Dual-Layer Surface Coating of PLGA-Based Nanoparticles Provides

Sep 24, 2014 - Department of Cancer Immunology & AIDS, Dana-Farber Cancer Institute, Boston, Massachusetts 02215, United States. ‡. Department of ...
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Dual-Layer Surface Coating of PLGA-Based Nanoparticles Provides Slow-Release Drug Delivery To Achieve Metronomic Therapy in a Paclitaxel-Resistant Murine Ovarian Cancer Model Zohreh Amoozgar,†,‡ Lei Wang,†,‡ Tania Brandstoetter,†,‡ Samuel S. Wallis,†,‡ Erin M. Wilson,†,‡ and Michael S. Goldberg*,†,‡ †

Department of Cancer Immunology & AIDS, Dana-Farber Cancer Institute, Boston, Massachusetts 02215, United States Department of Microbiology & Immunobiology, Harvard Medical School, Boston, Massachusetts 02215, United States

‡

S Supporting Information *

ABSTRACT: Development of drug resistance is a central challenge to the treatment of ovarian cancer. Metronomic chemotherapy decreases the extent of drug-free periods, thereby hindering development of drug resistance. Intraperitoneal chemotherapy allows for treatment of tumors confined within the peritoneum, but achieving sustained tumor-localized chemotherapy remains difficult. We hypothesized that modulating the surface properties of poly(lactic-co-glycolic acid) (PLGA)based nanoparticles could enhance their drug retention ability and extend their release profile, thereby enabling metronomic, localized chemotherapy in vivo. Paclitaxel was encapsulated in particles coated with a layer of polydopamine and a subsequent layer of poly(ethylene glycol) (PEG). These particles achieved a 3.8-fold higher loading content compared to that of nanoparticles formulated from linear PLGA−PEG copolymers. In vitro release kinetic studies and in vivo drug distribution profiles demonstrate sustained release of paclitaxel. Although free drug conferred no survival advantage, low-dose intraperitoneal administration of paclitaxel-laden surface-coated nanoparticles to drugresistant ovarian tumor-bearing mice resulted in significant survival benefits in the absence of any apparent systemic toxicity.

1. INTRODUCTION Ovarian cancer has the highest mortality rate among gynecological cancers.1 More than 90% of ovarian cancers are of epithelial origin and represent the most lethal form of the disease.1,2 Typically, ovarian cancer does not manifest with specific symptoms until the cancer has progressed and disseminated throughout the peritoneal cavity.1 The current standard therapy for ovarian cancer includes surgical debulking of the tumors followed by intravenous (IV) administration of taxanes and platinum-based chemotherapeutics in consecutive cycles to eliminate residual cancer cells.3 While many patients achieve a complete response to chemotherapy, the disease eventually relapses due to the emergence of multidrug-resistant (MDR) tumors.4 Therefore, a therapy that prevents onset of relapse is urgently needed. Intraperitoneal (IP) chemotherapy allows for higher local drug concentration at the site of disease and theoretically reduces systemic toxicity.5 IP chemotherapy improves patient survival by 8−16 months relative to delivery of the same regimen by IV administration,6 and it is endorsed by the National Cancer Institute (NCI).7 Despite the observed survival benefits, the utility of IP chemotherapy remains limited due to heightened local toxicity in the abdominal region as well as unresolved systemic toxicity caused by clearance of small molecule chemotherapeutics (e.g., taxanes) to the systemic circulation.8 These © XXXX American Chemical Society

problems have hampered patient desire to complete treatment cycles and have consequently lowered the acceptance of IP chemotherapy by clinicians. Various drug delivery systems have been developed to improve therapeutic outcomes in ovarian cancer. Such drug delivery systems aim (i) to achieve greater local drug concentration, (ii) to lower systemic toxicity by enhancing the drug residence time in the peritoneal cavity, and (iii) to sustain drug release to maintain continuous presence of drug. To achieve these aims, implantable and injectable depots, microsized drug delivery systems, and nanosized drug delivery systems have been developed. Injectable depots such as PoLigel increase drug bioavailability by decreasing first-pass metabolism and sustaining drug release in preclinical models.9 The placement or injection of implants requires surgical expertise, and implantation of solid or semisolid implants can cause tissue damage and can invade surrounding tissue over time.10 Unlike nanosized particulate systems, implantable gels cannot penetrate into the tumor parenchyma. Similarly, microspheres can extend drug release profiles but have very limited tumor penetration capability and can cause inflammation.11 Received: August 13, 2014 Revised: September 22, 2014

