Article pubs.acs.org/Biomac
Synthesis and Characterization of Hybrid Polymer/Lipid Expansile Nanoparticles: Imparting Surface Functionality for Targeting and Stability Michelle Stolzoff,† Iriny Ekladious,† Aaron H. Colby,† Yolonda L. Colson,§ Tyrone M. Porter,*,‡ and Mark W. Grinstaff*,† †
Departments of Biomedical Engineering and Chemistry and ‡Departments of Mechanical Engineering and Biomedical Engineering, Boston University, Boston, Massachusetts 02215, United States § Division of Thoracic Surgery, Department of Surgery, Brigham and Women’s Hospital, Boston, Massachusetts 02115, United States S Supporting Information *
ABSTRACT: The size, drug loading, drug release kinetics, localization, biodistribution, and stability of a given polymeric nanoparticle (NP) system depend on the composition of the NP core as well as its surface properties. In this study, novel, pH-responsive, and lipid-coated NPs, which expand in size from a diameter of approximately 100 to 1000 nm in the presence of a mildly acidic pH environment, are synthesized and characterized. Specifically, a combined miniemulsion and free-radical polymerization method is used to prepare the NPs in the presence of PEGylated lipids. These PEGylated-lipid expansile NPs (PEG-L-eNPs) combine the swelling behavior of the polymeric core of expansile NPs with the improved colloidal stability and surface functionality of PEGylated liposomes. The surface functionality of PEG-L-eNPs allows for the incorporation of folic acid (FA) and folate receptor-targeting. The resulting hybrid polymer/lipid nanocarriers, FA-PEG-L-eNPs, exhibit greater in vitro uptake and potency when loaded with paclitaxel compared to nontargeted PEG-L-eNPs.
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INTRODUCTION Nanoparticle (NP) drug delivery systems possess several advantages over conventional small molecule therapies.1−7 Among these is the potential to provide controlled release as well as sustained levels of drug within cells, thereby conferring greater efficacy and reduced toxicity. As a result of their size, surface properties, and core properties, NPs can be used to dramatically alter the pharmacokinetics of entrapped drugs. For example, entrapping doxorubicin within PEGylated liposomes increases the circulation time and cumulative maximum tolerated dose of the drug.8 Polymeric NPs are especially attractive in cancer treatment because of their ability to encapsulate water insoluble agents, provide controlled drug release rates, and target tumors.2,9 Additionally, NPs can increase the maximum tolerated dose of encapsulated drugs by circumventing the solubility limitations of certain anticancer agents.1 The composition of the NP polymeric core dictates important features such as drug loading, NP size, and drug release kinetics, while surface properties affect NP localization, biodistribution, and stability. The selection of the polymer for the core is critical. Currently, poly(lactic acid) (PLA) and poly(lactic-co-glycolic) (PLGA) polymers are the most widely studied polymers in nanocarrier systems due to their biocompatibility, biodegradability, and availability.9−14 As such, PLGA and PLA are used © XXXX American Chemical Society
for the encapsulation and delivery of a wide range of chemotherapeutic agents including paclitaxel, doxorubicin, and docetaxel.15−19 However, PLGA and PLA core systems are criticized for their relatively rapid “burst” release of encapsulated drug.20,21 It was previously demonstrated that half of the encapsulated agent in a PLGA−PEG NP system can be released within minuteswell before any significant tumor accumulation would occur in vivo.22,23 Consequently, burst release from circulating particles may compromise the effectiveness of the anticancer agent and increase the risk of systemic toxicities. To address this limitation, scientists are engineering stimuli-responsive polymeric NPs to achieve triggered drug release within solid tumors. Potential triggers for release include enzymatic activity, acidity, light activation, temperature, and other chemical or physical stimuli.24−31 For example, β-amino ester linkages within a polymer confer pH sensitivity, as the protonation of tertiary amines in these polymersnear pH 6.5causes the polymer core to become soluble in water where it was previously insoluble at pH 7.4.31,32 The relatively acidic pH of tumor extracellular fluid is well-documented and is now considered to be a phenotype of Received: March 12, 2015 Revised: May 30, 2015
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Figure 1. Schematic of eNP surface coatings and pH-triggered swelling.
PEGylating the surface of polymeric NPs. For example, NPs can be formulated from diblock copolymers composed of PEG and a polyester. However, the copolymer must be synthesized, and it is difficult to control the molecular weight of each block to minimize polydispersity and batch variability. As an alternative approach, Park et al. coated polymeric NPs with lipids functionalized with avidin and then conjugated biotinylated PEG to the shell.44 Unfortunately, avidin is immunogenic, which prompted the exploration of alternative linkers for PEG. PEG−lipid conjugates with more suitable chemistries have been synthesized and incorporated in the formulation of hybrid lipid−polymer assemblies for various biotechnological and biomedical applications.9,45−47 Additionally, PEG terminated with a targeting moiety can improve the specificity with which lipid−polymer hybrid assemblies deliver drugs.17 Thus, PEGylated lipid coatings on polymeric NPs serve multiple purposes with respect to drug delivery, and we sought to explore this further with our polymer-based NP platform. We have previously engineered a novel pH-responsive expansile NP (eNP) that expands from 100 to 1000 nm under mildly acidic conditions, such as those found in the endosome and lysosome, resulting in intracellular release of the encapsulated payload.48 The eNPs are synthesized from a monomer that has a 2,4,6-trimethoxybenzylidene acetal protecting group that is stable at neutral pH but hydrolyzes under mildly acidic conditions (pH ∼5). Hydrolysis reveals the once protected hydroxyl groups, which increases the hydrophilicity of the polymeric core. As a result, eNPs swell in aqueous solutions and rapidly release the entrapped payload (Figure 1). Previously, paclitaxel-loaded eNPs demonstrated
solid tumors, making pH-sensitivity a viable route of NP targeting to tumors.33 NPs can also be designed to release their payload intracellularly by, for example, conjugating the drug to the polymer backbone via an acid-responsive hydrazone bond and thus taking advantage of the acidic environment, or via a disulfide linker, which is cleaved by the reducing environment within the endocytic pathway, and present in tumors.34−39 Oxidative responsive polymeric NPs are also synthesized, for example, from poly(propylene sulfide), which reacts with hydrogen peroxide to give poly(sulfone) or poly(sulfoxide).31,40 Therefore, oxidative stress causes a hydrophobic to hydrophilic transition in these particles, which leads to drug release. Nanocarrier surface properties also dramatically impact NP behavior as a result of the high surface-to-volume ratios in these systems.1 Surface properties principally impact NP stability, circulation times, clearance, biodistribution, and localization.1 For example, NPs can be directly functionalized with targeting moieties, poly(ethylene glycol) (PEG) stealth layers, or a variety of surfactants. PEG-functionalized NPs are widely studied due to their reduced in vivo clearances compared to non-PEGylated NPs.41 An additional feature of using PEG is that it can be derivatized to contain targeting moieties, proteins, fluorescent dyes, or other molecules.2,19,42,43 The attachment of hydrophilic PEG chains onto the surface of NPs endows the NPs with stealth and biocompatibility as well as colloidal stability as the chains of PEG display a large excluded volume, creating steric repulsion.41 When combined with a targeting moiety, PEGylated NPs thereby avoid nonspecific protein binding and subsequent clearance but specifically bind to their receptor target. The advantages provided by PEG to colloidal drug carriers have prompted scientists to devise schemes for B
