Thermo-Responsive Core-Shell Composite Nanoparticles Synthesized via One-Step Pickering Emulsion Polymerization for Controlled Drug Delivery

Purpose: The focus of this work is to develop a unique drug delivery vehicle which can be taken up by cancer cells and can release the loaded drug. Methods: Core-shell composite nanoparticles have been prepared by one-step Pickering emulsion polymerization with a nonionic initiator, using silica as the sole stabilizer. More importantly, the Pickering emulsion polymerization is applied to synthesize polystyrene/poly(N-isopropylacrylamide) (PNIPAAm)-silica core-shell nanoparticles with N-isopropylacrylamide incorporated into the core as a co-monomer. Results: The composite nanoparticles are temperature sensitive and can be taken up by human prostate cancer (PC3 and PC3-PSMA) cells. An anticancer agent 17-(Allylamino)-17-demethoxygeldanamycin (17-AAG) has been loaded into the polymeric cores during formation of the nanoparticles and drug release has been successfully observed at elevated temperatures. The ability of the various nanoparticles for inducing death in human prostate cancer cells has been evaluated. Conclusion: The work has demonstrated the temperature sensitivity, controlled drug release properties of the synthesized core-shell nanoparticles, and their effectiveness for inducing death of human prostate cancer cells. Citation: Sanyal S, Huang H, Rege K, Dai LL (2011) Thermo-Responsive Core-Shell Composite Nanoparticles Synthesized via One-Step Pickering Emulsion Polymerization for Controlled Drug Delivery. J Nanomedic Nanotechnol 2:126. doi:10.4172/2157-7439.1000126


Introduction
Nanoparticles are receiving increasing attention since they have properties that are significantly different than those of bulk materials [1,2]. The high surface-to-volume ratio of nanoparticles makes them a suitable carrier or delivery system for drugs, proteins or genes which can be either delivered locally or targeted specifically [3]. Organicinorganic composites are vital in biological and medical applications such as in artificial bones, dental fillings, and drug delivery [4]. Among them, core-shell composite nanoparticles are a unique class of materials which are attractive for their potential applications as delivery vehicles (for drugs, dyes, cosmetics, ink etc.) [5,6]. Core-shell composite particles can be synthesized by methods of post-surface-reaction [7,8], electrostatic deposition [9], and layer-by-layer self-assembly [10][11][12].
Here we employ the concept of Pickering emulsions to synthesize core-shell composite particles. Pickering emulsion polymerization is superior in several aspects: (1) no sophisticated instrumentation is needed; (2) a commercialized nanoparticle powder or solution can be used without further treatment; (3) the synthesis can be completed in one-step; and (4) the produced particle dispersion is surfactant-free which makes it easier to purify and an excellent material for biological applications. Despite these advantages, efforts made to explore and utilize the Pickering approach have been scarce, although some related synthesis methods have been documented including miniemulsion polymerization [13,14], dispersion polymerization [15,16], inverse suspension polymerization [17,18], and inverse emulsion polymerization [19] stabilized by fine solid particles. It is worthwhile to note that the composite nanoparticle structure in this study is opposite to the often reported core-shell structure in which inorganic particles serve as the core and polymer serves as the shell [10,[20][21][22][23][24]. Here the polymer serves as the core and the inorganic particles serve as the shell. Such materials provide a new class of supramolecular building blocks and can "exhibit unusual, possibly unique, properties which cannot be obtained simply by co-mixing polymer and inorganic particles" [25].
"Smart" materials that respond to environmental changes, such as temperature or pH, are attractive means for designing "intelligent" drug carrier systems. In this study, N-isopropylacrylamide (NIPAAm) is incorporated as a co-monomer in order to impart temperature sensitivity to the core-shell nanoparticles. Poly (N-isopropylacrylamide) (PNIPAAm) is a well-understood temperature sensitive gel, which undergoes volume shrinkage at a transition temperature of approximately 32°C in pure water [26]. Below this temperature which is referred to as the lower critical solution temperature (LCST), the polymer chain is hydrophilic because the hydrogen bonding between the hydrophilic groups and water molecules dominates; above the LCST, the polymer chain becomes hydrophobic due to the weakened hydrogen bonding at elevated temperatures and the interactions among hydrophobic groups [27]. To synthesize PNIPAAm chains in solutions, techniques such as free radical initiation in organic solutions, redox initiation in aqueous medium, ionic initiators, and radiation of aqueous medium are used [28]. One advantage of using PNIPAAm is that it can be engineered to possess transition temperatures slightly above physiological temperatures (e.g., at 38-39°C) [29], that can be exploited for selective and controlled delivery at the tumor site using external stimuli (e.g. laser irradiation) which is suitable for triggered in vivo drug delivery. Recently a "nanopump" system was reported using block copolymer poly(L-lactide-star block-N-isopropylacrylamide) (PLLA-sb-PNIPAAm) for controlled drug release [30]. The PNIPAAm block leads to the shrinkage of the complex micelles and pumps the drug out when the temperature is above the LCST. However, the nanoparticles constructed out of the block polymer have a tendency to aggregate together if the temperature is above the LCST [31]. This was likely due to the loss of amphiphilic property of the block copolymer above the LCST as PNIPAAm blocks transformed into their hydrophobic character. In this work, we report a different synthetic pathway and use Pickering emulsion polymerization which provides adequate stability as evidenced by zeta potential measurements. Cancer drug candidate 17-(Allylamino)-17-demethoxygeldanamycin (17-AAG), an ansamycin antibiotic, which binds and inhibits Heat Shock Protein 90 (HSP 90), is incorporated into the core during synthesis and used as a hydrophobic drug in the release experiments. Figure 1 illustrates a synthesized core-shell "smart" nanoparticle and its release of drugs upon temperature change. Cellular uptake and cytotoxicity were studied in prostate cancer cells incubated at 37°C, with composite nanoparticles encapsulating 17-AAG and their controls.