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2.2. Preparation of Plain and PTX-Loaded NPs. NPs were prepared using a single-emulsion technique. Briefly, PLGA (200 mg) or PLGA−PEG was dissolved in 5 mL of DCM. The polymer solution was added to a 20 mL solution of 5% (w/v) PVA on ice and subsequently sonicated by Misonix (Farmingdale, NY) for 4 min (80% amplitude, 4 s on, 2 s off) to generate nanodroplets. For drug-loaded particles, PTX (20 mg) was dissolved in DCM along with the polymers. The NP droplets were solidified upon addition to 100 mL of DI water under magnetic stirring. The NP suspension was stirred overnight at room temperature to allow complete evaporation of organic solvents before flash freezing and lyophilization. NP size pre- and postdrying was measured by a Malvern Zetasizer Nano S series (Worcestershire, UK). 2.3. Surface Coating of NPs and Subsequent PEGylation. NPs (400 mg) were resuspended in 29 mL of Tris buffer (10 mM, pH 9). Dopamine (30 mg) was dissolved in 1 mL of water and added to the NP suspension under vigorous stirring at 4 °C for 3 h. Polydopamine-coated NPs (PD NP) were collected via centrifugation at 10 000 RCF for 1 h. Afterward, PD NP pellets were resuspended in 6 mL of Tris buffer (10 mM, pH 9), and NP size was measured (Malvern Zetasizer Nano S series, Worcestershire, UK). PEGylation was achieved via conjugation of PEG14 to the polydopamine surface by adding 15 or 30 mg of one of the following aminated-PEG subtypes: (i) methoxy-PEG-NH2 (5 kDa), (ii) 4 arm-PEG-NH2 (5 kDa), (iii) NH2-PEG-NH2 (7.5 kDa), (iv) 4 arm-PEG-NH2 (10 kDa), (v) 8 armPEG-NH2 (20 kDa), or (vi) Y-PEG-NH2 (40 kDa). The reaction continued under magnetic stirring for 2 h at 4 °C. PEGylated particles (PDP NP) were collected via centrifugation at 10 000 RCF for 1 h and washed with water via suspension and centrifugation as described above. Uncoated nanoparticles (PLGA NPs (P NP) or PLGA−PEG copolymer NPs (PP NPs)) or “bio-glued”-only nanoparticles (PD NPs) were collected and washed similarly to that for PDP-coated particles. NP size and surface charge were measured using a Malvern Zetasizer Nano S series (Worcestershire, UK). In summary, the following nanoparticles were formulated:

Nanosized drug delivery systems are of particular interest because they are the easiest to administer, generally do not cause substantial localized inflammation, and confer improved tumor penetration compared to that of implants and microspheres.11,12 Still, nanosized drug delivery systems have some limitations, including dose dumping and limited half-life due to vehicle clearance if they are smaller than 100 nm.13 These challenges must be overcome when designing effective and clinically useful nanosized delivery systems. We hypothesized that successful IP chemotherapy would combine elements of both the drug delivery system (sustained release and improved tumor penetration) and IP injection (localization to the disease site). We thus sought to design a nanosized delivery system with high drug loading, an extended drug release profile, and an appropriate size (100−200 nm) to achieve efficient localized therapy. It was anticipated that such nanoparticles (NPs) would be easy to administer, promote efficient tumor penetration, and maintain peritoneal localization. Preventing rapid clearance to the bloodstream is expected to reduce systemic toxicity. Such a system could theoretically allow one to reduce the administered dose and decrease the duration of drug-free rest periods for the tumor (“metronomic dosing”) while still generating a comparable therapeutic effect to that achieved with higher doses. Herein, we describe the use of a polymeric drug delivery system to engineer paclitaxel (PTX)-laden nanoparticles. We chose high molecular weight poly(lactic-co-glycolic acid) (PLGA) to achieve high initial drug loading12 and surface coated these NPs with polydopamine to prevent rapid diffusion of the encapsulated drug. A second layer of poly(ethylene glycol) (PEG) further prevents NP phagocytic uptake and rapid particle clearance. We manufactured particles with sizes of ∼150 nm to achieve tumor penetration with optimal IP residence time. We studied the release profile of such particles in vitro and the overall benefit of IP therapy in vivo utilizing an orthtopic syngeneic murine model of ovarian cancer.

P PD PP PDP

NP: PLGA NPs NP: Polydopamine-coated NPs NP: PLGA−PEG copolymer NPs NP: Polydopamine- and PEG-coated NPs

2.4. Confocal Microscopy and Flow Cytometry of NPMacrophage Interactions. Fluorescently labeled NPs were prepared by substituting 25% of the PLGA polymer with fluoresceinamineconjugated PLGA. NP size and zeta potential were measured prior to cell experiments using a Malvern Zetasizer Nano-ZS90 (Worcestershire, UK). J774A.1 mouse macrophages (ATCC) were grown in DMEM. All media contained 10% FBS, 100 units/mL penicillin, and 100 ug/mL streptomycin. J774A.1 cells were seeded at a density of 50 000 cells/cm2 in a 35 mm dish with a glass window (MatTek, Ashland, MA, USA). After overnight incubation, the media was replaced with a 0.1 mg/mL NP suspension in serum-free media, in which the cells were incubated for 1 h. Cells were then washed twice with 2 mL of serum-free media to remove free or loosely bound NPs prior to observation using a Leica confocal microscope (Wetzlar, Germany). DRAQ-5 nuclear stain (1−2 μL) was added 2−3 min prior to imaging. NPs and cell nuclei were excited using 488 and 633 nm lasers, respectively. The emission signals were read from 500 to 600 nm and 650 to 750 nm and expressed in green and blue, respectively. Fluorescently labeled NPs (0.5 mg/mL) were incubated with J774A.1 cells as described above. After 1 h of incubation, cells were scraped and washed in FACS buffer (PBS supplemented with 2% FBS). Cells were fixed and permeabilized with Cytofix/Cytoperm buffer. Next, a single-cell suspension was passed through a 70 μm mesh to remove possible cellular aggregates. A total of 25 000 events were recorded per sample in the FITC channel of a BD FACSAria. Data were analyzed using FlowJo software. 2.5. Evaluation of Surface Adsorption of Proteins to NPs. Lyophilized FBS was fluorescently labeled with FITC. Then, NPs (10 mg/mL) were incubated with DMEM media containing 50%