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ammonium persulfate and 2 μL of N,N,N′,N′-tetramethylethylenediamine (TEMED) were added to the emulsion postsonication to initiate monomer polymerization. Following polymerization, the suspension was stirred overnight while open to the atmosphere to allow for the evaporation of any remaining solvent. The resulting NPs were dialyzed against 5 mM pH 7.4 phosphate buffer for 24 h to remove excess lipid, salts, and unencapsulated drug. Characterization of NPs. Dynamic light scattering (DLS) measurements were taken on 100 μL aliquots of NPs diluted into 1 mL of 10 mM pH 7.4 phosphate buffer. The zeta potential of the NPs was then measured using a Brookhaven Instruments Corporation ZetaPALS zeta potential analyzer via the Smoluchowski method. The size was determined with a dilution in DI water resulting in ∼50 kcps (roughly 10−50 μL in 3 mL). The results for both the intensity- and number-weighted size distributions were recorded and compared. Size and concentration were measured on an IZON qNano, with nanopore sizes ranging from 100−800 nm. NP suspensions were diluted 1000× and calibrated to the appropriate polystyrene NP standards to determine the absolute sizes of the eNPs. Concentrations were calculated by comparing the change in translocation rate at three discrete pressures (typically 1, 2, and 3 cm of added pressure). By comparing the rate/concentration ratio of a known standard to that of the unknown sample, particle concentration can be determined. All results from the qNano are a combination of at least three separate runs, with at least 200 particle translocations each, run at the same voltage, nanopore stretch, and appropriate pressures. Scanning electron microscopy (SEM) was performed using a Zeiss SUPRA 55VP field emission SEM. Samples were prepared by diluting the NP suspension 100−500× in deionized water, which was dropped on a silicon wafer and allowed to air-dry overnight. All samples were coated with ∼5 nm of Au/Pd prior to imaging and imaged at an accelerating voltage of 2 keV. Images were processed using NIH ImageJ. Osmium Tetroxide Vapor Staining for Transmission Electron Microscopy. PEG-L-eNPs and eNPs were visualized using transmission electron microscopy (TEM; JEM-1010, JEOL) at an accelerating voltage of 80 keV. Particles were dried on 300-mesh copper-coated carbon grids (Electron Microscopy Sciences) and stained with 1% osmium tetroxide vapor (OsO4, Electron Microscopy Sciences) for 10 min. β-BODIPY Fluorescence. PEG-L-eNPs, eNPs, and lecithin micelles were incubated with 2.5 μg/mL of β-BODIPY 500/510 C12-HPC (2-(4,4-difluoro-5-methyl-4-bora-3a,4a-diaza-s-indacene-3dodecanoyl)-1-hexadecanoyl-sn-glycero-3-phosphocholine) (Molecular Probes), a lipophilic fluorescent dye that localizes within lipid membranes.55 Lecithin was dissolved in water at concentrations above (5 mg/mL) and below (0.1 mg/mL) its critical micelle concentration (CMC). PEG-L-eNPs were diluted to a concentration below the CMC of the incorporated lecithin, and the eNPs were diluted analogously. The treated samples were excited at 475 nm, and fluorescence spectra were read between 500−650 nm using a SpectraMax M3 spectrophotometer (Molecular Devices). Measurement of in Vitro Release Kinetics. Release medium was prepared using either 10 mM pH 5.0 acetate buffer or 10 mM pH 7.4 phosphate buffer containing 0.3 wt %/wt SDS following a published procedure.48 Pax-loaded PEG-L-eNPs, Pax-PEG-L-eNPs, (80 μg Pax equivalent) were diluted with 1.5 mL of release medium and placed into 10 000 molecular weight cutoff dialysis tubing. The tubing was then placed into 1 L of release medium with stirring. At 0, 1, 2, 4, 8, 12, and 24 h, 100 μL samples were withdrawn from the dialysis tubing, and Pax content was determined by high-performance liquid chromatography (HPLC). The experiment was performed in triplicate. Cell Culture. KB human epidermal carcinoma cells and A549 human lung carcinoma cells (ATCC) were cultured in Roswell Park Memorial Institute (RPMI) 1640 medium without folic acid (Gibco) and supplemented with 10% FBS and 1% penicillin−streptomycin. Cells were maintained at 37 °C with 5% CO2 in a humidified environment. In Vitro Cellular Uptake. KB and A549 cells were seeded in sixwell plates at a density of 5 × 104 cells/well and allowed to adhere for
dose-dependent toxicity against several cell lines in vitro as well as significant improvements in survival and delayed recurrence in murine tumor models.48−54 To date, the eNPs have been coated with sodium dodecyl sulfate (SDS), a surfactant that cannot be derivatized easily with targeting ligands. Coating eNPs with a lipid monolayer instead of SDS may allow for improved functionality and versatility, particularly in serumcontaining media. In this study, we use soy lecithin and PEGylated 1,2-distearoyl-sn-glycero-3-phosphoethanolamineN-[amino(polyethylene glycol)-2000] (DSPE-PEG-2k) with and without folic acid at the terminal end (Figure 1; Table 1). Table 1. eNP, PEG-L-eNP, and FA-PEG-L-eNP Formulations and Zeta Potentials
We herein combine the controlled triggered swelling of the eNPs with the biocompatibility, stability, and cellular targeting provided by the lipid monolayer. Specifically, we describe the (1) synthesis of eNPs possessing a lipid coating (i.e., PEG-LeNP); (2) characterization of the lipid coating; (3) characterization of the triggered swelling response in acidic medium and its comparison to SDS-coated eNPs; (4) drug release kinetics of paclitaxel (Pax) loaded PEG-L-eNPs; (5) inhibition of aggregation provided by the PEG/lipid shell resulting in increased colloidal stability; (6) synthesis of folic acid targeted conjugated PEG-L-eNPs (FA-PEG-L-eNPs); (7) uptake of the PEG-L-eNPs and FA-PEG-L-eNPs in KB human epidermal and A549 human lung carcinoma cells; (8) dependence of cell uptake and targeting on particle concentration; and (9) increased potency when Pax is loaded within FA-PEG-LeNPs compared to nontargeted PEG-L-eNPs.
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EXPERIMENTAL SECTION
Materials and Instrumentation. The lipid shell was composed of soy lecithin (Avanti), DSPE-PEG-2k (Lipoid GmbH), and 1,2distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-3400]-Folic acid (DSPE-PEG-3k-FA) (Nanocs). The eNP monomer and cross-linker were synthesized as previously described.48 All other chemicals were purchased from Sigma-Aldrich and used without further purification unless otherwise stated. All reactions were performed under nitrogen atmosphere. NMR spectra were recorded on a Varian Mercury spectrometer operating at 400 MHz, and all chemical shifts are calibrated against the residual solvent signals of CDCl3 (δ 7.26 ppm) and reported in ppm. NP Preparation. PEG-L-eNPs were prepared using a modification of a miniemulsion polymerization technique previously described and combined with a one-step method of lipid−polymer hybrid NP preparation.18,48 Nontargeted PEG-L-eNPs were prepared with a weighed amount of lecithin and DSPE-PEG-2k (molar ratio = 80:20) dispersed in 10 mM pH 7.4 phosphate buffer. Targeted FA-PEG-LeNPs were similarly prepared using lecithin, DSPE-PEG-2k, and DSPE-PEG-3k-FA (molar ratio = 80:19:1). The pH-responsive monomer and cross-linker, as well as Pax for drug-loaded eNPs (at 5 wt %), or rhodamine-methacrylate (at 0.01 wt %) for fluorescently labeled eNPs, were dispersed in the oil phase (5:1 acetonitrile/ dichloromethane). The aqueous phase was then added to the oil phase and sonicated in a water bath and under an argon blanket using a probe ultrasonicator (Sonics and Materials). Twenty μL of 200 mM C
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Biomacromolecules 24 h at 37 °C in 5% CO2. Cells were incubated with rhodaminelabeled FA-PEG-L-eNP and PEG-L-eNP for 1, 2, 4, and 24 h at concentrations of 1.69 × 107, 8.45 × 106, and 1.69 × 106 cells/well. Cells were then rinsed with phosphate buffered saline (PBS) and fixed with 4% formaldehyde for 20 min, and fluorescence was measured via a FACScalibur flow cytometer. Confocal Microscopy. KB and A549 cells were seeded in six-well plates at a density of 5 × 104 cells/well and allowed to adhere for 24 h onto coverslips at 37 °C in 5% CO2. After 4 h, the coverslips were washed two times with cold PBS, two times with cold PBS+ (containing Ca2+ and Mg2+), and fixed in 4% formaldehyde for 20 min. Cell membranes were stained with 100 μg/mL Concanavalin A633 conjugate, and nuclei were stained with 3 μg/mL of Hoechst trihydrochlorine trihydrate for 8 min. The coverslips were washed two times with cold PBS, two times with cold PBS+, and mounted with Prolong Gold Anti-Fade. Images were captured with an inverted confocal laser scanning microscope (Olympus IX81) and analyzed using Olympus Fluoroview Version 2.0b software. In Vitro Cell Cytotoxicity. KB cells were cultured in 96-well plates at 5.0 × 103 cells/well for 24 h. The culture media was removed and replaced with media containing the appropriate treatment. The treatment was then removed and replaced with fresh media at 4 or 24 h, and cell viability was assessed 72 h after treatment via the MTS in vitro cytotoxicity assay (CellTiter 96 Aqueous One, Promega) as described previously.53,56
Figure 3. TEM of osmium tetroxide stained PEG-L-eNP with a distinct lipid monolayer.
stain used to provide contrast for lipid membranes, was also used to stain SDS-coated eNPs. However, because of the absence of a lipid layer, the eNPs do not retain the lipid stain, and contrast is only observed as a result of the size of the NP core (Figure S1, Supporting Information). Additionally, fluorescence experiments with β-BODIPY 500/510 C12-HPC show an intensity increase and slight red shift in its emission when incubated with lecithin micelles or PEG-L-eNPs, indicating the incorporation of the dye into a lipid membrane (Figure 4). These observations are not observed in the presence of either the eNPs or lecithin below its CMC.