Composite nanoparticle synthesis
Water, IPA-ST, NIPAAm and styrene were agitated mechanically with an IKA Ultra Turrax T25 homogenizer at 10,800 rpm for 2 min in an ice bath to emulsify. Then, the emulsion was degassed with nitrogen and kept in nitrogen atmosphere under magnetic stirring. When the temperature was raised to 70°C, the initiator aqueous solution was added to start the polymerization which lasted for 5 hours. The dye BODIPY493/503, when used for cellular uptake experiments, was added prior to adding the initiator. A typical formulation of the thermoresponsive nanoparticle includes 0.66 g NIPAAm, 3.76 g styrene, 32 mL water, 4.1 g IPA-ST silica nanoparticle dispersion, 0.037 g initiator VA-086, and 1 µg BODIPY 493/503. For the release experiments, 17AAG was added to the emulsion before the polymerization took place, prior to adding the initiator at a concentration of 600 g/ml. The drug, 17AAG, remained stable, as determined from comparing its Fourier transform infrared spectroscopy (FTIR) spectra before and after heating for 5 hours at 70°C. Before characterization, release and uptake experiments, the synthesized nanoparticles were washed twice by centrifuging-redispersing cycles using an Eppendorf 5810R centrifuge.

Characterization methods
Particle size distribution and zeta potential of composite particles were measured using a Brookhaven 90Plus Particle Size Analyzer with the dynamic light scattering (DLS) technique. The washed composite particles were further dispersed to appropriate concentrations with water before measurements. Specimens for electron microscope were prepared by placing a droplet of polymer samples onto mica substrates and dried in air. The specimens were then sputter coated with gold and viewed by SEM-XL30 (FEI). Fourier transform infrared spectra (FTIR) were scanned over the range of 400-4000 cm -1 with potassium bromide pellet on a Bruker IFS 66V/S FTIR spectrometer. To remove the silica shell, hydrofluoric acid (HF) etching procedure described by Han and coworkers [33] was adopted.

Cellular uptake experiments
Composite core-shell nanoparticles were loaded with BODIPY 493/503 dye in the core using methods described above. Dye-loaded nanoparticles were employed in cellular uptake experiments, in concert with confocal microscopy. PC3 and PC3-PSMA prostate cancer cells were seeded in a 24-well plate at a density of 50,000 per well and incubated overnight at 37°C and 5% CO 2 . Dye-loaded nanoparticles were added to cancer cells in the absence of serum for 0.5, 1.0, 1.5, and 5 h following which, the nanoparticle-containing medium was removed and cells were washed five times using 1X phosphate buffered saline (PBS). Cells were stained with the nuclear stain Hoechst 33258 (MP Biomedicals) and then mounted for analysis using a laser scanning Nikon C2 confocal microscope (Nikon Instruments Inc., Melville, NY). Images were acquired and stacked using NIS-Elements Microscope Imaging Software (Nikon Corporation) at 60× water objective with a z-step of 0.4 μm slice and with PMT scanners at 1024 × 1024 pixels.

