PEG-PE-PLA NPs@MMC-SPC. Initially, MMC was complexed with SPC by a solvent evaporation method to prepare MMC-SPC. After the introduction of the organic solvent containing polymer into this complex, MMC was solubilized within the organic phase by the formation of the reverse micelles and subsequently loaded into the NPs by a reverse micelle−solvent evaporation technique combined with a self-assembly method. Besides, the relatively inefficient encapsulation of watersoluble small molecule drug into the PLA NPs may be a great challenge using traditional methods. For this reason, we introduced MMC-SPC into the FA-PLA-PE-PEG NPs, in which the fat-soluble tail of SPC was outward and MMC was protected inward. Moreover, the drug−phospholipid complex represented an attractive drug delivery approach to improve the therapeutic index of water-soluble cytotoxic drugs. However, the FA-PLA-PE-PEG NPs extravasated from the leaky vasculature with poor lymphatic drainage into the tumor interstitium via the EPR effect and rapidly internalized by the tumor cells via the FA receptor-mediated endocytosis. Characterization of FA-PEG-PE-PLA NPs@MMC-SPC. The particle size and zeta potential of FA-PEG-PE-PLA NPs@ MMC-SPC were performed at the weight ratio of drug− phospholipid complex to PLA polymer. As the weight ratio of MMC-SPC to PLA increased, the particle size of the MMCloaded NPs increased and the zeta potential of those decreased. We inferred a model of the distribution of drug−complex in FA-PEG-PE-PLA NPs@MMC-SPC (Figure 2). In the case of

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low MMC-SPC/PLA weight ratio, all of MMC-SPC distributed in the PLA with their lipophilic parts oriented toward the organic phase. In the case of the high one, some MMC-SPC was transported to the interface of the oil/water with their hydrophilic parts oriented toward the water phase. Moreover, the excess MMC-SPC might participate in the self-organization of the large size, vesicle-like structures (typical particle size is about 50−1000 nm with a broad distribution).16 Consequently, FA-PEG-PE-PLA NPs@MMC-SPC with a MMC-SPC/PLA weight ratio of 20% was applied for the following study. It was shown that FA-PEG-PE-PLA NPs@MMC-SPC presented a small particle size (215.6 ± 5.1 nm) with a PDI of 0.143 ± 0.021 (Figure 1B) and a spherical shape (Figure 1D−F), which might be favorable for the passive tumor targeting by EPR effect.34,35 Besides, FA-PEG-PE-PLA NPs@MMC-SPC exhibited a zeta potential of −25.88 ± 2.39 mV (Figure 1C), indicating a good dispersion stability. At the physiological environment, the plasma protein and red blood cells would have a net negative charge, which could reduce the nonspecific interaction with the negatively charged FA-PEG-PE-PLA NPs@MMC-SPC. The result was also benefited by reducing their rapid elimination from the blood circulation and promoting the passive accumulation at the tumor sites.25,36 Additionally, the coexistence of carboxyl and amino groups as “zwitterions” was a vital structural characteristic of the NPs since it could weaken the rapid RES clearance and facilitate the passive cell uptake.37 In Vitro Stability. An index to evaluate the stability of the NPs was their particle size changes in PBS or plasma. No obvious particle size changes of FA-PEG-PE-PLA NPs@MMCSPC over 120 h is shown in Figure S2 in the Supporting Information. Another major advantage of these MMC formulations was their excellent stability, which was a problem encountered with the MMC loaded liposomes. We inferred that the electrostatic repulsion, steric repulsion, and structural longevity of FA-PEG-PE-PLA NPs@MMC-SPC contributed to their effective physiological stability against ionic strength and protein adsorption. Once intravenously administrated, these NPs had potential to withstand the biological environment and promote the blood persistence. However, a major concern in pharmaceutical industry is the long-term storage stability. At 4 °C in 90 days, the high storage stability in terms of the particle size, zeta potential, and encapsulation efficiency of the lyophilized FA-PEG-PE-PLA NPs@MMC-SPC was observed in Figure S3 in the Supporting Information. It was suggested that the introduction of the drug−phospholipid complex and the polymer−lipid hybrid NPs structure greatly assisted in protecting the drug against the rapid degradation over a long period of time. In Vitro Drug Release. In vitro release profiles of MMC from the PEG-PE-PLA NPs@MMC-SPC and FA-PEG-PEPLA NPs@MMC-SPC are shown in Figure 3A. Outperforming the free drug, the two drug delivery systems displayed a biphasic pattern that was characterized by a burst of drug release followed by a sustained drug transport. The first one was explained by the presence of the NP surface-associated drug−phospholipid complex. The second one was predominantly driven by a diffusion-controlled mechanism: initially, MMC was dissociated from the complex effectively and was homogeneously dispersed within the internal core, and subsequently, the free MMC diffused from the polymer−lipid hybrid NPs. Since MMC was a time-dependent drug, a reservoir or depot effect of the MMC-loaded NPs indicated