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RESULTS AND DISCUSSION Size and Zeta Potential Characterization. The synthesis of PEG-L-eNPs was modified from the previously established miniemulsion synthesis to accommodate the lipid surfactants via a two-step synthesis.18,48,57 The lipid coating was composed of lecithin and DSPE-PEG-2k at a lecithin/DSPE-PEG-2k molar ratio of 80:20. The eNP monomer and cross-linker were dissolved in organic solvent (1:5 dichloromethane/acetonitrile), while the lipids were dissolved in 10 mM pH 7.4 phosphate buffer, heated to 65 °C, and added dropwise to the organic solution. Sonication, polymerization, and purification methods followed our previously published reports using the miniemulsion polymerization procedure.48 PEG-L-eNPs are spherical NPs of approximately 150 nm in diameter as determined by DLS measurements and by SEM (Figure 2). In comparison, eNPs are a more heterogeneous
Figure 4. Emission spectra of BODIPY 500/510 incubated with PEGL-eNPs and eNPs as well as with lecithin above (5 mg/mL) and below (0.1 mg/mL) its CMC. The intensity increase of BODIPY, which occurs when the fluorophore is incorporated into a lipid layer, is only observed when the dye is incubated with PEG-L-eNPs or lecithin micelles.
Figure 2. Scanning electron micrographs of eNPs (left) and PEG-LeNPs (right).
particle population of both 20−50 nm and ∼200 nm particles, with an average diameter of 150 nm. PEG-L-eNPs have a zeta potential of −12 mV compared to a −50 to −60 mV zeta potential for eNPs, which are highly charged due to their SDS surface coating (Table 1). Characterization of Lipid Coating. TEM imaging of osmium tetroxide stained PEG-L-eNPs indicates the presence of a ∼4.68 nm lipid layer, which is consistent with the thickness of a lecithin monolayer (Figure 3).58 Osmium tetroxide, a TEM
Analysis of NP Colloidal Stability. With the addition of PEG to the eNP lipid coating, it was expected that the PEG-LeNPs would have improved colloidal stability in vitro in the presence of serum proteins. Both eNPs and PEG-L-eNPs were incubated in a 10% bovine calf serum solution (diluted into 1× PBS), and their sizes were monitored via DLS over time. The D
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However, in the case of PEG−lipid hybrid NP systems, it is possible for the payload to partition into either the lipid layer or the polymeric core. To assess drug release kinetics, Pax-PEG-LeNPs were synthesized and incubated in both pH 7.4 and pH 5 buffers for 24 h. Pax-PEG-L-eNP samples were collected over time and analyzed for Pax content. Similar release profiles are observed for Pax-PEG-L-eNPs at both pH 5 and pH 7.4, with approximately 40% of the drug payload being released within 24 h (Figure 7). These data suggest that Pax partitions into the
eNPs and PEG-L-eNPs used in this study had initial diameters of ∼150 nm, but the eNPs experience significant aggregation within 30 min as evidenced by an increase in average particle size, while the PEG-L-eNPs maintain their stability over 24 h (Figure 5). The increased stability of the PEG-L-eNPs as
Figure 5. eNP (solid) and PEG-L-eNP (hashed) diameters, as measured by DLS, over a period of 24 h in 10% bovine calf serum. Data represented as mean ± SD; n = 3.
compared to the eNPs is attributed to the steric repulsion afforded by the PEG chains, which reduce aggregation as well as protein adsorption. Additionally, SDS, as a small, single chain surfactant, has lower affinity forand weaker interactions withthe hydrophobic NP core than does lecithin, contributing to the increased colloidal stability of lipid-coated eNPs. Similar results have been reported for other lipid−PEG hybrid NP systems.9,13,59 Analysis of pH-Responsive Swelling. To ensure that the lipid surface coating and modified preparation of the eNPs did not compromise their pH-responsive swelling, eNPs and PEGL-eNPs were incubated in either pH 7.4 or pH 5 buffers for 24 h. Each batch was prepared to have a similar starting diameter of ∼350 nm as measured via qNano, since larger starting NP sizes enable easier detection and monitoring via the qNano.60 After exposure to pH 5 buffer, both eNP formulations experience a two-fold increase in size after 24 h, confirming that there is not a significant difference in swelling between the two eNP formulations (Figure 6). Analysis of Drug Release Kinetics. Previously, we demonstrated that Pax-loaded eNPs release their payload rapidly at pH 5, while they maintain the drug within the core at pH 7.4.48 This behavior parallels the swelling of the core.
Figure 7. Release kinetics of Pax-PEG-L-eNPs exposed to pH 5 (red) or pH 7.4 (blue) buffer for 24 h. Data represent mean ± SD (N = 3).
lipid coating of the hybrid system rather than the eNP core. The release kinetics we report are consistent with those reported by others investigating hybrid lecithin−polymer NP systems.9,61,62 Additionally, Pax is readily able to partition into lipid layers as evidenced by its wide use in liposomal formulations.63−65 Although the pH-responsive swelling of the PEG-L-eNP core does not affect the drug release kinetics, the hydrogel structures formed intracellulary after the uptake of pH-responsive eNPs afford a unique opportunity where the eNPs act as intracellular drug depots.66 Characterization of in Vitro Cellular Uptake. To incorporate cell-specific targeting, 1 mol % of the folic acidconjugated PEGylated lipid, FA-DSPE-PEG-3k, was added to the lipid shell, for an 80:19:1 molar ratio of lecithin/DSPEPEG-2k/FA-DSPE-PEG-3k. To investigate the uptake of FAPEG-L-eNPs into cancer cells, rhodamine (Rho) methacrylate was incorporated into the polymer backbones of PEG-L-eNPs and FA-PEG-L-eNPs for detection of eNP uptake in cells via flow cytometry and confocal microscopy. Since polymer mass is not equally distributed among NPs, a NP concentration-based assessment of uptake was chosen. Until now, it has been difficult to accurately measure the concentration of NP suspensions. Recently, the qNano, which is a scanning ion occlusion sensing instrument (SIOS), was developed for accurate NP characterization including size, surface charge, and particle concentration.60,67 With this technique, doses of NPs can be quantified for more consistent results in in vitro and in vivo assays. Folate receptor positive (FR+) human epidermal KB cells and folate receptor negative (FR−) A549 cells were treated with 1.69 × 107, 8.45 × 106, and 1.69 × 106 particles/mL, as determined via qNano.60,67 Uptake was assessed via flow cytometry at 0, 1, 4, and 24 h after treatment. In KB cells, uptake of targeted NPs increases with time, while untargeted
Figure 6. Swelling comparison of eNPs (solid) and PEG-L-eNPs (hashed) over 24 h in either pH 7.4 (blue) or pH 5 (red) buffer. Data are represented as mean of four separate runs with >500 events recorded per run. E
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Biomacromolecules PEG-L-eNPs exhibit less significant increases in uptake over time (Figure 8). At 1 and 4 h, FR+ KB cells show significantly
Figure 10. Confocal microscopy confirms uptake of Rho labeled FAPEG-L-eNPs (top) and PEG-L-eNPs (bottom) after 4 h of incubation. KB cells were costained with DAPI and Concanavalin-A to mark the nucleus and cell membrane, respectively. Rho-eNP aggregates are observed within the cells and are shown in red.