Drug loading and release experiments
Composite core-shell nanoparticles loaded with 17AAG were used for the drug release experiments. The drug release study was done by centrifugation, after the nanoparticles were centrifuged and the supernatant was collected and tested for drug release. The amount of drug released was determined by UV analysis. Absorbance (at a wavelength of 332 nm) of a number of standard solutions of the reference substance at concentrations encompassing the sample concentrations was measured using a Biotek Microplate Reader and the calibration graph was constructed.

Cytotoxicity experiments
The cytotoxicity of nanoparticles (NPs) was determined using 3-(4, 5-dimethylthiazol-2-yl)-2, 5-diphenyl tetrazolium bromide (MTT) assay. Prostate cancer cells were cultured in medium (RPMI 1640 plus 10 % fetal bovine serum and 1% penicillin/streptomycin) under 5% CO 2 at 37°C. Cells were seeded in a 96-well plate at a cell density of 5,000 cells per well and allowed to attach overnight prior to NP treatment (0.01-1 μg/ml). At 72 h after treatment with different doses of NPs, 10 μl of MTT reagent was added to each well and incubated under 5% CO 2 at 37°C for 4 h, followed by addition of detergent. Absorbance at 570 nm was determined with microplate reader (Biotek Synergy 2). Studies were performed with at least 4 individual samples and repeated 3 times. Any interference from NPs alone (i.e. without cells) was normalized for all samples. Two-tailed Student's t-test was employed for statistical analysis.

Composite nanoparticle synthesis
Different ratios of styrene/NIPAAm were used in the formulation of the polystyrene/PNIPAAm-silica core-shell nanoparticles. It was found that when the concentration of PNIPAAm is high, the volume change of the nanoparticles is more significantly greater with change in temperature. Here we present the studies using the percentage of 15% NIPAAm. Pickering emulsion polymerization was performed using VA-086 as the initiator. In order to verify the sole stabilizing effect of silica nanoparticles, emulsifier-free emulsion polymerization using VA-086 as the initiator in the absence of silica nanoparticles was also conducted. No particle formation was observed in the product, as evidenced by SEM images. These experiments suggest that the initiator VA-086 has little effect on stabilizing in emulsion polymerization and therefore silica nanoparticles are the only source of stabilizer when present [34,35]. Figure 2(a) is a representative SEM image of the composite particles sampled at 5-hour reaction time which shows that the particles tend to be spherical. The roughness of the composite particle surfaces suggests that the composite particles are covered by silica nanoparticles; this is contrasted by the smooth surface of the hydrofluoric acid (HF)-treated particles in Figure 2 (b). HF dissolves the silica layer and leaves behind the smooth polymer surface. It is also evidenced by the blue line in the FTIR spectrum in Figure 2(c) which shows that the composite particles have a characteristic strong peak at 1104 cm -1 , corresponding to the asymmetrical vibration of the Si-O-Si bond. Such a peak is absent in the red line in Figure  2(c) which represents the HF-treated composite particles. FTIR is a strong analytical tool which gives information about specific chemical bonds simply by interpreting the infrared absorption spectrum; here it is used to identify the presence of silica. The core-shell structure of the nanoparticles has also been confirmed by transmission electron microscope (TEM) images of polystyrene-silica nanoparticle crosssections [34]. The measurement of zeta potential allows predictions about the storage stability of a colloidal dispersion. Particle aggregation is less likely to occur for charged particles (high zeta potential) due to electric repulsion. The mean zeta potential for the drug loaded polystyrene/PNIPAAm-silica nanoparticles was -48.1 mV. Therefore, this system has a relative good stability and dispersion quality.