Figure 3. MMC release profiles of different formulations in PBS at pH 7.4. (A) In vitro release profiles of MMC from the free MMC, PEGPE-PLA NPs@MMC-SPC, and FA-PEG-PE-PLA NPs@MMC-SPC in PBS at pH 7.4 (mean ± SD, n = 3). (B) In vitro release profiles of MMC from FA-PEG-PE-PLA NPs@MMC-SPC versus FA-PEG-PEPLA NPs@MMC in PBS at pH 7.4 (mean ± SD, n = 3).

that these NPs were anticipated to be safe, long-lasting, and effective MMC delivery systems. Additionally, the encapsulation efficiency of FA-PEG-PEPLA NPs@MMC-SPC (94.5 ± 3.2%) was higher than that of FA-PEG-PE-PLA NPs@MMC (48.7 ± 6.2%). It was particularly noted that the strategy of drug−phospholipid complex facilitated the smart and efficient preparation of MMC-based drug−phospholipid complex-loaded NPs with a high encapsulation level. This result is a significant improvement on the traditional drug-loaded NPs. On the basis of the encouraging result, we hypothesized that during the NPs preparation, MMC-SPC complex (used to bridge the gap between MMC and SPC) and lipophilic SPC (used to act as a molecular fence) may serve the function of retaining the loading drug, providing a significantly better protection of the drug. To test our hypothesis, we investigated the in vitro release profile of these two drug delivery systems (Figure 3B), the even more sustained drug release from FAPEG-PE-PLA NPs@MMC-SPC proved that SPC and its strong physical interaction with MMC in drug−phospholipid complex played a vital role against both water penetration inward and drug diffusion outward. In a word, we believed that the F



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introduction of drug−phospholipid complex could endow the drug with increased hydrophobicity not only to increase the drug encapsulation efficiency but also to prolong the drug release, which might actually contribute to a longer circulation in the bloodstream. In Vitro Cellular Uptake. To demonstrate that FA-PEGPE-PLA NPs@MMC-SPC could use their function to specifically target cancer cells and induce cell apoptosis/ death, we chose two cell lines of HeLa cells (FA receptor positive) and A549 cells (FA receptor negative) and respectively incubated them with FA-PEG-PE-PLA NPs@ MMC-SPC. As revealed in Figure 4A, the red fluorescence distributed in the cytoplasma, indicating that FA-PEG-PE-PLA NPs@MMC-SPC had been uptaken by HeLa cells after incubation for 8 h. Through staining the nuclei with blue fluorescence probe, it was found that the amount of FA-PEGPE-PLA NPs@MMC-SPC uptaken by HeLa cells was significantly more in comparison with A549 cells, as demonstrated by the weak red fluorescence (Figure 4Ab,Ad). The reason was that the overexpressed FA receptor residing on the outer surface of HeLa cells membrance facilitated the effective recognition of FA-PEG-PE-PLA NPs@MMC-SPC, resulting in a dramatically increased cellular uptake. Similarly, between FA-PEG-PE-PLA NPs@MMC-SPC and PEG-PE-PLA NPs@MMC-SPC, the significant difference of their cellular uptake in HeLa cells (Figure 4Aa,Ab) but insignificant difference of that in A549 cells (Figure 4Ac,Ad) could be explained by their distinct internalization mechanisms.38 The result further confirmed the cell-specific effect of FA-PEG-PE-PLA NPs@MMC-SPC, which could contribute to selectively increasing the anticancer activity. Flow Cytometry Tests. The quantitative analysis of the cellular uptake was determined by flow cytometry tests. The intracellular fluorescence intensity increased as the incubation time was extended from 1 to 4 h (Figure 4B). Also, FA-PEGPE-PLA NPs@MMC-SPC@DiD exhibited the higher cellular uptake compared with PEG-PE-PLA NPs@MMC-SPC@DiD at DiD-eq dose. PEG-PE-PLA NPs@MMC-SPC could be transported into cells across the lipid bilayer structured cellular membranes by a nonspecific absorption depending on the particle size and the presence of lipids on the NPs surface. On the contrary, FA-PEG-PE-PLA NPs@MMC-SPC could be uptaken by cells via a selective ligand−receptor interaction. To this end, the quantitative result correlated well with the qualitative results giving further proof of the improved cellular uptake efficiency of FA-PEG-PE-PLA NPs@MMC-SPC. Versatile Therapeutic Drug Delivery. To investigate the state of FA-PEG-PE-PLA NPs@MMC-SPC in different buffers, the NPs were immersed in PBS (pH 7.4, pH 7.4 with 10% FBS, pH 6.5, pH 6.5 with 10% FBS, pH 5.5, and pH 5.5 with 10% FBS) for 4 h at 37 °C. The particle size of the NPs increased as pH decreased (Figure 5A,B). As pH decreased from 7.4 to 6.5 and 5.5, initially, DPPE was partly protonated, subsequently the intra- and intermolecular hydrophobic interactions might become weaker, and lastly the NPs were disintegrated and aggregated into a larger formation.39 Additionally, the drug release behavior of FA-PEG-PE-PLA NPs@MMC-SPC was time- and pH-dependent (Figure 5C). In agreement with the previous reports,40,41 these results indicated that phosphatidylethanolamine was pH-sensitive, which was favorable for achieving controlled drug release.42,43 Such a pH-triggered release behavior of MMC from FA-PEGPE-PLA NPs@MMC-SPC presented the potential use in drug