Figure 8. Uptake of rhodamine-labeled (A) PEG-L-eNPs and (B) FAPEG-L-eNPs at a concentration of 8.45 × 106 particles/mL in FR+ KB cells as quantified by flow cytometry.
more FA-PEG-L-eNP uptake than PEG-L-eNP uptake at all three concentrations (Figure 9). At 24 h, FR+ KB cells exhibit significantly enhanced uptake of FA-PEG-L-eNPs at the two lower NP concentrations. However, at the highest NP concentration, at 24 h, there is not a significant difference between uptake of targeted and nontargeted particles, which suggests saturation of folate receptors at this concentration within 24 h. In the FR− A549 cells, there is no significant difference between the uptake of targeted eNPs and nontargeted eNPs at all three time points, consistent with the lack of folate receptors on these cells. Similarly, uptake was assessed via confocal microscopy. KB cells were treated with Rho-labeled FA-PEG-L-eNPs and Rholabeled PEG-L-eNPs at a concentration of 1.69 × 107 particles/ mL, and imaged 4 h after treatment (Figure 10). Confocal images reveal that KB cells exhibit enhanced uptake of targeted NPs compared to nontargeted NPs, and z-stack images show that FA-PEG-L-eNPs are localized within the cells. In Vitro Cytotoxicity Analysis. Previously, nontargeted Pax-loaded eNPs demonstrated efficacy in multiple cell lines in vitro including breast, lung, ovarian, and mesothelioma cancers.48−54 To assess the toxicity of the lipid-coated eNPs in a folate-receptor expressing cell line, Pax-loaded PEG-LeNPs and FA-PEG-L-eNPs were incubated with KB cells and compared with Pax dissolved in cremophor/ethanol (Pax-C/E) at 4 and 24 h of incubation (Figure 11). Since Pax-PEG-L-eNP uptake varies minimally between 4 and 24 h (Figure 8), these particles exhibit similar potency at 4 and 24 h, with IC50 values of 26.87 and 23.11 ng/mL, respectively. The effect of Pax-FA-
Figure 11. In vitro cytotoxicity of eNP formulations. Folate receptor overexpressing KB cells were treated for 4 or 24 h with Pax-PEG-LeNPs, Pax-FA-PEG-L-eNPs, or Pax-C/E, and viability was assessed 72 h after treatment. Drug free NP formulations did not affect cell viability (data not shown). Data displayed as mean ± SD; n = 6−8.
PEG-L-eNPs at 4 h is similar to that of Pax-PEG-L-eNPs, with an IC50 of 20.42 ng/mL. Although the uptake of FA-targeted particles is significantly greater than that of nontargeted particles at 4 h (Figure 9), this difference does not translate into an effect on the potency of the NP formulation. However, at 24 h, Pax-FA-PEG-L-eNPs are more potent than Pax-PEG-LeNPs, with an IC50 of 12.87 ng/mL. Unloaded NP formulations do not affect cell viability (data not shown). At 4 h, the potency
Figure 9. Concentration-based assessment of PEG-L-eNP and FA-PEG-L-eNP uptake in (A) FR+ KB cells and (B) FR− A549 cells, as determined by flow cytometry. Data are displayed as mean ± SD (n = 4 for KB cells; n = 3 for A549 cells). ∗, p < 0.05 for each pair at discrete time points and concentrations. #, no significant difference. F
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CONCLUSION The lipid−polymer hybrid NP system presented here combines the pH-responsive swelling of the previously described eNPs with the improved colloidal stability and surface functionality of liposomes. Folic-acid conjugated PEG-L-eNPs demonstrate enhanced uptake in vitro in a folate receptor expressing cell line, and the addition of PEGylated lipids allows for significantly improved colloidal stability in serum relative to SDS-coated eNPs. When loaded with the anticancer agent, paclitaxel, the resulting hybrid polymer/lipid nanocarriers exhibit greater potency compared to nontargeted PEG-L-eNPs. Importantly, the swelling behavior of the eNP polymer core is maintained, which affords the unique opportunity to form intracellular hydrogel structures. Hybrid polymer−lipid NP systems offer opportunities for improved stability and enable active targeting via the lipid surface coating while maintaining the advantages of a polymeric core.
of the clinical formulation of Pax, Pax-C/E, is the lowest, with an IC50 of 111.2 ng/mL. At 24 h, Pax-C/E is the most potent, with an IC50 of 1.99 ng/mL. This result is dependent on the experimental conditions, as the amount of drug uptake (or efflux) over time will affect efficacy. Targeting ligands, including antibodies, peptides, and small molecules, are incorporated into NP coatings to improve cell uptake along with PEGylation to prevent nonspecific protein.19,68,69 Many receptors and surface proteins are overexpressed in cancer cells relative to healthy cells and serve well as targets for specific drug delivery.10,19,70 The folate receptor is one such target, which possesses a high affinity for folic acid.10,15,17,19,70−74 In agreement with the results reported here, previous studies report enhanced uptake and potency of folate-targeted drug-loaded entities, such as micellar paclitaxel, both in vitro as well as in vivo.15,75−78 It is also known that avidity varies as a function of cellular receptor density as well as NP ligand valency.79 Additionally, folate-targeted NPs and imaging probes are being explored in in vivo imaging applications as a result of their enhanced uptake and association with FR+ cancer cells, further documenting its beneficial use.80−82 However, since the endocytosis of folate targeted entities occurs through recycling centers characterized by pH ∼6.5, the extent of FA-PEG-L-eNP expansion will be reduced, as eNP swelling is pH-dependent.83 The new hybrid PEG-L-eNPs explored in this study are prepared using a minor modification of the miniemulsion polymerization technique and add to the collection of previously prepared and characterized hybrid lipid−polymer assemblies.1,9,13,45,59,84,85 Given the diversity of polymeric NPs successfully coated with lipids, the lipid-coating strategy is likely general and applicable to various NP systems. Several factors must be considered in the development of hybrid lipid−PEG NP systems with a nonresponsive or responsive core. For example, the formulation components (e.g., ratio of lecithin/ DSPE-PEG/DSPE-PEG-FA) are known to affect release kinetics, NP uptake, and colloidal stability; optimization studies should be undertaken after the initial development of a new system. With regards to the responsive core in our studies, the pH-responsive swelling of the polymeric core is not affected by the lipid−PEG coating, and the drug release kinetics are not affected by the pH conditions. These results suggest that Pax partitions primarily into the lipid layer of the PEG-L-eNPs and not the core. Therefore, it is important to consider the partitioning of the payload into the core versus the coating when designing future hybrid polymer−lipid NP systems, as this may critically affect the release kinetics. These results also highlight the potential opportunity, for those working in this area, to use the polymer core as well as the lipid shell for drug loading and thus engineer NP systems with multiple payloads or multimodal release kinetics. Finally, the selection of the targeting agent is an important consideration for cellular uptake, NP function, and drug release. For example, with folate targeting, in contrast to antibody mediated targeting or passive targeting, endocytosis occurs with a minimal reduction in pH to ∼6.5 as opposed to a pH of ∼5.83 While folate is one of the most extensively used ligands for targeted drug delivery, the incorporation of other targeting moieties, aptamers, or antibodies can be explored and has been described by others.1,13,59,85
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ASSOCIATED CONTENT
S Supporting Information *
Transmission electron micrograph of eNP. The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.biomac.5b00336.
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AUTHOR INFORMATION
Corresponding Authors
*E-mail: [email protected]. Phone: (617)-358-3429. Fax: (617)358-3186. *E-mail: [email protected]. Phone: (617)-353-7366. Fax: (617)353-5866. Author Contributions
M.S. and I.E. contributed equally to this work. 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 This work was supported in part by funding from National Science Foundation (DMR-1006601, DGE-1247312), Boston University’s Nanomedicine Program and Cross-Disciplinary Training in Nanotechnology for Cancer (NIH R25 CA153955), Brigham and Women’s Hospital, and Boston University.
