Article pubs.acs.org/Biomac
Injectable Microgel-Hydrogel Composites for Prolonged SmallMolecule Drug Delivery Daryl Sivakumaran, Danielle Maitland, and Todd Hoare* Department of Chemical Engineering, McMaster University, 1280 Main Street West, Hamilton, Ontario, Canada L8S 4L7 S Supporting Information *
ABSTRACT: The design and application of soft nanocomposite injectable hydrogels containing entrapped microgels for smallmolecule drug delivery is demonstrated. Copolymer microgels based on N-isopropylacrylamide and acrylic acid were synthesized that exhibited both ionic and hydrophobic affinity for binding to bupivacaine, a cationic local anesthetic used as a model drug. Microgels were subsequently immobilized within an in situ-gelling hydrogel network cross-linked via hydrazide-aldehyde chemistry to generate hydrogel−microgel soft nanocomposites. Drug release could be sustained for up to 60 days from these nanocomposite hydrogels, significantly longer than that achievable using the constituent hydrogel or microgels alone (99.8%), and bupivacaine hydrochloride (99%) were all purchased from Sigma Aldrich (Oakville, ON). Sodium dodecyl sulfate (SDS, electrophoresis grade) was obtained from Bioshop Canada (Burlington, ON). Dimethyl sulfoxide (DMSO, reagent grade) was purchased from Caledon Laboratory Chemicals (Georgetown, ON). 3T3 Mus musculus mouse cells, C2C12 mouse muscle myoblast cells, and RAW 264.7 macrophage mouse cells were acquired from ATCC: Cederlane Laboratories (Burlington, ON). Media contents included Dulbecco’s modified Eagle’s medium-high glucose (DMEM), fetal bovine serum (FBS), horse serum (HS), and penicillin streptomycin (PS), all of which were obtained from Invitrogen Canada (Burlington, ON). Recovery cell culture freezing and trypsin-EDTA were purchased from Invitrogen Canada (Burlington, ON). Thiazolyl blue tetrazolium bromide (MTT) was purchased from Sigma Aldrich (Oakville, ON). Microgel Synthesis. Acrylic acid-functionalized poly(NIPAM) microgels (AA-NIPAM) were synthesized via a mixed precipitationemulsion polymerization according to the formulations given in Table 1. Microgels were synthesized based on methods described by Hoare and Pelton.38 Polymerizations were performed in a three-necked flask with magnetic stir bar and attached condenser. Specified amounts of NIPAM monomer, acrylic acid, MBA, and SDS were dissolved in 150 mL of Milli-Q water. Polymerizations were conducted at a temperature of 70 °C under nitrogen purge at a mixing rate of 200 rpm. After 30 min of heating, APS was dissolved in 10 mL of Milli-Q water and injected into the mixture to initiate polymerization. Polymerization was allowed to continue for 12 h, followed by dialysis for purification. Dialysis was performed against Milli-Q water with membrane tubing from Spectrum Laboratories with a molecular weight cutoff of 12 000 4113
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cellulose membranes (Spectrum Laboratories) with a molecular weight cutoff of 100 kDa. Elution studies were conducted based on a microgel suspension volume of 0.67 mL that contained 5 mg/mL bupivacaine and 40 mg/mL AA-NIPAM microgels. Note that drug not taken up by the microgels was not specifically removed in these release studies such that the loading protocols used for both the microgels and microgel-hydrogel composites are matched. A 6 mL elution medium of PBS was used as the release medium, with the full volume replaced at each sampling time to ensure infinite sink conditions. Sampling was conducted every 30 min for the first 3 h and subsequently measured every hour for the 3 h following. Samples were collected until drug was no longer detected in the eluent using a DU 800 UV/visible spectrophotometer (Beckman Coulter) operating at a wavelength of 262 nm. Hydrogel and Microgel-Hydrogel Composite Drug Release Studies. Drug release studies for hydrogels and hydrogel−microgel composites were conducted in 12-well plates using cell culture inserts (2.5 cm diameter, 8 μm pore size) to contain the hydrogel. Prior to testing, the cell inserts were perforated 20 times using a 20 gauge needle to ensure facile flow of PBS in and out of the insert. A total of 2 mL of PBS was added to each well as the release medium, with six hydrogel or hydrogel−microgel nanocomposite samples assayed per composition investigated. Samples