Figure 4. Cellular uptake of FA-PEG-PE-PLA NPs@MMC-SPC. (A) Qualitative cellular uptake analysis. LCSM images of (Aa, Ab) HeLa cells or (Ac, Ad) A549 cells incubated with (Aa, Ac) PEG-PE-PLA NPs@MMC-SPC and (Ab, Ad) FA-PEG-PE-PLA NPs@MMC-SPC. (B) Quantitative cellular uptake analysis. Flow cytometer tests of HeLa cells incubated with PEG-PE-PLA NPs@MMC-SPC and FA-PEG-PEPLA NPs@MMC-SPC for 1, 2, and 4 h. *P < 0.05.

delivery for the antiproliferative effect, due to the intensive release of MMC in an acidic compartment inside the cells while limiting the premature release in the blood circulation. In vitro tests were also performed in 10% FBS to preferably simulate the in vivo situation. It was noteworthy that the release rate was increased in 10% FBS (Figure 5Cb) compared to PBS (Figure 5Ca). This finding suggested that the presence of enzymes in the FBS might facilitate this release, which was in agreement with a recent study.25 HeLa cells were treated with FA-PEG-PE-PLA NPs@MMCSPC@DiD for 12 h, and the intracellular delivery is shown in Figure 5D. The NPs were found to be localized in the G



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Figure 5. Time- and pH-dependent MMC release of FA-PEG-PE-PLA NPs@MMC-SPC. (A) Particle size of FA-PEG-PE-PLA NPs@MMC-SPC in PBS at different pH for 4 h: (a) 7.4, (b) 6.5, and (c) 5.5. (B) Particle size of FA-PEG-PE-PLA NPs@MMC-SPC in PBS at different pH with 10% FBS for 4 h: (a) 7.4, (b) 6.5, and (c) 5.5. (C) In vitro release profiles of MMC from FA-PEG-PE-PLA NPs@MMC-SPC in (a) PBS at different pH and (b) PBS at different pH with 10% FBS (mean ± SD, n = 3). (D) Lysosomal colocalization of FA-PEG-PE-PLA NPs@MMC-SPC@DiD in HeLa cells after incubation for 12 h at 37 °C was observed by LCSM through (a) the DiD channel and (b) the LysoTracker green channel. The overlap of the images in panels a and b are shown in panel c. The nuclei were stained with Hoechst 33258 (blue). The DiD loaded FA-PEG-PE-PLA NPs@ MMC-SPC appeared in red. Lysotracker Green was used to identify the lysosomes.

On one hand, FA-PEG-PE-PLA NPs@MMC-SPC revealed a more potent cytotoxicity on HeLa cells (target cells) but a lesser one on A549 cells (nontarget cells) compared to PEGPE-PLA NPs@MMC-SPC at MMC-eq dose (3, 6, 9, and 12 μg/mL) (Figure 6), which could be explained by the selective targeting effect of FA-PEG-PE-PLA NPs@MMC-SPC through FA receptor-mediated endocytosis (see Figure 4A). On the other hand, FA-PEG-PE-PLA NPs@MMC-SPC induced a higher cytotoxicity on target cells compared to the free MMC at a low MMC-eq dose (3 μg/mL) (Figure 6), while the cytotoxicity of FA-PEG-PE-PLA NPs@MMC-SPC was not significantly different from that of the free MMC at a high MMC-eq dose (6, 9, and 12 μg/mL), this phenomenon could be explained by that the drug penetration rate may be dependent on both NPs internalization and drug concentration gradient between the internal and external environments of the cell membrane. It should be noted that in the case of a low concentration of MMC, the NPs internalization played a rather important role in the the drug penetration rate. In the case of a high concentration of MMC, however, the free drugs can be rapidly transported into cells by passive diffusion due to the high concentration gradient and instantaneously inhibit the cell growth without the drug release process.41 To the end, the delayed drug release behavior of the PEG-PE-PLA NPs@ MMC-SPC and FA-PEG-PE-PLA NPs@MMC-SPC can result in the progressive increase of intracellular drug level for target cell death. More noticeably, the targeting efficacy of FA-PEG-

lysosomes, as evidenced by the yellow spots in the merged image obtained from the images of the NPs (red) and lysosomes (green). This indicated that the NPs were internalized via the endocytosis pathway into the lysosomes. However, a part of NPs were no longer localized in the lysosomes, suggesting successful escape from the lysosomes, which was also in agreement with a recent study.44 Thus, it was reasonable to believe that FA-PEG-PE-PLA NPs@MMC-SPC entered into the lysosomes (pH 4.0−5.0) inside HeLa cells, and MMC could easily escape from the lysosomes and eventually enter the nuclei. To this end, as illustrated in Figure 1G, once the NPs reached the target site with maximizing the drug protection and reducing the systemic toxicity, there would be a selective cellular uptake (see Figure 4), effective lysosomal escape (see Figure 5D), and a sustained/controlled drug release (see Figures 3 and 5C), indicative of an effective and long-lasting effect of MMC in the nuclei (Figure S4 in the Supporting Information). In Vitro Cytotoxicity Assays. To evaluate the cytotoxicity of the MMC-based drug−phospholipid complex-loaded NPs, we chose two cancer cell lines: HeLa cells and A549 cells. No significant toxic effect against HeLa cells or A549 cells was found in the drug-free NPs (see Figure S5 in the Supporting Information), indicating that the PEG-PE-PLA NPs and FAPEG-PE-PLA NPs possessed good biocompatibility. H