2. MATERIALS AND METHODS 2.1. Materials. Dopamine hydrochloride, dichloromethane (DCM), Tween 80, Tris buffer, methanol, and acetonitrile (ACN) were purchased from Sigma-Aldrich (St. Louis, MO, USA). 3-(4,5Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT), Dulbecco’s modified Eagle medium (DMEM), opti-MEM, Gibco fetal bovine serum (FBS), and penicillin−streptomycin (10 000 U/mL) were purchased from Life Technologies (Grand Island, NY, USA). Poly(lactic-co-glycolic acid) (lactic acid/glycolic acid = 50:50, carboxylic acid terminus) with molecular weights of 3 and 125 kDa were purchased from Akina (West Lafayette, IN, USA) and Durect Corporation (Birmingham, AL, USA), respectively. Linear copolymer of PLGA−PEG (55 kDa PLGA, 5 kDa PEG, 8% conjugated PEG) was purchased from Akina (West Lafayette, IN, USA). Poly(vinyl alcohol) (PVA, MW 6000) was obtained from Polysciences Inc. (Warrington, PA, USA). Poly(ethylene glycol)s (single and multibranched, MW = 5000, 10 000, 20 000, 40 000, and 80 000, with amine end functionality) were purchased from JenKem Technology USA (Allen, TX, USA). Paclitaxel (PTX) was obtained from Selleckchem (Houston, TX, USA). Tube-a-Lyzers was purchased from Spectrum Laboratories Inc. (Rancho Dominguez, CA, USA). Cytofix/Cytoperm was purchased from BD Biosciences (San Diego, CA, USA). BR5FVB1-Akt (murine ovarian cancer cell line) was kindly provided by Dr. Sandra Orsulic (Samuel Oschin Comprehensive Cancer Institute, CedarsSinai Medical Center, Los Angeles, CA, USA). SKOV3 (human ovarian cancer cell line) and Calu6 (human lung cancer cell line) were acquired from ATCC (Manassas, VA, USA). B

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Figure 1. Nanoparticle formulation and properties. (A) PLGA-based nanoparticles are formed by single-emulsion method and then coated with polydopamine and subsequently with poly(ethylene glycol) (4 arm-PEG-NH2, named PDP). (B) Plain and surface-coated particles were of sizes below 200 nm and remained stable in suspension for 1 week at 4 °C. (C) Uncoated PLGA particles are spherical with a smooth surface. Deposition of polydopamine on the particle surface does not affect shape or surface smoothness for nanoparticles made of high molecular weight PLGA (125 kDa). FITC-FBS for 1 h at 37 °C. NPs were transferred to centrifuge tubes (MWCO 100 kDa) and washed twice with excess PBS. The NP suspension was transferred to Tube-a-Lyzers (MWCO 100 kDa) and dialyzed in water twice for 4 h intervals and a third time overnight. NPs were then flash frozen and lyophilized. NPs were resuspended in water at a concentration of 1 mg/mL, and fluorescence was measured (485 nm excitation, 535 nm emission) at room temperature using a Spectra Max M5 plate reader (Molecular Devices, Sunnyvale, CA, USA) 2.6. Morphology of NPs. Lyophilized NPs were attached to the sample mount using a double-sided conductive carbon adhesive tab and coated with 5 nm platinum using a Cressington HR 208 sputter coater (Cressington Scientific Instruments Ltd., Watford, UK). Samples were imaged with a Hitachi S-4800 field emission scanning electron microscope (Hitachi, Japan) using the high-resolution throughthe-lens detector (TLD) operating at 3 kV accelerating voltage and ∼2.9−3.0 mm working distance. 2.7. Evaluation of Drug Loading and Encapsulation Efficiency of PTX in PP and PDP NPs. Lyophilized NPs were accurately weighed, dissolved in a 50:50 mixture of ACN and water, and analyzed using reversed-phase high-pressure liquid chromatography (HPLC) equipped with an Ascentis C18 column (St. Louis, MO, USA). The mobile phase for PTX-laden NPs is a 50:50 mixture of ACN and water, with a flow rate of 1 mL/min. PTX peaks were detected using a multiwavelength UV detector at 227 nm. PTX content was calculated as a weight percentage of the drugs in the NPs. 2.8. PTX-Laden NP Release Kinetics. NPs of a known concentration (1 mg/mL) were resuspended in 1 mL of release buffer, consisting of PBS (10 mM phosphate, pH 7.4) containing 0.1% Tween 80, and incubated in a rotating shaker at 37 °C. At regular time points (1, 3, 7, 12, 24, 48, and 72 h), the NP suspension was centrifuged at 10 000 rpm for 10 min, and 0.9 mL of the supernatant was collected and replaced with fresh buffer. The remaining NPs (after 72 h) were lyophilized and dissolved in a 50:50 mixture of ACN and water. The