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Synthesis and Characterization of Hybrid Polymer/Lipid Expansile Nanoparticles: Imparting Surface Functionality for Targeting and Stability Michelle Stolzoff,† Iriny Ekladious,† Aaron H. Colby,† Yolonda L. Colson,§ Tyrone M. Porter,*,‡ and Mark W. Grinstaff*,† †
Departments of Biomedical Engineering and Chemistry and ‡Departments of Mechanical Engineering and Biomedical Engineering, Boston University, Boston, Massachusetts 02215, United States § Division of Thoracic Surgery, Department of Surgery, Brigham and Women’s Hospital, Boston, Massachusetts 02115, United States S Supporting Information *
ABSTRACT: The size, drug loading, drug release kinetics, localization, biodistribution, and stability of a given polymeric nanoparticle (NP) system depend on the composition of the NP core as well as its surface properties. In this study, novel, pH-responsive, and lipid-coated NPs, which expand in size from a diameter of approximately 100 to 1000 nm in the presence of a mildly acidic pH environment, are synthesized and characterized. Specifically, a combined miniemulsion and free-radical polymerization method is used to prepare the NPs in the presence of PEGylated lipids. These PEGylated-lipid expansile NPs (PEG-L-eNPs) combine the swelling behavior of the polymeric core of expansile NPs with the improved colloidal stability and surface functionality of PEGylated liposomes. The surface functionality of PEG-L-eNPs allows for the incorporation of folic acid (FA) and folate receptor-targeting. The resulting hybrid polymer/lipid nanocarriers, FA-PEG-L-eNPs, exhibit greater in vitro uptake and potency when loaded with paclitaxel compared to nontargeted PEG-L-eNPs.
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INTRODUCTION Nanoparticle (NP) drug delivery systems possess several advantages over conventional small molecule therapies.1−7 Among these is the potential to provide controlled release as well as sustained levels of drug within cells, thereby conferring greater efficacy and reduced toxicity. As a result of their size, surface properties, and core properties, NPs can be used to dramatically alter the pharmacokinetics of entrapped drugs. For example, entrapping doxorubicin within PEGylated liposomes increases the circulation time and cumulative maximum tolerated dose of the drug.8 Polymeric NPs are especially attractive in cancer treatment because of their ability to encapsulate water insoluble agents, provide controlled drug release rates, and target tumors.2,9 Additionally, NPs can increase the maximum tolerated dose of encapsulated drugs by circumventing the solubility limitations of certain anticancer agents.1 The composition of the NP polymeric core dictates important features such as drug loading, NP size, and drug release kinetics, while surface properties affect NP localization, biodistribution, and stability. The selection of the polymer for the core is critical. Currently, poly(lactic acid) (PLA) and poly(lactic-co-glycolic) (PLGA) polymers are the most widely studied polymers in nanocarrier systems due to their biocompatibility, biodegradability, and availability.9−14 As such, PLGA and PLA are used © XXXX American Chemical Society
for the encapsulation and delivery of a wide range of chemotherapeutic agents including paclitaxel, doxorubicin, and docetaxel.15−19 However, PLGA and PLA core systems are criticized for their relatively rapid “burst” release of encapsulated drug.20,21 It was previously demonstrated that half of the encapsulated agent in a PLGA−PEG NP system can be released within minuteswell before any significant tumor accumulation would occur in vivo.22,23 Consequently, burst release from circulating particles may compromise the effectiveness of the anticancer agent and increase the risk of systemic toxicities. To address this limitation, scientists are engineering stimuli-responsive polymeric NPs to achieve triggered drug release within solid tumors. Potential triggers for release include enzymatic activity, acidity, light activation, temperature, and other chemical or physical stimuli.24−31 For example, β-amino ester linkages within a polymer confer pH sensitivity, as the protonation of tertiary amines in these polymersnear pH 6.5causes the polymer core to become soluble in water where it was previously insoluble at pH 7.4.31,32 The relatively acidic pH of tumor extracellular fluid is well-documented and is now considered to be a phenotype of Received: March 12, 2015 Revised: May 30, 2015
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Figure 1. Schematic of eNP surface coatings and pH-triggered swelling.
PEGylating the surface of polymeric NPs. For example, NPs can be formulated from diblock copolymers composed of PEG and a polyester. However, the copolymer must be synthesized, and it is difficult to control the molecular weight of each block to minimize polydispersity and batch variability. As an alternative approach, Park et al. coated polymeric NPs with lipids functionalized with avidin and then conjugated biotinylated PEG to the shell.44 Unfortunately, avidin is immunogenic, which prompted the exploration of alternative linkers for PEG. PEG−lipid conjugates with more suitable chemistries have been synthesized and incorporated in the formulation of hybrid lipid−polymer assemblies for various biotechnological and biomedical applications.9,45−47 Additionally, PEG terminated with a targeting moiety can improve the specificity with which lipid−polymer hybrid assemblies deliver drugs.17 Thus, PEGylated lipid coatings on polymeric NPs serve multiple purposes with respect to drug delivery, and we sought to explore this further with our polymer-based NP platform. We have previously engineered a novel pH-responsive expansile NP (eNP) that expands from 100 to 1000 nm under mildly acidic conditions, such as those found in the endosome and lysosome, resulting in intracellular release of the encapsulated payload.48 The eNPs are synthesized from a monomer that has a 2,4,6-trimethoxybenzylidene acetal protecting group that is stable at neutral pH but hydrolyzes under mildly acidic conditions (pH ∼5). Hydrolysis reveals the once protected hydroxyl groups, which increases the hydrophilicity of the polymeric core. As a result, eNPs swell in aqueous solutions and rapidly release the entrapped payload (Figure 1). Previously, paclitaxel-loaded eNPs demonstrated
solid tumors, making pH-sensitivity a viable route of NP targeting to tumors.33 NPs can also be designed to release their payload intracellularly by, for example, conjugating the drug to the polymer backbone via an acid-responsive hydrazone bond and thus taking advantage of the acidic environment, or via a disulfide linker, which is cleaved by the reducing environment within the endocytic pathway, and present in tumors.34−39 Oxidative responsive polymeric NPs are also synthesized, for example, from poly(propylene sulfide), which reacts with hydrogen peroxide to give poly(sulfone) or poly(sulfoxide).31,40 Therefore, oxidative stress causes a hydrophobic to hydrophilic transition in these particles, which leads to drug release. Nanocarrier surface properties also dramatically impact NP behavior as a result of the high surface-to-volume ratios in these systems.1 Surface properties principally impact NP stability, circulation times, clearance, biodistribution, and localization.1 For example, NPs can be directly functionalized with targeting moieties, poly(ethylene glycol) (PEG) stealth layers, or a variety of surfactants. PEG-functionalized NPs are widely studied due to their reduced in vivo clearances compared to non-PEGylated NPs.41 An additional feature of using PEG is that it can be derivatized to contain targeting moieties, proteins, fluorescent dyes, or other molecules.2,19,42,43 The attachment of hydrophilic PEG chains onto the surface of NPs endows the NPs with stealth and biocompatibility as well as colloidal stability as the chains of PEG display a large excluded volume, creating steric repulsion.41 When combined with a targeting moiety, PEGylated NPs thereby avoid nonspecific protein binding and subsequent clearance but specifically bind to their receptor target. The advantages provided by PEG to colloidal drug carriers have prompted scientists to devise schemes for B