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Figure 7. Pharmacokinetics profiles after intravenous injection of the free MMC, PEG-PE-PLA NPs@MMC-SPC, and FA-PEG-PE-PLA NPs@MMC-SPC in rats (mean ± SD, n = 3).

free MMC, with a longer elimination half-life (t1/2), higher area under the curve (AUC), and a substantially lower value for clearance (CL). The results suggested the superior blood retention of the MMC-loaded FA-PEG-PE-PLA NPs, which may play a decisive role in promoting the accumulation in tumor via the EPR effect. The retarded clearance of plasma MMC associated with the reduced macrophage uptake in the FA-PEG-PE-PLA NPs@MMC-SPC group compared to the free MMC group might be explained by their in vitro controlled/sustained drug release based on the hybrid NPs structure (see Figures 3 and 5C) and in vivo prolonged circulation based on the PEGylated nanoscaled delivery systems. Ex Vivo Tumor Targeted Imaging. DiR, a near-infrared fluorescent dye, was used to evaluate the biodistribution of the NPs. The tumor model was established by inoculating H22 cells in the right axillary region of BALB/c nude mice. As shown in Figure 8A, the tumor fluorescent intensity in the mice treated with FA-PEG-PE-PLA NPs@MMC-SPC@DiR was higher than PEG-PE-PLA NPs@MMC-SPC@DiR. However, FA-PEG-PEPLA NPs@MMC-SPC@DiR showed the decreased fluorescent intensity in the liver compared to PEG-PE-PLA NPs@MMCSPC@DiR. These results indicated the less RES uptake and more tumor accumulation of FA-PEG-PE-PLA NPs@MMCSPC@DiR over PEG-PE-PLA NPs@[email protected],46 As clearly shown in Figure 8B, the fluorescent intensity from the tumor in FA-PEG-PE-PLA NPs@MMC-SPC@DiR group was significantly greater than that in the PEG-PE-PLA NPs@ MMC-SPC@DiR group, which was a further evidence of the tumor targeting of FA-PEG-PE-PLA NPs@MMC-SPC. Quantitative analysis for main tissue distribution of MMC in H22 tumor-bearing mice after 24 h injection of the PEG-PEPLA NPs@MMC-SPC or FA-PEG-PE-PLA NPs@MMC-SPC were shown in the inset of Figure 8B (in vivo biodistribution of MMC in the tumor and normal tissues was also shown in Figure S7 in the Supporting Information). The MMC concentration within the tumor of mice treated with FAPEG-PE-PLA NPs@MMC-SPC was significantly higher than that of PEG-PE-PLA NPs@MMC-SPC. In addition, MMC was nearly well-distributed in the normal tissues even after 24 h of injection. All findings suggested that these nanoscaled MMC delivery systems permitted specific tumor targeting efficiency

Figure 6. Cell viability of (A) HeLa cells or (B) A549 cells incubated with the free MMC, PEG-PE-PLA NPs@MMC-SPC, and FA-PEGPE-PLA NPs@MMC-SPC at different concentrations (3, 6, 9, and 12 μg/mL) for 24 h (mean ± SD, n = 6).

PE-PLA NPs@MMC-SPC would increase the tumor accumulation of the more amounts of drugs loaded within the drug delivery systems during circulation in vivo, which was confirmed by ex vivo tumor targeted imaging (discussed below). In Vivo Pharmacokinetics. No significant hemolytic effects for hemolysis analysis were observed even at the highest NP concentration of 0.5 mg/mL in PBS (Figure S6 in the Supporting Information), this is a proof to confirm that the PEG-PE-PLA NPs or FA-PEG-PE-PLA NPs possessed excellent blood compatibility ( 0.05 vs PEG-PE-PLA NPs@MMC-SPC.