collected supernatant release samples and remaining NPs were analyzed by HPLC. 2.9. Cell Viability Assay. Cell cytotoxicity of PTX-laden NPs, empty NPs, and free PTX was evaluated using BR5FVB1-Akt (murine ovarian cancer), SKOV3 (human ovarian cancer), and Calu6 (human lung cancer) cell lines. Cells were seeded into a 96-well plate at a density of 10 000 cells per well and incubated overnight in 200 μL of complete media. The media was replaced with 200 μL of fresh media and dosed with particles to concentrations of 100, 10, 1, and 0.1 μg/ mL for either 3 or 72 h. After the appropriate incubation time, the media was replaced with 100 μL of fresh media containing 13% MTT and incubated for 3.5 h. Finally, 100 μL of solubilization/stop solution (composed of 20% SDS, 0.02% (v/v) acetic acid, and 50% (v/v) dimethyl sulfoxide (DMSO)) was added to each well, and the absorbance was read at 560 nm using a PerkinElmer 1420 VICTOR 3V multilabel counter (PerkinElmer, Waltham, MA, USA). Cell viability was calculated by dividing the absorbance of the treated cells by that of the untreated cells, after subtracting the background absorbance of cell-free MTT media from each. 2.10. Determination of PTX Levels in Murine Blood and Peritoneal Cavity. Six week-old female FVB mice were obtained from Charles River Laboratory (Boston, MA, USA). Mice were housed in a pathogen-free facility in accordance with the standards of the Dana-Farber Cancer Institute. All animal experiments were approved by the Institute’s IACUC. Mice were kept in groups of 4−5 per cage in a 12 h light/12 h dark cycle and housed at 25 °C with 50% relative humidity. Mice (n = 4 per group) were given IP doses of 5 mg/kg free PTX or PTX−PDP NP as a rapid bolus. Mice were anaesthetized, and blood samples were collected via retro-orbital bleeding 3 h post dosing. Blood samples were kept on ice until centrifugation. Samples were centrifuged at 1400 RCF for 10 min, and plasma was collected. After blood sample collection, mice were sacrificed, and their peritoneal cavities were washed thrice post-mortem with PBS. Plasma and peritoneal washes were flash frozen with liquid nitrogen and C

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lyophilized. The lyophilized powders were rehydrated in 300 μL of water and extracted with tert-butyl methyl ether (tBME). The organic solvent was dried under nitrogen. The samples were resuspended in a 50:50 mixture of ACN and water and analyzed by HPLC as described in Section 2.7. 2.11. IP Tumor Development and Subsequent Survival Study in an Immunocompetent Murine Model of Ovarian Cancer. The BR5FVB1-Akt mouse ovarian cancer cell line was cultured in DMEM supplemented with 10% FBS and 1% penicillin− streptomycin at 37 °C with 5% CO2. Subconfluent BR5FVB1-Akt cells were trypsinized, washed, and suspended in opti-MEM as a single-cell suspension. Six week-old female FVB mice were injected IP with a total volume of 400 μL opti-MEM containing 3.75 × 106 BR5FVB1Akt cells. Four days after IP inoculation of cancer cells, the mice (n = 8) were treated with PTX (5 mg/kg dissolved in 50 μL DMSO and 450 μL PBS), PTX−PDP NP (5 mg/kg in PBS), or controls (DMSO/PBS mix or PDP NP in PBS) twice weekly for a total of five doses. The mice, which received no further treatment, were subsequently monitored for health. Mice were sacrificed by CO2 asphyxiation upon ascites development. Their peritoneal cavities were opened postmortem to observe tumor location and extent of metastasis in the peritoneal region. 2.12. Statistical Analysis. Statistical analysis was performed with GraphPad Prism 5 (La Jolla, CA, USA). Statistical significance among groups was determined by one-way ANOVA, and pairs were compared by Tukey’s multicomparison test. Survival groups were compared using a Mantel−Cox test. Samples were considered to be significantly different if p < 0.05.