were incubated in an orbital shaker (VWR) at 37 °C at an oscillation rate of 100 rpm. Gels were switched into fresh media every 30 min for the first 2 h, then every hour for the next 5 h and intermittent measurements following over the course of the gel lifetime, with all drug concentrations assayed representing 1 mg/mL for CMC-B (reduced to 80−90% cell viability at 2 mg/mL), whereas other components retained high viability even at higher concentrations. Of particular note, despite some reports of cytotoxicity associated with aldehydecontaining polymers and the concurrent risk of protein crosslinking,48 the B (aldehyde-functionalized), polymers tested herein maintain high cell viability even at extremely high concentrations; indeed, cell viability of ∼50% (Dex-B) to ∼80% (CMC-B) is maintained even at a concentration of 10 mg/mL relative to both C2C12 and 3T3 cells (Supporting Information, Figure S6). Similar results were obtained for 3T3 cells and RAW 247.6 macrophages (Supporting Information, Figures S7 and S8). These results are consistent with prior studies involving PNIPAM-based microgels34 and hydrazide−aldehyde cross-linked hydrogels.15 Cell viability of 3T3 and C2C12 cells in the presence of a CMC-A44/Dex-B hydrogel and a composite CMC-A44/DexB/AA-6% hydrogel was also assessed using two test methods: (1) plating cells in the polystyrene well and placing a hydrogel disk on top of the cells and (2) extruding the hydrogel to fill the polystyrene well and plating cells on top of the hydrogel. These two geometries were selected to mitigate potential complications in terms of nutrient accessibility (hydrogel on top), cell-hydrogel contact (hydrogel on top), and cell adhesion (hydrogel on bottom) that may skew the results of any cell 4118
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viability study. No significant cytotoxicity was observed when cells were plated below or above a CMC-A44/Dex-B hydrogel, whereas the nanocomposite hydrogel induced only minimal cytotoxicity at extremely high mass concentrations relative to the total volume of media used (16 mg of hydrogel per 1 mL of media) (Supporting Information, Table S3). It should be noted that cell morphology when plated on the hydrogel relative to tissue culture polystyrene was slightly rounder in nature and not as well spread out, suggesting that some of the decrease in cell viability may be attributable to poor adhesion between the adhesion-dependent 3T3 and C2C12 cell lines and the hydrophilic hydrogel phases.
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with lower degrees of functionalization (and correspondingly larger deswelling transitions between the preparation conditions (25 °C) and the testing conditions (37 °C), Table S1 of the Supporting Information) exhibit higher burst releases of drug at early times. This is consistent with previous studies of drug release from bulk PNIPAM hydrogels, 51 further confirming that the entrapped microgel phase and bulk hydrogel phase are relatively independent. In addition, the observed swelling of the microgel as the loaded bupivacaine concentration decreases (Table S2 of the Supporting Information) would assist in facilitating a slow increase in the effective diffusion coefficient of drug inside the microgel phase over time, promoting a more constant drug release rate over an extended duration. Bupivacaine was the model drug chosen in this work both for its favorable physical properties (high ionic and hydrophobic partitioning affinity for AA-NIPAM microgels) as well as its clinical relevance, as long-term local anesthetic delivery is of significant interest for the management of postsurgical or chronic pain.52 However, the principles described herein should apply to any target drug, with appropriate design of the bulk or embedded phases to include affinity groups for whatever target drug is of interest. The ∼60-day period of release for bupivacaine observed for AA-20% nanocomposite hydrogels is significantly longer than release durations previously reported from hydrogel-based materials for bupivacaine53 or other smallmolecule drugs.54 Together with the facile injectability of this soft nanocomposite material, the long-term drug release achieved suggests significant potential for using these soft nanocomposite hydrogels clinically for long-term local drug delivery.