Thermoresponsiveness of the composite particles
The synthesized polystyrene/PNIPAAm-silica core-shell nanoparticles are responsive to thermal stimuli. Figure 3(a) shows the dependence of average diameter of the composite particles on temperature. The average particle size at 25°C is approximately 92 nm. The size decreases sharply as the temperature reaches 32°C, around the LCST for homopolymer PNIPAAm and size change is nearly reversible upon cooling. Control experiments of polystyrene-silica nanoparticles did not show a size transition over a temperature range of 25-45°C (data not shown). The transition temperature is not shifted by the silica nanoparticle encapsulation. This is consistent with the recently reported composite microspheres with a PNIPAAm core and a silica shell which also show a volume transition starting at 32°C [17]. It is likely due to the fact that silica particles are physically adsorbed on the surfaces of PNIPAAm microspheres thus no chemical bond formation with silica occurs which might change the transition temperature. Moreover, the copolymerization with styrene has no significant effect on the transition temperature. One hypothesis is the relative phase separation of PNIPAAm and polystyrene within the core. Duracher et al. [36] studied PNIPAAm-polystyrene particles and suggest a PNIPAAm-rich shell and a polystyrene-rich core structure. Such phase separation may also occur in the core of the composite particles here although detailed morphology is unknown.

Drug release, cell-uptake and cytotoxicity
The drug candidate used in the release experiment is 17-AAG, an ansamycin antibiotic which binds to Hsp90 (Heat Shock Protein 90). Hsp90 plays a key role in regulating the physiology of cells exposed to environmental stress and in maintaining the malignant phenotype of tumor cells. Hsp90 client proteins are important in the regulation of the cell cycle, cell growth, cell survival, apoptosis, and oncogenesis. 17-AAG binds with a high affinity into the domain of adenosine triphosphate (ATP) binding pocket in Hsp90 and induces the degradation of proteins that require it for conformational maturation. Heat Shock Protein 90 (Hsp90) is of significant interest in cancer therapy because it helps in cell survival and tumor cell proliferation. The main obstacle to the delivery of 17-AAG is its poor water solubility  There is no release at 25°C which is below the transition temperature of the composite particle. The cumulative fractional release is shown as an inset where the maximum amount of drug release value is taken at 40°C for 10 hours; diffusion model parameters have been calculated from these data. and it requires complicated formulations with Cremophor EL (CrEL), DMSO, or EtOH before parenteral administration [37]. This is undesirable as CrEL is known to induce hypersensitivity reactions and anaphylaxis, and requires patient pretreatment with antihistamines and steroids before administration [38]. A safer administration of 17-AAG can be made by a surfactant free delivery system rather than using harmful surfactants to solubilize the drug. . Therefore, it requires a viable drug carrier for its time-controlled release. The release study of 17-AAG from the nanoparticles was first performed in deionized water. The composite nanoparticles encapsulated with 17-AAG to be tested were divided into two parts. The first portion was kept at room temperature and the second was in a temperature-controlled water bath at 40°C during the duration of the experiment. At regular time intervals, the samples were taken out and then centrifuged at 10,000 rpm for 3 min in a Beckman Coulter tabletop centrifuge and 100 µL of the supernatant was drawn out from release system for analysis. We have also performed DLS experiments on the supernatant to verify the absence of nanoparticles (data not shown; readings of 0 nm).
Cumulative drug release measurements were performed. Figure  3(b) depicts the actual release at 25°C and 40°C and cumulative fractional release (40°C) kinetics of 17-AAG from the drug-loaded nanoparticles. No significant release of the drug was observed at room temperature (25°C). However, at a higher temperature of 40°C, the drug release from the nanoparticles reached a maximum after 7 h. The cumulative fractional drug release (inset of Figure 3 (t) = kt n where M (t) is the amount of drug M (∞) released at time t, M(∞) is the total amount of drug released at a time ∞ which is taken to be the saturation time when no further amount of drug is released, k is a constant related to the physical properties of the system, and the index, n, is the diffusional component that depends on the release mechanism. When n<0.5, the solute is released by Fickian diffusion; when 0.5<n<1.0, the solute is released by non-Fickian diffusion and when n=1, there is zero order release [39]. The calculated n value is 0.73 which indicates the non-Fickian diffusion. The mathematical model indicates that the drug diffusion behavior is non-Fickian and the rate of drug release is due to the combined effect of drug diffusion and polymer response due to increase in temperature.
