(513g) Injectable Nanocomposite Hydrogels with Engineered “Smart” Properties

INTRODUCTION: The development of reliable, adaptable, and easily administered drug delivery systems to deliver drugs at the rate and location desired in the body has tremendous potential to improve clinical outcomes and quality of life. Successfully achieving controlled rate and location drug delivery depends entirely on precisely engineering the nature of the vehicle used to deliver the drug. Specifically, the chemical, physical, mechanical, and biological properties of the materials used to entrap, encapsulate, or bind a specific drug must be optimized to achieve a targeted drug release rate via diffusion and/or material degradation. Based on this requirement, hydrogels hold tremendous promise for addressing current drug delivery challenges. The crosslinking density of hydrogels can be modified to control the average pore size and thus the rate of drug diffusion from the hydrogel (1). Concurrently, the chemistry of the crosslinks can be tuned to control the degradation time of the hydrogel in vivo (2). Hydrogels typically exhibit excellent biocompatibility given their physicochemical and mechanical similarity to cell extracellular matrix (3). In addition, the incorporation of ?smart? materials such as poly(N-isopropylacrylamide) (PNIPAM) that can be reversibly triggered to change the porosity of hydrogels offers potential for developing drug release systems that are dynamically tunable as a function of temperature, pH, or some other physical stimulus (4). However, drug delivery from hydrogels is inherently limited in that the properties which contribute to the generally excellent biocompatibility of hydrogels are in many cases also detrimental to effective local drug delivery. For example, the highly hydrated microstructure of hydrogels imparts good biocompatibility but also facilitates very rapid drug release; furthermore, the bulk dimensions of hydrogels effectively prolong drug delivery relative to microparticle or nanoparticle-based formulations but render injection of hydrogels to the desired site of drug action to be difficult if not impossible. Thus, the development of injectable, self-gelling hydrogel formulations that contain local microdomains with controllable pore sizes, drug affinities, and biodegradation profiles has significant potential to address current drug delivery challenges. EXPERIMENTAL: Hydrazide-aldehyde chemistry was applied to synthesize two reactive pre-polymers that rapidly form covalent crosslinks via hydrazone bond formation upon mixing (5). Hydrazide-derivatized polymers (denoted as A-polymers) were synthesized by reacting carboxymethyl cellulose (molecular weight 400-600kDa) or PNIPAM-co-acrylic oligomers with molecular weights below the kidney cut-off limit (Mn = 19kDa, polydispersity = 1.6) with adipic acid dihydrazide using EDC/NHS chemistry. Aldehyde-derivatized polymers (denoted as B-polymers) were synthesized by oxidizing dextran or carboxymethyl cellulose using sodium periodate. When any A polymer was mixed with any B polymer via co-injection through a needle at concentrations of at least 2 wt%, a hydrogel was formed. Thermoresponsive microphases based on PNIPAM microgels were synthesized via a mixed precipitation-emulsion free radical polymerization mechanism in a dilute (~1 wt% monomer) aqueous solution. By co-injecting AA-NIPAM microgels together with one of the reactive polymers, physically-entrapped composite hydrogel-microgel networks were formed. The mechanical properties of the resulting hydrogels and hydrogel-microgel composites were analyzed using rotational viscometry. Bupivacaine, a cationic local anesthetic which can be loaded at high concentrations into AA-NIPAM microgels via both ionic and hydrophobic partitioning (6), was used as the model drug. Drug release from the composite hydrogels was analyzed using a transwell plate technique and quantified using UV/VIS spectrophotometry. Biocompatibility was assayed with the MTT assay, using fibroblasts and myoblasts as model cells. RESULTS: Hydrogels were formed within ~30s post-injection via mixing any A polymer with any B polymer using the double-barrelled syringe. Thus, hydrogels of varying compositions can readily be fabricated by co-extruding any mixture of A polymers with any mixture of B polymers. The water content (and thus mechanical strength) of the resulting hydrogels was directly related to the hydrophilicity of the individual polymer components of the hydrogel. Thus, by mixing various A and B polymer components, injectable, self-gelling