NPs yielded the more effective tumor growth inhibition compared to the free MMC. Notably, there was no significant difference between the tumor growth inhibition of the free MMC and PEG-PE-PLA NPs@MMC-SPC (p > 0.05) before the fourth treatment, implying that the PEG-PE-PLA NPs@ MMC-SPC had a superior anticancer effect compared with the free MMC, but this advantage was only apparent after multiple dose intravenous administration over a long period of time. At the end of experiment, the tumors were excised and weighed. As shown in Figure 9B, it was found that the MMC-loaded NPs had superior therapeutic efficacy compared with free MMC (p < 0.05). Moreover, the targeting properties greatly improved the therapeutic efficacy of the MMC-loaded polymer−lipid hybrid NPs because of the higher accumulation at the tumor site and the greater cellular uptake by the tumor cells.28 Although the cytotoxicity of the free MMC was higher than (or similar to) that of the MMC-loaded NPs in vitro, the MMCloaded NPs presented an enhanced therapeutic effect against H22 tumor probably due to the plasma stability (see Figure S2 in the Supporting Information), the EPR effect (see Figure 8A), extended blood circulation time (see Figure 7), receptormediated targeting (see Figure 4), and sustained drug release (see Figures 3 and 5C). Additional evidence of the enhanced anticancer activity of the MMC-loaded NPs was analyzed by histological examination (Figure 9D). The abundant tumor cells were observed in the tumor treated with 0.9% NaCl. On the contrary, the extensive dead cells without nuclei were observed in the tumor treated with the free MMC or MMC-loaded NPs. Moreover, it should be noted that FA-PEG-PE-PLA NPs@MMC-SPC yielded greater effectiveness in inducing the cell apoptosis/death and reducing the cell proliferation, which was explained by the size effect-mediated enhanced accumulation at the tumor site and the targeting effect-mediated increased cellular uptake in the tumor cells. For any drug delivery systems, the systemic toxicity that is usually encountered in the free MMC-mediated treatment8,9 should be considered to ensure safety and effectiveness. In this work, the administration of the free MMC resulted in the listlessness/laziness and severe body weight loss of mice (Figure 9C), indicative of the undesirable side effects of chemotherapy. Anyway, the toxicity caused by the free MMC was the principal reason to induce the mice death (see Table S1 in the Supporting Information). On the contrary, no obvious side effects were shown in the mice treated with the MMCloaded NPs. Also, all the mice treated with the MMC-loaded NPs were alive during the experimental period. Overall, it was indicated that these NPs with the superior anticancer effects as well as lower toxicity would greatly improve the efficacy of quality of life therapy.

Figure 8. Ex vivo tumor targeted imaging of FA-PEG-PE-PLA NPs@ MMC-SPC. (A) Ex vivo fluorescence imaging of normal organs and tumor excised from H22 tumor-bearing mice treated with (a) PEGPE-PLA NPs@MMC-SPC@DiR or (b) FA-PEG-PE-PLA NPs@ MMC-SPC@DiR at 12 h postinjection (S, spleen; Li, liver; T, tumor; K, kidney; Lu, lung; and H, heart). (B) Quantificative tumor target characteristics of PEG-PE-PLA NPs@MMC-SPC@DiR and FAPEG-PE-PLA NPs@MMC-SPC@DiR at 12 h postinjection (mean ± SD, n = 3). Inset: Quantitative analysis for main tissue distribution of MMC in H22 tumor-bearing mice treated with PEG-PE-PLA NPs@ MMC-SPC or FA-PEG-PE-PLA NPs@MMC-SPC at 24 h postinjection (mean ± SD, n = 3). *P < 0.05.

and extended blood circulation time, which could contribute to selectively increasing the therapeutic efficiency. In Vivo Anticancer Effects. To evaluate the in vivo antitumor effects, we treated Kunming mice bearing H22 tumor with 0.9% NaCl, free MMC, PEG-PE-PLA NPs@MMC-SPC, and FA-PEG-PE-PLA NPs@MMC-SPC. Compared to the mice treated with 0.9% NaCl as control, the tumor volume in the mice treated with the free MMC or MMC-loaded NPs was significantly decreased after a schedule of multiple doses (Figure 9A), indicating the significantly effective tumor growth inhibition. It was encouraging to find that the MMC-loaded

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CONCLUSIONS The present research proposed and evaluated the cell-specific and pH-responsive MMC-based drug−phospholipid complex J



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Figure 9. Antitumor effects of different formulations (0.9% NaCl, free MMC, PEG-PE-PLA NPs@MMC-SPC, and FA-PEG-PE-PLA NPs@MMCSPC) in mice bearing H22 tumor. *P < 0.05. (A) Tumor volume changes of the tumor-bearing mice during the treatment. The arrows represent the day on which the intravenous administration was performed. (B) Quantitative results of the tumor weight excised from the tumor-bearing mice sacrificed on day 9. (C) Body weight changes of the tumor-bearing mice during the treatment. (D) Histological section of the tumor of the mice after the treatment.

MMC-SPC or FA-PEG-PE-PLA NPs@FITC-MMC-SPC, effect of the MMC drug-free NPs on cell viability, hemolysis assay of the MMC drug-free NPs, and survival study. This material is available free of charge via the Internet at http:// pubs.acs.org.

loaded PEG-lipid-PLA hybrid nanoscaled drug delivery systems by loading MMC−phospholipid complex within FA-PEG-PEPLA NPs in vitro and in vivo. FA-PEG-PE-PLA NPs@MMCSPC presented an appropriate particle size and zeta potential, a good stability, high encapsulation efficiency, sustained/controlled release, and steady-state pharmacokinetics, which could improve drug efficacy and reduce drug toxicity. Importantly, these NPs can escape from the lysosomes of cancer cells and ensure the nuclear delivery of anticancer drug. Besides, FAPEG-PE-PLA NPs@MMC-SPC exhibited highly selective tumor accumulation and greatly superior therapeutic efficiency in vivo. We therefore suggested that FA-PEG-PE-PLA NPs@ MMC-SPC were robust and attractive drug delivery systems for effective chemotherapy against FA receptor overexpressing cancer.