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ammonium persulfate and 2 μL of N,N,N′,N′-tetramethylethylenediamine (TEMED) were added to the emulsion postsonication to initiate monomer polymerization. Following polymerization, the suspension was stirred overnight while open to the atmosphere to allow for the evaporation of any remaining solvent. The resulting NPs were dialyzed against 5 mM pH 7.4 phosphate buffer for 24 h to remove excess lipid, salts, and unencapsulated drug. Characterization of NPs. Dynamic light scattering (DLS) measurements were taken on 100 μL aliquots of NPs diluted into 1 mL of 10 mM pH 7.4 phosphate buffer. The zeta potential of the NPs was then measured using a Brookhaven Instruments Corporation ZetaPALS zeta potential analyzer via the Smoluchowski method. The size was determined with a dilution in DI water resulting in ∼50 kcps (roughly 10−50 μL in 3 mL). The results for both the intensity- and number-weighted size distributions were recorded and compared. Size and concentration were measured on an IZON qNano, with nanopore sizes ranging from 100−800 nm. NP suspensions were diluted 1000× and calibrated to the appropriate polystyrene NP standards to determine the absolute sizes of the eNPs. Concentrations were calculated by comparing the change in translocation rate at three discrete pressures (typically 1, 2, and 3 cm of added pressure). By comparing the rate/concentration ratio of a known standard to that of the unknown sample, particle concentration can be determined. All results from the qNano are a combination of at least three separate runs, with at least 200 particle translocations each, run at the same voltage, nanopore stretch, and appropriate pressures. Scanning electron microscopy (SEM) was performed using a Zeiss SUPRA 55VP field emission SEM. Samples were prepared by diluting the NP suspension 100−500× in deionized water, which was dropped on a silicon wafer and allowed to air-dry overnight. All samples were coated with ∼5 nm of Au/Pd prior to imaging and imaged at an accelerating voltage of 2 keV. Images were processed using NIH ImageJ. Osmium Tetroxide Vapor Staining for Transmission Electron Microscopy. PEG-L-eNPs and eNPs were visualized using transmission electron microscopy (TEM; JEM-1010, JEOL) at an accelerating voltage of 80 keV. Particles were dried on 300-mesh copper-coated carbon grids (Electron Microscopy Sciences) and stained with 1% osmium tetroxide vapor (OsO4, Electron Microscopy Sciences) for 10 min. β-BODIPY Fluorescence. PEG-L-eNPs, eNPs, and lecithin micelles were incubated with 2.5 μg/mL of β-BODIPY 500/510 C12-HPC (2-(4,4-difluoro-5-methyl-4-bora-3a,4a-diaza-s-indacene-3dodecanoyl)-1-hexadecanoyl-sn-glycero-3-phosphocholine) (Molecular Probes), a lipophilic fluorescent dye that localizes within lipid membranes.55 Lecithin was dissolved in water at concentrations above (5 mg/mL) and below (0.1 mg/mL) its critical micelle concentration (CMC). PEG-L-eNPs were diluted to a concentration below the CMC of the incorporated lecithin, and the eNPs were diluted analogously. The treated samples were excited at 475 nm, and fluorescence spectra were read between 500−650 nm using a SpectraMax M3 spectrophotometer (Molecular Devices). Measurement of in Vitro Release Kinetics. Release medium was prepared using either 10 mM pH 5.0 acetate buffer or 10 mM pH 7.4 phosphate buffer containing 0.3 wt %/wt SDS following a published procedure.48 Pax-loaded PEG-L-eNPs, Pax-PEG-L-eNPs, (80 μg Pax equivalent) were diluted with 1.5 mL of release medium and placed into 10 000 molecular weight cutoff dialysis tubing. The tubing was then placed into 1 L of release medium with stirring. At 0, 1, 2, 4, 8, 12, and 24 h, 100 μL samples were withdrawn from the dialysis tubing, and Pax content was determined by high-performance liquid chromatography (HPLC). The experiment was performed in triplicate. Cell Culture. KB human epidermal carcinoma cells and A549 human lung carcinoma cells (ATCC) were cultured in Roswell Park Memorial Institute (RPMI) 1640 medium without folic acid (Gibco) and supplemented with 10% FBS and 1% penicillin−streptomycin. Cells were maintained at 37 °C with 5% CO2 in a humidified environment. In Vitro Cellular Uptake. KB and A549 cells were seeded in sixwell plates at a density of 5 × 104 cells/well and allowed to adhere for
dose-dependent toxicity against several cell lines in vitro as well as significant improvements in survival and delayed recurrence in murine tumor models.48−54 To date, the eNPs have been coated with sodium dodecyl sulfate (SDS), a surfactant that cannot be derivatized easily with targeting ligands. Coating eNPs with a lipid monolayer instead of SDS may allow for improved functionality and versatility, particularly in serumcontaining media. In this study, we use soy lecithin and PEGylated 1,2-distearoyl-sn-glycero-3-phosphoethanolamineN-[amino(polyethylene glycol)-2000] (DSPE-PEG-2k) with and without folic acid at the terminal end (Figure 1; Table 1). Table 1. eNP, PEG-L-eNP, and FA-PEG-L-eNP Formulations and Zeta Potentials
We herein combine the controlled triggered swelling of the eNPs with the biocompatibility, stability, and cellular targeting provided by the lipid monolayer. Specifically, we describe the (1) synthesis of eNPs possessing a lipid coating (i.e., PEG-LeNP); (2) characterization of the lipid coating; (3) characterization of the triggered swelling response in acidic medium and its comparison to SDS-coated eNPs; (4) drug release kinetics of paclitaxel (Pax) loaded PEG-L-eNPs; (5) inhibition of aggregation provided by the PEG/lipid shell resulting in increased colloidal stability; (6) synthesis of folic acid targeted conjugated PEG-L-eNPs (FA-PEG-L-eNPs); (7) uptake of the PEG-L-eNPs and FA-PEG-L-eNPs in KB human epidermal and A549 human lung carcinoma cells; (8) dependence of cell uptake and targeting on particle concentration; and (9) increased potency when Pax is loaded within FA-PEG-LeNPs compared to nontargeted PEG-L-eNPs.
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EXPERIMENTAL SECTION
Materials and Instrumentation. The lipid shell was composed of soy lecithin (Avanti), DSPE-PEG-2k (Lipoid GmbH), and 1,2distearoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethylene glycol)-3400]-Folic acid (DSPE-PEG-3k-FA) (Nanocs). The eNP monomer and cross-linker were synthesized as previously described.48 All other chemicals were purchased from Sigma-Aldrich and used without further purification unless otherwise stated. All reactions were performed under nitrogen atmosphere. NMR spectra were recorded on a Varian Mercury spectrometer operating at 400 MHz, and all chemical shifts are calibrated against the residual solvent signals of CDCl3 (δ 7.26 ppm) and reported in ppm. NP Preparation. PEG-L-eNPs were prepared using a modification of a miniemulsion polymerization technique previously described and combined with a one-step method of lipid−polymer hybrid NP preparation.18,48 Nontargeted PEG-L-eNPs were prepared with a weighed amount of lecithin and DSPE-PEG-2k (molar ratio = 80:20) dispersed in 10 mM pH 7.4 phosphate buffer. Targeted FA-PEG-LeNPs were similarly prepared using lecithin, DSPE-PEG-2k, and DSPE-PEG-3k-FA (molar ratio = 80:19:1). The pH-responsive monomer and cross-linker, as well as Pax for drug-loaded eNPs (at 5 wt %), or rhodamine-methacrylate (at 0.01 wt %) for fluorescently labeled eNPs, were dispersed in the oil phase (5:1 acetonitrile/ dichloromethane). The aqueous phase was then added to the oil phase and sonicated in a water bath and under an argon blanket using a probe ultrasonicator (Sonics and Materials). Twenty μL of 200 mM C
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Biomacromolecules 24 h at 37 °C in 5% CO2. Cells were incubated with rhodaminelabeled FA-PEG-L-eNP and PEG-L-eNP for 1, 2, 4, and 24 h at concentrations of 1.69 × 107, 8.45 × 106, and 1.69 × 106 cells/well. Cells were then rinsed with phosphate buffered saline (PBS) and fixed with 4% formaldehyde for 20 min, and fluorescence was measured via a FACScalibur flow cytometer. Confocal Microscopy. KB and A549 cells were seeded in six-well plates at a density of 5 × 104 cells/well and allowed to adhere for 24 h onto coverslips at 37 °C in 5% CO2. After 4 h, the coverslips were washed two times with cold PBS, two times with cold PBS+ (containing Ca2+ and Mg2+), and fixed in 4% formaldehyde for 20 min. Cell membranes were stained with 100 μg/mL Concanavalin A633 conjugate, and nuclei were stained with 3 μg/mL of Hoechst trihydrochlorine trihydrate for 8 min. The coverslips were washed two times with cold PBS, two times with cold PBS+, and mounted with Prolong Gold Anti-Fade. Images were captured with an inverted confocal laser scanning microscope (Olympus IX81) and analyzed using Olympus Fluoroview Version 2.0b software. In Vitro Cell Cytotoxicity. KB cells were cultured in 96-well plates at 5.0 × 103 cells/well for 24 h. The culture media was removed and replaced with media containing the appropriate treatment. The treatment was then removed and replaced with fresh media at 4 or 24 h, and cell viability was assessed 72 h after treatment via the MTS in vitro cytotoxicity assay (CellTiter 96 Aqueous One, Promega) as described previously.53,56
Figure 3. TEM of osmium tetroxide stained PEG-L-eNP with a distinct lipid monolayer.
stain used to provide contrast for lipid membranes, was also used to stain SDS-coated eNPs. However, because of the absence of a lipid layer, the eNPs do not retain the lipid stain, and contrast is only observed as a result of the size of the NP core (Figure S1, Supporting Information). Additionally, fluorescence experiments with β-BODIPY 500/510 C12-HPC show an intensity increase and slight red shift in its emission when incubated with lecithin micelles or PEG-L-eNPs, indicating the incorporation of the dye into a lipid membrane (Figure 4). These observations are not observed in the presence of either the eNPs or lecithin below its CMC.