3. RESULTS 3.1. Production and Characterization of NPs. NPs were formulated with sizes below 200 nm and a narrow distribution range (Figure 1A,B). These results were observed regardless of the PLGA molecular weight used. Surface coating with polydopamine decreased surface charge from −15.00 ± 9.21 to −25.00 ± 7.31 mV. Subsequent PEGylation further reduced surface charge toward neutrality. Depositing polydopamine followed by PEGylation did not change the surface morphology of NPs made from high molecular weight PLGA (MW = 125 kDa) (Figure 1C). However, the same coating procedure resulted in surface roughness for NPs made of low molecular weight PLGA (MW = 3 kDa), which was not observed for uncoated NPs (Supporting Information Figure 1). The coating did not affect NP size, which remained stable for 1 week at 4 °C. 3.2. Selection of PEG Chemistry for Coating Based on Macrophage Evading Capacity of NPs and Extent of Protein Adsorption. NPs were PEGylated by linkage of the PEG’s amine terminus to the polydopamine-coated particles. Two concentrations of PEG (1 and 3 mg/mL) were used for coating. One mg/mL PEGylation did not result in macrophage (J774A.1) evasion, but 3 mg/mL of (i) 4 arm-PEG-NH2 (5 kDa), (ii) NH2−PEG-NH2 (7.5 kDa), or (iii) 4 arm-PEG-NH2 (10 kDa) prevented uptake by macrophages. Among these, 4 arm-PEGNH2 (5 kDa) most strongly prevented phagocytosis. Flow cytometry analysis additionally showed greater uptake of NPs that were coated with PEGs of larger molecular weights (10 and 40 kDa) (Figure 2B and Supporting Information Figure 2). Next, it was shown that effective coatings could also minimize protein adsorption (Figure 2C), as summarized here. Micrograms of protein adsorbed to 1 mg of each NP are as follows: PP NP = 4.8 ± 0.4, methoxy-PEG-NH2 (5 kDa) = 12.6 ± 0.9, 4 arm-PEG-NH2 (5 kDa) = 3.6 ± 0.5, NH2-PEG-NH2 (7.5 kDa) = 6.0 ± 0.4, 4 arm-PEG-NH2 (10 kDa) = 6.2 ± 0.6, 8 arm-PEGNH2 (20 kDa) = 14.6 ± 0.5, and Y-PEG-NH2 (40 kDa) = 16.4 ± 1.1.

Figure 2. Presence and type of PEGylation alter opsonization of nanoparticles. (A) PEGylated particles are able to evade phagocytosis by macrophages. J774A.1 macrophages were incubated with uncoated or PEGylated nanoparticles for 1 h. PEGylated particles (PDP NP) were not taken up by macrophages. Overlaid images of NP (green), nuclei (blue), and transmission images. Flow cytometry reveals that only certain types of PEGylation (B) decrease NP uptake and (C) lower the amount of protein adsorbed to 1 mg of NPs. Samples can be divided in two groups. Group 1 (effective): PP NP, 4 arm-PEG-NH2 (5 kDa), NH2-PEG-NH2 (7.5 kDa), and 4 arm-PEG-NH2 (10 kDa). Group 2 (not effective): methoxy-PEG-NH2 (5 kDa), 8 arm-PEGNH2 (20 kDa), and Y-PEG-NH2 (40 kDa). Groups 1 and 2 are significantly different from each other, with p value < 0.001. There is no significant difference within group 1 or 2. Data are expressed as averages with standard deviations of five identically and independently prepared samples.

3.3. Efficiency of PTX Encapsulation in PP and PDP NPs and Kinetics of PTX Release. PTX was encapsulated in NPs by addition to organic solvent (DCM) during emulsification (Figure 3A). All NPs were dried, coated, and/ or washed similarly. The yield of NPs produced by coating particles was ∼1.5 times higher than that of NPs produced from linear PLGA−PEG copolymers. Most of the encapsulated PTX was released from PP (copolymer) NPs during washes, resulting in lower loading and encapsulation efficiencies compared to those of coated NPs. Indeed, dual-coated NPs contained 3.8fold more PTX than PP NPs. The release kinetics of PTX from NPs was examined in PBS doped with Tween 80 to mimic physiological conditions, as described previously15 (Figure 3B). Within the first 3 h, PP NP released 60.8 ± 2.7% of its PTX payload, whereas coated NPs (PD NP and PDP NP) released 40.8 ± 2.7 and 31.1 ± 1.8% PTX, respectively. The coated NPs released the remaining PTX slowly over 72 h. D

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Figure 3. Nanoparticles successfully encapsulate PTX and released this payload over an extended period of time. (A) Loading of PTX was measured after particle formation, purification, and freeze-drying. Drug was extracted using organic solvents, and its quantity was measured by HPLC according to the AUC of its absorbance peak (227 nm) compared to a standard curve. The loading efficiency (LE) is determined as the percentage of drug in the total mass of the NP. The encapsulation efficiency (EE) is determined by the ratio of drug present within particles compared to the amount of drug added to the formulation emulsion prior to particle formation and purification. (B) Release of drugs in serum-simulated media at sink condition was measured at predetermined intervals. Only 31.1% of PTX was released from PTX−PDP NPs in the first 3 h, with the rest steadily released over 72 h. With the exception of total release (“remaining samples”), PTX−PP NP, PTX−PD NP, and PTX−PDP NP samples were significantly different from each other (p < 0.05) for each of the given time points.