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AUTHOR INFORMATION

Corresponding Author

*(Z.H.) E-mail: [email protected]. Tel: 0865922189650. Fax: 0865922183058. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This study was supported by the Natural Science Foundation of China (Grant No. 31271071 and 81000660).

ASSOCIATED CONTENT

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S Supporting Information *

ABBREVIATIONS USED MMC, mitomycin C; DLS, dynamic light scattering; SLS, static light scattering; PDI, polydispersity index; SEM, scanning electron microscopy; LCSM, laser confocal scanning microscopy; HPLC, high-performance liquid chromatography

More details of schematic illustration of FA-PEG-PE-PLA NPs@MMC-SPC, in vitro stability of FA-PEG-PE-PLA NPs@ MMC-SPC, in vitro storage stability of the lyophilized FAPEG-PE-PLA NPs@MMC-SPC, LCSM images of HeLa cells after incubation for 12 h with the PEG-PE-PLA NPs@FITCK



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(18) Alam, M. A.; Al-Jenoobi, F. I.; Al-Mohizea, A. M. Commercially bioavailable proprietary technologies and their marketed products. Drug Discovery Today 2013, 18, 936−949. (19) Low, P. S.; Henne, W. A.; Doorneweerd, D. D. Discovery and development of folic-acid-based receptor targeting for imaging and therapy of cancer and inflammatory diseases. Acc. Chem. Res. 2008, 41, 120−129. (20) Weitman, S. D.; Lark, R. H.; Coney, L. R.; Fort, D. W.; Frasca, V.; Zurawski, V. R., Jr.; Kamen, B. A. Distribution of the folate receptor GP38 in normal and malignant cell lines and tissues. Cancer Res. 1992, 52, 3396−3401. (21) Cui, F.; Shi, K.; Zhang, L.; Tao, A.; Kawashima, Y. Biodegradable nanoparticles loaded with insulin-phospholipid complex for oral delivery: preparation, in vitro characterization and in vivo evaluation. J. Controlled Release 2006, 114, 242−250. (22) Peng, Q.; Zhang, Z. R.; Gong, T.; Chen, G. Q.; Sun, X. A rapidacting, long-acting insulin formulation based on a phospholipid complex loaded PHBHHx nanoparticles. Biomaterials 2012, 33, 1583− 1588. (23) Wei, W.; Shi, S.-J.; Liu, J.; Sun, X.; Ren, K.; Zhao, D.; Zhang, X.N.; Zhang, Z.-R.; Gong, T. Lipid nanoparticles loaded with 10hydroxycamptothecin−phospholipid complex developed for the treatment of hepatoma in clinical application. J. Drug Targeting 2010, 18, 557−566. (24) Zhou, C.; Xia, X.; Liu, Y.; Li, L. The preparation of a complex of insulin-phospholipids and their interaction mechanism. J. Pept. Sci. 2012, 18, 541−548. (25) Li, M.; Song, W.; Tang, Z.; Lv, S.; Lin, L.; Sun, H.; Li, Q.; Yang, Y.; Hong, H.; Chen, X. Nanoscaled poly( L -glutamic acid)/ doxorubicin-amphiphile complex as pH-responsive drug delivery system for effective treatment of nonsmall cell lung cancer. ACS Appl. Mater. Interfaces 2013, 5, 1781−1792. (26) Lee, R. J.; Low, P. S. Folate-mediated tumor cell targeting of liposome-entrapped doxorubicin in vitro. Biochim. Biophys. Acta 1995, 1233, 134−144. (27) Wu, J.; Liu, Q.; Lee, R. J. A folate receptor-targeted liposomal formulation for paclitaxel. Int. J. Pharm. 2006, 316, 148−153. (28) Tian, J.; Ding, L.; Xu, H. J.; Shen, Z.; Ju, H.; Jia, L.; Bao, L.; Yu, J. S. Cell-specific and pH-activatable rubyrin-loaded nanoparticles for highly selective near-infrared photodynamic therapy against cancer. J. Am. Chem. Soc. 2013, 135, 18850−18858. (29) Wang, H.; Zhao, Y.; Wu, Y.; Hu, Y. L.; Nan, K.; Nie, G.; Chen, H. Enhanced anti-tumor efficacy by co-delivery of doxorubicin and paclitaxel with amphiphilic methoxy PEG-PLGA copolymer nanoparticles. Biomaterials 2011, 32, 8281−8290. (30) Wei, T.; Liu, J.; Ma, H.; Cheng, Q.; Huang, Y.; Zhao, J.; Huo, S.; Xue, X.; Liang, Z.; Liang, X. J. Functionalized nanoscale micelles improve drug delivery for cancer therapy in vitro and in vivo. Nano Lett. 2013, 13, 2528−2534. (31) Liu, G.; Shen, H.; Mao, J.; Zhang, L.; Jiang, Z.; Sun, T.; Lan, Q.; Zhang, Z. Transferrin modified graphene oxide for glioma-targeted drug delivery: in vitro and in vivo evaluations. ACS Appl. Mater. Interfaces 2013, 5, 6909−6914. (32) Duan, X.; Xiao, J.; Yin, Q.; Zhang, Z.; Yu, H.; Mao, S.; Li, Y. Smart pH-sensitive and temporal-controlled polymeric micelles for effective combination therapy of doxorubicin and disulfiram. ACS Nano 2013, 7, 5858−5869. (33) Chen, J. X.; Xu, X. D.; Chen, W. H.; Zhang, X. Z. Multifunctional envelope-type nanoparticles assembled from amphiphilic peptidic prodrug with improved anti-tumor activity. ACS Appl. Mater. Interfaces 2014, 6, 593−598. (34) Graf, N.; Bielenberg, D. R.; Kolishetti, N.; Muus, C.; Banyard, J.; Farokhzad, O. C.; Lippard, S. J. αVβ3 integrin-targeted PLGA-PEG nanoparticles for enhanced anti-tumor efficacy of a Pt(IV) prodrug. ACS Nano 2012, 6, 4530−4539. (35) O’Neal, D. P.; Hirsch, L. R.; Halas, N. J.; Payne, J. D.; West, J. L. Photo-thermal tumor ablation in mice using near infrared-absorbing nanoparticles. Cancer Lett. 2004, 209, 171−176.