hydrogels with a wide range of physical properties and biodegradation rates can easily be formed. Microgels were quantitatively incorporated into the self-gelling bulk hydrogel network at concentrations of up to 50wt%. Mechanical testing indicated that all hydrogels exhibited characteristic shear-thinning behavior at a strain rate of 5%. G' and G? data also indicated that both the hydrogels and microgel-impregnated hydrogels exhibited a predominantly elastic response to oscillatory strains over a wide range of oscillation frequencies (G'/G? > 8). Interestingly, even though microgels are incorporated only via physical entrapment inside the bulk hydrogel, microgel-hydrogel composites exhibited a slightly more elastic strain response than the hydrogels alone. We expect that the higher local chain density of the microgel phase relative to the bulk hydrogel accounts for this observation. Increasing the degree of hydrazide or aldehyde functionalization in the hydrogel pre-polymers (by increasing the EDC/NHS ratio used to functionalize the polymer and/or using an A polymer with a higher density of reactive carboxylic acid groups) results in lower water contents and more elastic hydrogel networks. Drug release from the microgel-hydrogel composite networks was observed to be significantly slower than drug release from the hydrogels or microgels alone. Bupivacaine release was sustained over a period of up to 60 days using microgel-hydrogel composites (see figure). In contrast, both the bulk hydrogel phase alone and the microgels alone release effectively all entrapped bupivacaine within 1-2 weeks, depending on the crosslink density and functional group content of each gel phase. Drug release rates scaled directly with the anionic functional group content of the microgel phase; the higher the degree of acrylic acid functionalization in the microgel, the higher the ionic binding between the cationic drug and the microgel and the slower the drug release. Indeed, release rates from the composite microgels appear to be primarily driven by ion exchange and/or ion partitioning between the microgel and drug as opposed to simple diffusion. As evidence of this, bulk hydrogels prepared with CMC B instead of Dextran B swelled significantly more over the duration of the drug release experiment but released drug at a statistically identical rate. Thus, the composition of the microgel phase and not the bulk hydrogel phase appears to predominantly control the drug release kinetics. The composite hydrogels, hydrogel pre-polymers, and microgels all showed no significant cytotoxicity to fibroblasts or myoblasts at concentrations up to 2mg/mL according to the MTT assay. Thus, these materials have potential utility as effective in vivo drug delivery vehicles. It should be noted that a 10-fold higher total concentration of microgel could be tolerated by all cell types tested if the microgels are entrapped inside a bulk hydrogel phase compared to microgel added directly to the cell media without hydrogel encapsulation. Thus, entrapment of microgels inside a hydrogel phase significantly increases the potential to safely use high concentrations of thermoresponsive microgels in vivo. CONCLUSIONS: Microstructured, in situ-gelling composite hydrogels based on hydrazide-aldehyde crosslinking chemistry have been designed, characterized, and tested. The bulk hydrogel phase can be prepared by mixing any number of aldehyde and hydrazide-functionalized polymers together to achieve a range of desired physical and biological properties. Drug release durations of up to 60 days can be achieved using hydrogel-microgel composites, significantly longer durations of release than can be achieved using hydrogels or microgels alone. As the microgels and polymers induce no significant cytotoxic effects, these injectable materials are excellent candidates as local drug delivery vehicles. ACKNOWLEDGEMENTS: Funding for this research was provided by McMaster University and the Natural Sciences and Engineering Research Council of Canada (NSERC). REFERENCES: (1) Kim, S. W.; Bae, Y. H.; Okano, T. Pharmaceutical Research 1992, 9, (3), 283-290; (2) Murthy, N.; Xu, M. C.; Schuck, S.; Kunisawa, J.; Shastri, N.; Frechet, J. M. J. Proceedings of the National Academy of Sciences of the United States of America 2003, 100, (9), 4995-5000; (3) Park, H.; Park, K. Pharmaceutical Research 1996, 13, (12), 1770-1776; (4) Hoare, T.; Pelton, R. Langmuir 2008, 24, (3), 1005-1012; (5) Bulpitt, P.; Aeschlimann, D. J. Biomed. Mat. 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