DISCUSSION
The results presented herein indicate that soft nanocomposite hydrogels prepared by physically entrapping microgels inside injectable bulk hydrogels offer broad control over both the total duration and instantaneous rate of drug delivery. Specifically, tuning the cross-link density, chemistry, and charge density of both the hydrogel and microgel phases allows for independent engineering of each hydrogel component to optimize the release profile of drug from the system. Drug release from soft nanocomposite hydrogels was found to be regulated by a combination of electrostatic binding (partitioning affinity) and diffusion effects. Electrostatic binding between cationic bupivacaine and anionic acrylic acid groups of AA-NIPAM microgels promotes higher proportional uptake of bupivacaine into the microgel phase than the bulk hydrogel phase; this is supported both by the bupivacaine binding dynamic light scattering data in Table S2 of the Supporting Information and the small bulk hydrogel-related burst release observed for the composite hydrogel in Figure 3. In this way, drug partitioning between the two soft phases regulates the initial distribution of bupivacaine between the bulk and microgel phases to minimize the magnitude of the initial burst release achieved. In addition, increased densities of anionic functional groups in either the bulk hydrogel phase or the microgel phase (i.e., increases in the affinity of the hydrogel or microgel phase for the drug) reduce the rate of drug release over the full lifetime of the nanocomposite hydrogel. Diffusion contributions to drug release are determined by the cross-link density of the hydrogel and how that cross-link density ultimately regulates hydrogel swelling or deswelling over the course of the release period. Composite hydrogels that swell more (or deswell less) release drug faster due to the higher average diffusion coefficient of bupivacaine within the more swollen network (Figure 6). Hydrophobic partitioning of bupivacaine into the moderately hydrophobic domains of the NIPAM residues may also contributes to drug partitioning in to the AA-NIPAM phase due to the similarity in log P values between bupivacaine and the N-isopropylacrylamide residue in the microgel.49 However, this effect was not directly assayed through the experimental data presented and is not required to rationalize the experimental results obtained, suggesting that this effect (if present) is small relative to the ionic partitioning effects in this system. Microgel swelling and deswelling effects, independent of the macroscopic composite hydrogel, may also contribute to the drug release profiles observed. Given that microgels are simply physically entrapped inside hydrogels in this work, they are free to shrink according to changes in their environment, largely unconstrained by the surrounding hydrogel phase.50 Microgels
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CONCLUSIONS Soft nanocomposite hydrogels based on injectable carbohydrate bulk hydrogels containing entrapped, poly(N-isopropylacrylamide)-based microgel nanophases facilitated the release of the small-molecule drug bupivacaine over the course of several weeks, significantly longer durations than could be achieved using the hydrogels or microgels alone. The release rate, duration of release, and magnitude of the drug burst can all be tuned by adjusting the affinity of each of the two soft phases for the target drug as well as the cross-link density of the bulk hydrogel phase, which regulates both drug diffusion and hydrogel degradation. Both the gel precursors and the nanocomposite hydrogels exhibit minimal cytotoxicity, even at high concentrations. Such materials, adapted for any target drug, may address clinical needs in terms of facilitating longterm local delivery of small-molecule drugs using an injectable, in situ-gellable hydrogel vehicle.
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ASSOCIATED CONTENT
S Supporting Information *
Potentiometric titration data for hydrazide-functionalized carboxymethyl cellulose and aldehyde modified carbohydrates, dynamic light scattering sizing data for microgels and druginduced microgel-drug deswelling, rheological data of the hydrogel and hydrogel−microgel composites, and cell toxicity of hydrogel and microgel components when exposed to 3T3 cells and RAW 264.7 macrophages are all provided. This material is free of charge via the Internet at http://pubs.acs. org/. 4119
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AUTHOR INFORMATION Corresponding Author *E-mail:
[email protected].
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ACKNOWLEDGMENTS
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REFERENCES
Thomas Oszustowicz is acknowledged for his assistance in collecting electrophoretic mobility and particle size data. Funding from the Natural Sciences and Engineering Research Council of Canada (NSERC) and the Ontario Ministry of Research and Innovation (Early Researcher Award program) is gratefully acknowledged.
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