REFERENCES

(1) Jemal, A.; Bray, F.; Center, M. M.; Ferlay, J.; Ward, E.; Forman, D. Global cancer statistics. CA Cancer J. Clin. 2011, 61, 69−90. (2) Iyer, V. N.; Szybalski, W. Mitomycins and porfiromycin: Chemical mechanism of activation and cross-linking of DNA. Science 1964, 145, 55−58. (3) Tomasz, M.; Lipman, R.; Chowdary, D.; Pawlak, J.; Verdine, G. L.; Nakanishi, K. Isolation and structure of a covalent cross-link adduct between mitomycin C and DNA. Science 1987, 235, 1204−1208. (4) Bradner, W. T. Mitomycin C: a clinical update. Cancer Treat. Rev. 2001, 27, 35−50. (5) Ekins, S.; Kim, R. B.; Leake, B. F.; Dantzig, A. H.; Schuetz, E. G.; Lan, L.-B.; Yasuda, K.; Shepard, R. L.; Winter, M. A.; Schuetz, J. D.; Wikel, J. H.; Wrighton, S. A. Three-dimensional quantitative structureactivity relationships of inhibitors of P-glycoprotein. Mol. Pharmacol. 2002, 61, 964−973. (6) Ekins, S.; Kim, R. B.; Leake, B. F.; Dantzig, A. H.; Schuetz, E. G.; Lan, L.-B.; Yasuda, K.; Shepard, R. L.; Winter, M. A.; Schuetz, J. D.; Wikel, J. H.; Wrighton, S. A. Application of three-dimensional quantitative structure-activity relationships of P-glycoprotein inhibitors and substrates. Mol. Pharmacol. 2002, 61, 974−981. (7) Gontero, P.; Sargent, J. M.; Hopster, D. J.; Lewandowic, G. M.; Taylor, C. G.; Elgie, A. W.; Williamson, C. J.; Sriprasad, S. I.; Muir, G. H. Ex vivo chemosensitivity to mitomycin C in bladder cancer and its relationship with P-glycoprotein and apoptotic factors. Anticancer Res. 2002, 22, 4073−4080. (8) Gabizon, A. A.; Tzemach, D.; Horowitz, A. T.; Shmeeda, H.; Yeh, J.; Zalipsky, S. Reduced toxicity and superior therapeutic activity of a mitomycin C lipid-based prodrug incorporated in pegylated liposomes. Clin. Cancer Res. 2006, 12, 1913−1920. (9) Gabizon, A.; Amitay, Y.; Tzemach, D.; Gorin, J.; Shmeeda, H.; Zalipsky, S. Therapeutic efficacy of a lipid-based prodrug of mitomycin C in pegylated liposomes: studies with human gastro-entero-pancreatic ectopic tumor models. J. Controlled Release 2012, 160, 245−253. (10) Hou, Z.; Li, Y.; Huang, Y.; Zhou, C.; Lin, J.; Wang, Y.; Cui, F.; Zhou, S.; Jia, M.; Ye, S.; Zhang, Q. Phytosomes loaded with mitomycin C-soybean phosphatidylcholine complex developed for drug delivery. Mol. Pharmaceutics 2013, 10, 90−101. (11) Cheung, R. Y.; Ying, Y.; Rauth, A. M.; Marcon, N.; Yu, W. X. Biodegradable dextran-based microspheres for delivery of anticancer drug mitomycin C. Biomaterials 2005, 26, 5375−5385. (12) Hou, Z.; Wei, H.; Wang, Q.; Sun, Q.; Zhou, C.; Zhan, C.; Tang, X.; Zhang, Q. New method to prepare mitomycin C loaded PLAnanoparticles with high drug entrapment efficiency. Nanoscale Res. Lett. 2009, 4, 732−737. (13) El-Gogary, R. I.; Rubio, N.; Wang, J. T.; Al-Jamal, W. T.; Bourgognon, M.; Kafa, H.; Naeem, M.; Klippstein, R.; Abbate, V.; Leroux, F.; Bals, S.; Van Tendeloo, G.; Kamel, A. O.; Awad, G. A.; Mortada, N. D.; Al-Jamal, K. T. Polyethylene glycol conjugated polymeric nanocapsules for targeted delivery of quercetin to folateexpressing cancer cells in vitro and in vivo. ACS Nano 2014, 8, 1384− 1401. (14) Ling, G.; Zhang, P.; Zhang, W.; Sun, J.; Meng, X.; Qin, Y.; Deng, Y.; He, Z. Development of novel self-assembled DS-PLGA hybrid nanoparticles for improving oral bioavailability of vincristine sulfate by P-gp inhibition. J. Controlled Release 2010, 148, 241−248. (15) Shi, J.; Xiao, Z.; Votruba, A. R.; Vilos, C.; Farokhzad, O. C. Differentially charged hollow core/shell lipid-polymer-lipid hybrid nanoparticles for small interfering RNA delivery. Angew. Chem., Int. Ed. 2011, 50, 7027−7031. (16) Khan, J.; Alexander, A.; Ajazuddin; Saraf, S.; Saraf, S. Recent advances and future prospects of phyto-phospholipid complexation technique for improving pharmacokinetic profile of plant actives. J. Controlled Release 2013, 168, 50−60. (17) Kidd, P. M. Bioavailability and activity of phytosome complexes from botanical polyphenols: the silymarin, curcumin, green tea, and grape seed extracts. Altern. Med. Rev. 2009, 14, 226−246. L