Figure 4. Cell viability of BR5FVB1-Akt cells exposed to PTX and PTX-laden nanoparticles for 3 or 72 h in a dose-dependent manner. PTX present in the media is quickly taken up by cells and kills cells irrespective of additional exposure time. PTX−PP NPs release the drug quickly and thus exhibit similar behavior to that of free drug. PTX−PD NPs, which are coated with a cellular adhesive material, maintain contact with cells after media substitution (after 3 h) and continue the drug release. In accordance with their in vitro release profiles, PTX−PDP NPs exhibit limited cytotoxicity after the short exposure, as only 30% of their PTX payload is released in the first 3 h.

3.4. Time-Dependent Cell Cytotoxicity in Vitro. The cytotoxic effects of free PTX, PTX−PP NPs, PTX−PD NPs, and PTX−PDP NPs on BR5FVB1-Akt (mouse ovarian cancer), SKOV3 (human ovarian cancer), and Calu6 (human lung cancer) cells were evaluated in a time-dependent manner. When cells were exposed to free PTX or drug-laden NPs for 72 h, the dose response was similar. The IC50 of PTX in SKOV3 and Calu6 cells was the same (∼10 nM); however, BR5FVB1-Akt cells are 10-fold more resistant to PTX, with an IC50 of ∼100 nM. However, when cells were exposed to free PTX or NPs for a short period of time (3 h), a significant difference in cytotoxicity pattern emerged. Unlike for free PTX and PD NPs, toxicity was substantially reduced for PDP NPs. A decrease in toxicity was similarly observed following acute exposure to PP NPs, but to a much lesser extent (Figure 4 and Supporting Information Figure 3). 3.5. PTX Content in Murine Blood and Peritoneal Cavity. To quantify PTX concentration in murine blood and peritoneal cavity, samples were collected via retro-orbital bleeding (pre-mortem) and peritoneal lavage (post-mortem) (Figure 5). Mice that received free PTX had three times as much PTX in their systemic circulation (9.5 ± 1.9 μg/mL) as that of mice treated with PTX−PDP NPs (3.0 ± 1.8 μg/mL). In contrast, an ∼8-fold larger dose of PTX (32.8 ± 9.5 μg/mL) was detected in the

Figure 5. PTX concentration in blood and peritoneal lavage fluid 3 h after free PTX or PTX−PDP NP treatment. PTX (5 mg/kg) in solution or as PTX−PDP NP was injected IP to six week-old female FVB mice (n = 4 per group). Blood and peritoneal lavage fluid were collected 3 h post treatment. These data show that PTX is retained in the peritoneal cavity for at least 3 h when administered in nanoparticles compared to that of free drug. ***: p < 0.001, for PTX measurements in peritoneum; **: p < 0.02 for blood samples (Tukey’s test). PTX−PDP NP (peritoneum, ■); PTX−PDP NP (blood/mL, □); PTX (peritoneum, ●); PTX (blood/mL, ○).

peritoneal cavities of mice when PTX−PDP NPs were administered compared to that when free PTX was administered (4.5 ± 1.3 μg/mL). E

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3.6. IP Tumor Development and Mouse Survival Following Treatment. All FVB mice inoculated with BR5FVB1-Akt cells developed tumors. Tumors developed gradually and manifested by formation of ascites in the peritoneal cavity. Mice were sacrificed upon ascites development. Tumors were detectable post-mortem as nodules of various sizes and number on the peritoneal wall or adhered to the intestine. Control groups (drug-free NPs and DMSO:PBS) and mice that received free PTX (5 mg/kg) developed tumors at similar time points during the fourth week after tumor cell injection. PTX−PDP NP significantly improved survival, as tumor onset in mice treated with these particles was delayed by 1 to 2 weeks (Figure 6).