It was hypothesized that polystyrene/PNIPAAm-silica core-shell nanoparticles are sufficiently small to be taken up by cancer cells. Cell uptake experiments were performed using human prostate cancer cells (50,000/well in 24 well plates). The uptake of the composite nanoparticles by PC3 and PC3-PSMA cells following incubation for 0.5, 1, 1.5 (Supplementary Figure 1) and 5 h (Figure 4) was visualized using confocal microscopy. In Figure 4, at a nanoparticle dosage of 0 (control), 0.025 and 0.05 µg/ml, the human prostate cancer cells take up dye-loaded composite nanoparticles and traffic them throughout the cytoplasm while remaining viable after incubation for 5 h. Visible green fluorescence in Figure 4 suggests that composite nanoparticles are internalized in both the PC3 and PC3-PSMA cells. Initial internalization of nanoparticles in both PC3 and PC3-PSMA cells was observed after 1 h of incubation (Supplementary figure). A control was kept under observation at the different times with no nanoparticles. Uptake was observed at lower concentration of 0.025 µg/ml and at a higher concentration of 0.05 µg/ml. The higher concentration case shows presence of some debris which may be due to some cell death as indicated by cytotoxicity experiments. Nanoparticles could be transported into the cell by either specific (receptor-targeted) or nonspecific cellular uptake mechanisms depending on the surface properties. Since the nanoparticles are not conjugated with any antibody, the uptake behavior here is nonspecific. Nonspecific uptake of nanoparticles has recently attracted much attention in the development of new strategies for designing efficient nano-carriers, though specific uptake is a more developed strategy as more control is possible and effects on cell functions are easier to predict. Our findings suggest that the thermally sensitive composite nanoparticles can be taken up by prostate cancer cells, which opens new opportunities in controlled drug delivery. Further research in our laboratory will involve conjugation of targeting biomolecules on the surface of these nanoparticles in order to facilitate selective delivery to prostate cancer cells. Eventually, this strategy can result in lower toxicities towards non-malignant cells and other organs in vivo.
Drug-loaded nanoparticles are minimally toxic at lower doses (0.01 µg/ ml), but induce death in 35 -90% of the PC3 population at doses from 0.03-0.3 µg/ml. This is statistically significant from other nanoparticle formulations tested under similar conditions (Figure 5a), indicating that uptake and drug release of the 17-AAG containing nanoparticles resulted in death of PC3 cells. The nanoparticles were not as efficacious in inducing death in PC3-PSMA cells compared to PC3 cells (Figure 5b). While 17-AAG containing nanoparticles demonstrated higher average losses in cell viability in PC3-PSMA cells compared to other nanoparticle formulations, these differences were not statistically significant except at the highest concentration (0.3 µg/ml). This can be explained in part by the susceptibility of this cell line to nanoparticles without the drug; while 'bare' nanoparticles do not result in loss of viability of PC3 cells, they induce death in a large population of PC3-PSMA cells in a dosedependent fashion. In addition, previous results in our laboratory have demonstrated differential therapeutic efficacy in these cell lines due to differential intracellular trafficking and localization of nanoparticles [40,41]. These differences in closely related cell lines underscore the role of the cancer cell phenotype in determining efficacy of delivered nanoparticle therapeutics and are currently under investigation in our laboratories. In addition, it is hypothesized that conjugation of prostate cancer cell specific biomolecules (e.g. antibodies to the Prostate-Specific Membrane Antigen [42]) will help receptor-mediated uptake of these nanoparticles leading to increased efficacies.

Conclusions
The focus of this work was to employ a unique drug delivery vehicle which can be taken up by cancer cells and can release the loaded drug. Polystyrene/PNIPAAm-silica core-shell nanoparticles were successfully synthesized with N-isopropylacrylamide incorporated into the core of the nanocomposite as a co-monomer. The work has demonstrated the temperature sensitivity, controlled drug release properties of the synthesized core-shell nanoparticles, and their effectiveness for inducing death of human prostate cancer cells. The collapse of the PNIPAAm resulted in the shrinkage of composite particles at temperatures above the LCST which makes them promising as a unique drug delivery vehicle. Future work will focus on further understanding the delivery mechanism, investigation of targeted nanoparticles, and modification of the transition temperature of the composite particles to just above physiological temperature (i.e. 37°C) for in vivo drug delivery for destruction of the human prostate cancer tumors.