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(36) Peer, D.; Karp, J. M.; Hong, S.; Farokhzad, O. C.; Margalit, R.; Langer, R. Nanocarriers as an emerging platform for cancer therapy. Nat. Nanotechnol. 2007, 2, 751−760. (37) Yezhelyev, M. V.; Qi, L.; O’Regan, R. M.; Nie, S.; Gao, X. Proton-sponge coated quantum dots for siRNA delivery and intracellular imaging. J. Am. Chem. Soc. 2008, 130, 9006−9012. (38) Rejman, J.; Oberle, V.; Zuhorn, I. S.; Hoekstra, D. Sizedependent internalization of particles via the pathways of clathrin- and caveolae-mediated endocytosis. Biochem. J. 2004, 377, 159−169. (39) Jaskolla, T.; Fuchs, B.; Karas, M.; Schiller, J. The new matrix 4chloro-alpha-cyanocinnamic acid allows the detection of phosphatidylethanolamine chloramines by MALDI-TOF mass spectrometry. J. Am. Soc. Mass Spectrom. 2009, 20, 867−874. (40) Xu, Q.; Liu, Y.; Su, S.; Li, W.; Chen, C.; Wu, Y. Anti-tumor activity of paclitaxel through dual-targeting carrier of cyclic RGD and transferrin conjugated hyperbranched copolymer nanoparticles. Biomaterials 2012, 33, 1627−1639. (41) Han, S.; Liu, Y.; Nie, X.; Xu, Q.; Jiao, F.; Li, W.; Zhao, Y.; Wu, Y.; Chen, C. Efficient delivery of antitumor drug to the nuclei of tumor cells by amphiphilic biodegradable poly(L-aspartic acid-co-lactic acid)/ DPPE co-polymer nanoparticles. Small 2012, 8, 1596−1606. (42) Fattal, E.; Couvreur, P.; Dubernet, C. “Smart” delivery of antisense oligonucleotides by anionic pH-sensitive liposomes. Adv. Drug Delivery Rev. 2004, 56, 931−946. (43) Connor, J.; Yatvin, M. B.; Huang, L. pH-sensitive liposomes: acid-induced liposome fusion. Proc. Natl. Acad. Sci. U.S.A. 1984, 81, 1715−1718. (44) Jin, E.; Zhang, B.; Sun, X.; Zhou, Z.; Ma, X.; Sun, Q.; Tang, J.; Shen, Y.; Van Kirk, E.; Murdoch, W. J.; Radosz, M. Acid-active cellpenetrating peptides for in vivo tumor-targeted drug delivery. J. Am. Chem. Soc. 2013, 135, 933−940. (45) Zhao, Y.; Sun, X.; Zhang, G.; Trewyn, B. G.; Slowing, I. I.; Lin, V. S. Interaction of mesoporous silica nanoparticles with human red blood cell membranes: size and surface effects. ACS Nano 2011, 5, 1366−1375. (46) Acharya, S.; Sahoo, S. K. PLGA nanoparticles containing various anticancer agents and tumour delivery by EPR effect. Adv. Drug Delivery Rev. 2011, 63, 170−183.

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Mitomycin C-Soybean Phosphatidylcholine Complex-Loaded Self

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