drug release. This can ultimately result in poor drug penetration to solid tumors, which is a potential mechanism of chemotherapy failure.23 While microspheres prolong drug release, they similarly confer limited tumor penetration.24 NPs accumulate in the tumor, but many NP systems release encapsulated drug too rapidly. The resultant necessity for frequent dosing limits their practical utility. We hypothesized that metronomic dosing for treatment of ovarian cancer can be achieved using a suitable drug delivery system that sustains drug release and penetrates deep into the tumor to impart significant therapeutic benefits. We thus engineered NPs with high drug loading and a significantly extended release profile compared to that of established PLGAbased particles. PTX-laden PLGA-based NPs were formulated by singleemulsion technique25 and freeze-dried prior to removal of the surfactant (PVA), polydopamine surface coating, and PEGylation. The method described herein allows for large-scale production of NPs with tight control over particle size. Importantly, the presence of PVA during polydopamine coating prevents particle aggregation, which is a possible explanation for the unchanged particle size observed following polydopamine coating. Upon removal of unreacted dopamine and PVA, NPs were coated with PEG via chelation of PEG-amine termini to the polydopaminated surface of the particles. By combining surface coating and purification of PVA into a single step, we reduced the number of washes needed, enabling robust control of particle size. During surface coating, PVA hydroxyl end functionality did not compete with PEG-amine due to limited affinity to polydopamine chelation as well as the low concentration of PVA (trace) compared to that of PEG-amine (3 mg/mL). Additionally, the temperature was maintained at 4 °C during coating stages to increase the activation energy and therefore shift the equilibrium toward the formation of the dominant product (PEGylation) and thereby eliminate side products and aggregates. Consequently, by reducing the speed of reaction, we were able to coat the particles uniformly with a thin layer of PEG that did not increase nanoparticle size or surface roughness. The longevity of NPs in blood or the peritoneal cavity requires sufficient PEGylation to prevent phagocytic uptake by macrophages. In our setting, multiarm aminated PEGs of sizes between 5 and 10 kDa minimized macrophage uptake (Supporting Information Figure 2). In contrast, monoamine PEG (5 kDa) resulted in only a partial stealth coat. For functional studies, we selected 4 arm-PEG-NH2, owing to its low uptake by macrophages and low protein adsorption compared to that of other PEGs (Figure 2B,C). Previous studies26 showed that 10 wt % PEG density is optimal for polylactic acid (PLA) and PLGAbased NPs. However, these studies focused on NPs made of linear copolymers of PLGA−PEG that, upon particle formation, form an exterior layer of PEG around the NPs. For our surface-coated NPs, 2 wt % PEG was sufficient to result in macrophage evasion. This may be due to a tight PEG loop forming on the NP surface by adherence of the 4 aminated termini per PEG molecule to the NP surface, which could inhibit surface interactions with opsonizing proteins that mark foreign bodies for rapid elimination via phagocytosis. An efficient delivery system should maximize the drug to polymer ratio. Here, PTX−PDP NPs contained 3.8-fold more PTX than NPs formulated from linear PLGA−PEG copolymer (PTX−PP NP). Since the number of purification steps is the same for both PDP and PP NPs, enhanced drug loading can be attributed to the dual surface coating. Our data suggest that the

Figure 6. Dual-coated nanoparticles enhance survival of ovarian tumor-bearing mice. PTX-resistant BR5FVB1-Akt cells were injected IP into six week-old female FVB mice. Four days post IP inoculation of cancer cells, mice (n = 8) were treated with PTX (□, 5 mg/kg dissolved in 50 μL DMSO and 450 μL PBS), PTX−PDP NP (●, 5 mg/kg in PBS), or controls (•, DMSO/PBS mix or PDP NP in PBS) twice weekly for a total of 5 doses. Metronomic dosing of PTX, achieved by sustained released from PTX−PDP NP, significantly enhances survival of ovarian tumor-bearing mice. In contrast, free PTX at the same dose (5 mg/kg) or control (PBS:DMSO vehicle) had no effect on mouse survival. ***: p < 0.001 □.

4. DISCUSSION Since ovarian cancer is mostly confined to the peritoneal cavity,16 IP chemotherapy is utilized to expose tumors to high drug concentrations.8b,17 Despite the pharmacokinetic advantage of this approach,6 small molecule drugs are eventually absorbed into the systemic circulation.18 Two approaches, metronomic dosing and localized drug delivery platforms, are used to enhance exposure of tumors to chemotherapeutics. In metronomic dosing, chemotherapeutics are administered at doses significantly below the maximum tolerated dose, allowing for minimized drug-free periods.19 It has been observed that continuous presence of chemotherapeutics increases expression of the endogenous angiogenesis inhibitor thrombospodin-1,20 although other antitumor mechanisms may also contribute. Unfortunately, the need for such frequent administration of chemotherapeutics limits the clinical utility of metronomic dosing. Localized drug delivery systems have also been developed to improve ovarian tumor responsiveness to chemotherapy.9,21 Many of these strategies aim to deliver drugs via passive diffusion from the vehicle to the peritoneal cavity (e.g., implantable gels and devices9,21) or directly to the tumor (e.g., particulate systems22). Among these delivery vehicles, gels may not control and sustain the release of hydrophobic drugs due to the hydrophilic nature of hydrogels.21 Specifically, PTX forms drug aggregates in hyaluronic-based hydrogels21 that hamper sustained F

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Aside from the therapeutic benefits, no weight loss or other visible side effects were observed in mice treated with PTX− PDP NP.