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RESULTS AND DISCUSSION Size and Zeta Potential Characterization. The synthesis of PEG-L-eNPs was modified from the previously established miniemulsion synthesis to accommodate the lipid surfactants via a two-step synthesis.18,48,57 The lipid coating was composed of lecithin and DSPE-PEG-2k at a lecithin/DSPE-PEG-2k molar ratio of 80:20. The eNP monomer and cross-linker were dissolved in organic solvent (1:5 dichloromethane/acetonitrile), while the lipids were dissolved in 10 mM pH 7.4 phosphate buffer, heated to 65 °C, and added dropwise to the organic solution. Sonication, polymerization, and purification methods followed our previously published reports using the miniemulsion polymerization procedure.48 PEG-L-eNPs are spherical NPs of approximately 150 nm in diameter as determined by DLS measurements and by SEM (Figure 2). In comparison, eNPs are a more heterogeneous
Figure 4. Emission spectra of BODIPY 500/510 incubated with PEGL-eNPs and eNPs as well as with lecithin above (5 mg/mL) and below (0.1 mg/mL) its CMC. The intensity increase of BODIPY, which occurs when the fluorophore is incorporated into a lipid layer, is only observed when the dye is incubated with PEG-L-eNPs or lecithin micelles.
Figure 2. Scanning electron micrographs of eNPs (left) and PEG-LeNPs (right).
particle population of both 20−50 nm and ∼200 nm particles, with an average diameter of 150 nm. PEG-L-eNPs have a zeta potential of −12 mV compared to a −50 to −60 mV zeta potential for eNPs, which are highly charged due to their SDS surface coating (Table 1). Characterization of Lipid Coating. TEM imaging of osmium tetroxide stained PEG-L-eNPs indicates the presence of a ∼4.68 nm lipid layer, which is consistent with the thickness of a lecithin monolayer (Figure 3).58 Osmium tetroxide, a TEM
Analysis of NP Colloidal Stability. With the addition of PEG to the eNP lipid coating, it was expected that the PEG-LeNPs would have improved colloidal stability in vitro in the presence of serum proteins. Both eNPs and PEG-L-eNPs were incubated in a 10% bovine calf serum solution (diluted into 1× PBS), and their sizes were monitored via DLS over time. The D
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However, in the case of PEG−lipid hybrid NP systems, it is possible for the payload to partition into either the lipid layer or the polymeric core. To assess drug release kinetics, Pax-PEG-LeNPs were synthesized and incubated in both pH 7.4 and pH 5 buffers for 24 h. Pax-PEG-L-eNP samples were collected over time and analyzed for Pax content. Similar release profiles are observed for Pax-PEG-L-eNPs at both pH 5 and pH 7.4, with approximately 40% of the drug payload being released within 24 h (Figure 7). These data suggest that Pax partitions into the
eNPs and PEG-L-eNPs used in this study had initial diameters of ∼150 nm, but the eNPs experience significant aggregation within 30 min as evidenced by an increase in average particle size, while the PEG-L-eNPs maintain their stability over 24 h (Figure 5). The increased stability of the PEG-L-eNPs as
Figure 5. eNP (solid) and PEG-L-eNP (hashed) diameters, as measured by DLS, over a period of 24 h in 10% bovine calf serum. Data represented as mean ± SD; n = 3.
compared to the eNPs is attributed to the steric repulsion afforded by the PEG chains, which reduce aggregation as well as protein adsorption. Additionally, SDS, as a small, single chain surfactant, has lower affinity forand weaker interactions withthe hydrophobic NP core than does lecithin, contributing to the increased colloidal stability of lipid-coated eNPs. Similar results have been reported for other lipid−PEG hybrid NP systems.9,13,59 Analysis of pH-Responsive Swelling. To ensure that the lipid surface coating and modified preparation of the eNPs did not compromise their pH-responsive swelling, eNPs and PEGL-eNPs were incubated in either pH 7.4 or pH 5 buffers for 24 h. Each batch was prepared to have a similar starting diameter of ∼350 nm as measured via qNano, since larger starting NP sizes enable easier detection and monitoring via the qNano.60 After exposure to pH 5 buffer, both eNP formulations experience a two-fold increase in size after 24 h, confirming that there is not a significant difference in swelling between the two eNP formulations (Figure 6). Analysis of Drug Release Kinetics. Previously, we demonstrated that Pax-loaded eNPs release their payload rapidly at pH 5, while they maintain the drug within the core at pH 7.4.48 This behavior parallels the swelling of the core.
Figure 7. Release kinetics of Pax-PEG-L-eNPs exposed to pH 5 (red) or pH 7.4 (blue) buffer for 24 h. Data represent mean ± SD (N = 3).
lipid coating of the hybrid system rather than the eNP core. The release kinetics we report are consistent with those reported by others investigating hybrid lecithin−polymer NP systems.9,61,62 Additionally, Pax is readily able to partition into lipid layers as evidenced by its wide use in liposomal formulations.63−65 Although the pH-responsive swelling of the PEG-L-eNP core does not affect the drug release kinetics, the hydrogel structures formed intracellulary after the uptake of pH-responsive eNPs afford a unique opportunity where the eNPs act as intracellular drug depots.66 Characterization of in Vitro Cellular Uptake. To incorporate cell-specific targeting, 1 mol % of the folic acidconjugated PEGylated lipid, FA-DSPE-PEG-3k, was added to the lipid shell, for an 80:19:1 molar ratio of lecithin/DSPEPEG-2k/FA-DSPE-PEG-3k. To investigate the uptake of FAPEG-L-eNPs into cancer cells, rhodamine (Rho) methacrylate was incorporated into the polymer backbones of PEG-L-eNPs and FA-PEG-L-eNPs for detection of eNP uptake in cells via flow cytometry and confocal microscopy. Since polymer mass is not equally distributed among NPs, a NP concentration-based assessment of uptake was chosen. Until now, it has been difficult to accurately measure the concentration of NP suspensions. Recently, the qNano, which is a scanning ion occlusion sensing instrument (SIOS), was developed for accurate NP characterization including size, surface charge, and particle concentration.60,67 With this technique, doses of NPs can be quantified for more consistent results in in vitro and in vivo assays. Folate receptor positive (FR+) human epidermal KB cells and folate receptor negative (FR−) A549 cells were treated with 1.69 × 107, 8.45 × 106, and 1.69 × 106 particles/mL, as determined via qNano.60,67 Uptake was assessed via flow cytometry at 0, 1, 4, and 24 h after treatment. In KB cells, uptake of targeted NPs increases with time, while untargeted
Figure 6. Swelling comparison of eNPs (solid) and PEG-L-eNPs (hashed) over 24 h in either pH 7.4 (blue) or pH 5 (red) buffer. Data are represented as mean of four separate runs with >500 events recorded per run. E
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Biomacromolecules PEG-L-eNPs exhibit less significant increases in uptake over time (Figure 8). At 1 and 4 h, FR+ KB cells show significantly
Figure 10. Confocal microscopy confirms uptake of Rho labeled FAPEG-L-eNPs (top) and PEG-L-eNPs (bottom) after 4 h of incubation. KB cells were costained with DAPI and Concanavalin-A to mark the nucleus and cell membrane, respectively. Rho-eNP aggregates are observed within the cells and are shown in red.
Figure 8. Uptake of rhodamine-labeled (A) PEG-L-eNPs and (B) FAPEG-L-eNPs at a concentration of 8.45 × 106 particles/mL in FR+ KB cells as quantified by flow cytometry.
more FA-PEG-L-eNP uptake than PEG-L-eNP uptake at all three concentrations (Figure 9). At 24 h, FR+ KB cells exhibit significantly enhanced uptake of FA-PEG-L-eNPs at the two lower NP concentrations. However, at the highest NP concentration, at 24 h, there is not a significant difference between uptake of targeted and nontargeted particles, which suggests saturation of folate receptors at this concentration within 24 h. In the FR− A549 cells, there is no significant difference between the uptake of targeted eNPs and nontargeted eNPs at all three time points, consistent with the lack of folate receptors on these cells. Similarly, uptake was assessed via confocal microscopy. KB cells were treated with Rho-labeled FA-PEG-L-eNPs and Rholabeled PEG-L-eNPs at a concentration of 1.69 × 107 particles/ mL, and imaged 4 h after treatment (Figure 10). Confocal images reveal that KB cells exhibit enhanced uptake of targeted NPs compared to nontargeted NPs, and z-stack images show that FA-PEG-L-eNPs are localized within the cells. In Vitro Cytotoxicity Analysis. Previously, nontargeted Pax-loaded eNPs demonstrated efficacy in multiple cell lines in vitro including breast, lung, ovarian, and mesothelioma cancers.48−54 To assess the toxicity of the lipid-coated eNPs in a folate-receptor expressing cell line, Pax-loaded PEG-LeNPs and FA-PEG-L-eNPs were incubated with KB cells and compared with Pax dissolved in cremophor/ethanol (Pax-C/E) at 4 and 24 h of incubation (Figure 11). Since Pax-PEG-L-eNP uptake varies minimally between 4 and 24 h (Figure 8), these particles exhibit similar potency at 4 and 24 h, with IC50 values of 26.87 and 23.11 ng/mL, respectively. The effect of Pax-FA-
Figure 11. In vitro cytotoxicity of eNP formulations. Folate receptor overexpressing KB cells were treated for 4 or 24 h with Pax-PEG-LeNPs, Pax-FA-PEG-L-eNPs, or Pax-C/E, and viability was assessed 72 h after treatment. Drug free NP formulations did not affect cell viability (data not shown). Data displayed as mean ± SD; n = 6−8.
PEG-L-eNPs at 4 h is similar to that of Pax-PEG-L-eNPs, with an IC50 of 20.42 ng/mL. Although the uptake of FA-targeted particles is significantly greater than that of nontargeted particles at 4 h (Figure 9), this difference does not translate into an effect on the potency of the NP formulation. However, at 24 h, Pax-FA-PEG-L-eNPs are more potent than Pax-PEG-LeNPs, with an IC50 of 12.87 ng/mL. Unloaded NP formulations do not affect cell viability (data not shown). At 4 h, the potency
Figure 9. Concentration-based assessment of PEG-L-eNP and FA-PEG-L-eNP uptake in (A) FR+ KB cells and (B) FR− A549 cells, as determined by flow cytometry. Data are displayed as mean ± SD (n = 4 for KB cells; n = 3 for A549 cells). ∗, p < 0.05 for each pair at discrete time points and concentrations. #, no significant difference. F
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CONCLUSION The lipid−polymer hybrid NP system presented here combines the pH-responsive swelling of the previously described eNPs with the improved colloidal stability and surface functionality of liposomes. Folic-acid conjugated PEG-L-eNPs demonstrate enhanced uptake in vitro in a folate receptor expressing cell line, and the addition of PEGylated lipids allows for significantly improved colloidal stability in serum relative to SDS-coated eNPs. When loaded with the anticancer agent, paclitaxel, the resulting hybrid polymer/lipid nanocarriers exhibit greater potency compared to nontargeted PEG-L-eNPs. Importantly, the swelling behavior of the eNP polymer core is maintained, which affords the unique opportunity to form intracellular hydrogel structures. Hybrid polymer−lipid NP systems offer opportunities for improved stability and enable active targeting via the lipid surface coating while maintaining the advantages of a polymeric core.
of the clinical formulation of Pax, Pax-C/E, is the lowest, with an IC50 of 111.2 ng/mL. At 24 h, Pax-C/E is the most potent, with an IC50 of 1.99 ng/mL. This result is dependent on the experimental conditions, as the amount of drug uptake (or efflux) over time will affect efficacy. Targeting ligands, including antibodies, peptides, and small molecules, are incorporated into NP coatings to improve cell uptake along with PEGylation to prevent nonspecific protein.19,68,69 Many receptors and surface proteins are overexpressed in cancer cells relative to healthy cells and serve well as targets for specific drug delivery.10,19,70 The folate receptor is one such target, which possesses a high affinity for folic acid.10,15,17,19,70−74 In agreement with the results reported here, previous studies report enhanced uptake and potency of folate-targeted drug-loaded entities, such as micellar paclitaxel, both in vitro as well as in vivo.15,75−78 It is also known that avidity varies as a function of cellular receptor density as well as NP ligand valency.79 Additionally, folate-targeted NPs and imaging probes are being explored in in vivo imaging applications as a result of their enhanced uptake and association with FR+ cancer cells, further documenting its beneficial use.80−82 However, since the endocytosis of folate targeted entities occurs through recycling centers characterized by pH ∼6.5, the extent of FA-PEG-L-eNP expansion will be reduced, as eNP swelling is pH-dependent.83 The new hybrid PEG-L-eNPs explored in this study are prepared using a minor modification of the miniemulsion polymerization technique and add to the collection of previously prepared and characterized hybrid lipid−polymer assemblies.1,9,13,45,59,84,85 Given the diversity of polymeric NPs successfully coated with lipids, the lipid-coating strategy is likely general and applicable to various NP systems. Several factors must be considered in the development of hybrid lipid−PEG NP systems with a nonresponsive or responsive core. For example, the formulation components (e.g., ratio of lecithin/ DSPE-PEG/DSPE-PEG-FA) are known to affect release kinetics, NP uptake, and colloidal stability; optimization studies should be undertaken after the initial development of a new system. With regards to the responsive core in our studies, the pH-responsive swelling of the polymeric core is not affected by the lipid−PEG coating, and the drug release kinetics are not affected by the pH conditions. These results suggest that Pax partitions primarily into the lipid layer of the PEG-L-eNPs and not the core. Therefore, it is important to consider the partitioning of the payload into the core versus the coating when designing future hybrid polymer−lipid NP systems, as this may critically affect the release kinetics. These results also highlight the potential opportunity, for those working in this area, to use the polymer core as well as the lipid shell for drug loading and thus engineer NP systems with multiple payloads or multimodal release kinetics. Finally, the selection of the targeting agent is an important consideration for cellular uptake, NP function, and drug release. For example, with folate targeting, in contrast to antibody mediated targeting or passive targeting, endocytosis occurs with a minimal reduction in pH to ∼6.5 as opposed to a pH of ∼5.83 While folate is one of the most extensively used ligands for targeted drug delivery, the incorporation of other targeting moieties, aptamers, or antibodies can be explored and has been described by others.1,13,59,85
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ASSOCIATED CONTENT
S Supporting Information *
Transmission electron micrograph of eNP. The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.biomac.5b00336.
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AUTHOR INFORMATION
Corresponding Authors
*E-mail: [email protected]. Phone: (617)-358-3429. Fax: (617)358-3186. *E-mail: [email protected]. Phone: (617)-353-7366. Fax: (617)353-5866. Author Contributions
M.S. and I.E. contributed equally to this work. 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 This work was supported in part by funding from National Science Foundation (DMR-1006601, DGE-1247312), Boston University’s Nanomedicine Program and Cross-Disciplinary Training in Nanotechnology for Cancer (NIH R25 CA153955), Brigham and Women’s Hospital, and Boston University.
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DOI: 10.1021/acs.biomac.5b00336 Biomacromolecules XXXX, XXX, XXX−XXX
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DOI: 10.1021/acs.biomac.5b00336 Biomacromolecules XXXX, XXX, XXX−XXX