polydopamine layer, along with the 4 arm-PEG-NH2 layer, on the surface of the particles forms a surface barrier that reduces drug loss during purification steps and extends drug retention in vitro and in vivo. The high loading efficiency achieved with this system reduces the quantity of particle dose required, as the amount of excipient in the formulation is reduced. Importantly, we observed that our dual-layer coating with polydopamine and PEG not only improves loading efficiency but also extends the release profile of PTX compared to that of conventional PP NPs. The surface coating of PTX−PDP NPs reduces the cumulative release of PTX by 1.9-fold in the first 7 h (41% PTX released in PTX−PDP NP) and sustains the release for 3 days. In contrast, 75% of PTX is released from traditional PP NPs in the first 7 h, and the remainder is released within 1 day (Figure 3B). Continuous low-dose administration of anticancer drugs (metronomic dosing) limits development of MDR tumors and enhances survival.27 Because PTX release is slow from PTX− PDP NPs, this system could be used for metronomic dosing in situ. Since the release kinetic studies were performed in release buffer rather than full blood containing cellular components, we further evaluated drug release in the presence of BR5FVB1-Akt cells. PTX release from NPs results in cellular cytotoxicity in a dose-dependent manner. PTX release is time-dependent, so BR5FVB1-Akt cells exposed to free PTX or PTX-laden NPs for 72 h (when total PTX release is achieved) exhibit the same dose−response curve. In contrast, cells exposed to PTX−PDP NP for a short amount of time (3 h) had lower toxicity compared to that of free PTX, PTX−PD NPs, or PTX−PP NPs (Figure 4). This result correlates with the observed release profile, with the exception of PTX−PD NPs. The shortexposure (3 h) cytotoxicity of PTX−PD NP could be attributed to polydopamine-related cellular adhesion (Figure 2A) that would enable the NPs to remain with the cells and continue PTX release even following media exchange. Importantly, it has been shown that polydopamine detachment occurs as pH values below 5,28 a value that is below the pH of many tumors.15 Therefore, it is unlikely that the release profile of PDP NPs is greatly affected by tumor pH. We validated the in vitro release profile and in vivo PTX distribution by comparing drug retention in the peritoneal cavity following IP administration of PTX−PDP NPs or free PTX (Figure 5). At 3 h postinjection, PTX blood concentration in mice that received free PTX (9.5 ± 1.9 μg) was 3-fold higher than PTX blood concentration in mice that received PTX− PDP NPs (3 ± 1.9 μg). At the same time point, PTX in the peritoneal cavity was 4.5 ± 1.3 and 32.8 ± 9.5 μg for free PTX and PTX−PDP NP, respectively. Together, these data demonstrate that PTX−PDP NP can retain drug locally in the peritoneal cavity. Such increased drug residence time allows for metronomic dosing and translates to significant survival benefit in orthotopic syngeneic tumor-bearing mice (Figure 6). We inoculated BR5FVB1-Akt cells into FVB mice and treated them with 5 mg/kg PTX, either as free PTX or PTX−PDP NP. Typically, 5 or more doses of 20 mg/kg of PTX are used in murine models of ovarian cancer.29 Sustained release of low-dose chemotherapy using PTX−PD NPs extended survival significantly, whereas free drug did not. Importantly, we demonstrated that metronomic dosing was associated with significant survival benefits in BR5FVB1-Akt tumor cells (IC50 ∼ 100 nM), which are an order of magnitude more resistant to PTX therapy than SKOV3 ovarian cancer cells (PTX IC50 ∼ 10 nM) (Figure 4).

5. CONCLUSIONS Metronomic dosing of PTX in ovarian tumor-bearing mice was achieved using a dual-layer coated NP drug delivery system. Surface coating of particles enabled high drug loading and lowered the rate of drug release from NPs in vitro compared to that of traditional PLGA-based NPs formulated from linear PLGA−PEG copolymer. Interestingly, dual-coated particles with only 2% PEGylation evaded macrophage phagocytosis, despite previous reports suggesting that at least 5% PEGylation was required to prevent uptake (with 10% being optimal) using the copolymer.26 The in vitro PTX release profile observed translated to meaningful in vivo results, as particles administered IP were retained in the peritoneal cavity and released PTX slowly. As a consequence, IP administration of PTX−PDP NPs to tumor-bearing mice conferred significant survival benefit with no visible side effects. The favorable properties of the dualcoated NPs greatly enhanced the therapeutic index of this commonly used chemotherapeutic and led to substantial efficacy using only 25% of the recommended PTX dose for murine models of ovarian cancer. Together, these data highlight the impact of (i) localizing the therapy and (ii) prolonging drug release to achieve therapeutic benefits with limited off-target side effects. Future studies of tumor components, especially the tumor microenvironment, after therapy would enable mechanistic understanding of the therapeutic processes.

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ASSOCIATED CONTENT

* Supporting Information S

Particle morphology, uptake, and release kinetics data. This material is available free of charge via the Internet at http:// pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Author Contributions

The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors would like to thank Aid for Cancer Research (ACR) for its financial support of Dr. Amoozgar (ACR Fellow) and the Ovarian Cancer Research Fund (OCRF) for its support of both this work and Dr. Goldberg (Liz Tilberis Scholar).

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ABBREVIATIONS NP, nanoparticle; PD, polydopamine; PEG, poly(ethylene glycol); PLGA, poly(lactic-co-glycolic acid); PTX, paclitaxel; P NP, PLGA nanoparticles; PD NP, polydopamine-coated nanoparticles; PP NP, PLGA−PEG copolymer nanoparticles; PDP NP, polydopamine- and PEG-coated nanoparticles G

dx.doi.org/10.1021/bm5011933 | Biomacromolecules XXXX, XXX, XXX−XXX

Biomacromolecules

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Article

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dx.doi.org/10.1021/bm5011933 | Biomacromolecules XXXX, XXX, XXX−XXX