Developing a drug delivery system for treatment of vocal fold scarring
Vocal fold scarring is an affliction that results in the formation of a disorganized and stiff extracellular matrix (ECM) with abnormal ECM component densities & structures including a significant increase in collagen deposition. It is caused by improper healing post injury and results in profound changes in the biomechanical properties of the vocal folds impairing their ability to generate a normal mucosal wave during phonation. Finding an effective treatment for vocal fold scarring has been elusive. Currently, treatments seek temporary solutions that correct glottal incompetence and reduce stiffness caused by the scar through the augmentation of the vocal folds using various injectable materials, such as hyaluronic acid or calcium hydroxylapatite. These solutions do not actually treat the actual scar, but only relieve its effects over the short term seeing as current injectable materials readily degrade. To solve this problem researchers have been testing the impact various therapeutics have on helping prevent or regress vocal fold scarring. Unfortunately, realizing the full potential of such therapeutics has not been fully achieved due to many barriers. The largest being that, even if anti-scarring agents are directly injected into a vocal fold scar, these therapeutics are readily cleared or metabolized by the vocal fold tissue over a short period of time. This necessitates the need to re-inject the therapeutic multiple times in order to maintain its efficacy. Though this is possible, issues of patient safety and tolerance quickly become significant problems one must address if one desires to pursue such a rigorous treatment method. The overall goal of my research was to address this issue of effectively delivering therapeutics to an area of vocal fold scarring to facilitate healthy unscarred tissue regeneration. To accomplish this goal my early work focused on the development of in vitro unscarred and scarred culture systems in which to eventually test our newly developed delivery systems. Initial work with these systems focused on characterizing them in order to better understand the role RGD integrin binding site concentration and cell phenotype had on directing extracellular matrix gene expression in vocal fold fibroblasts. What we learned was that indeed RGD integrin binding site concentration does affect both depending on the circumstances. With changes in fibroblast phenotype from unscarred to scarred resulting in ECM gene up-regulation for all genes tested, except for HAS2 and decorin. Meanwhile, changes in RGD concentration only increased elastin and collagen type 3 alpha 1 expression as RGD concentration was increased in scarred vocal fold fibroblasts only. HAS2 was also down-regulated in scarred fibroblasts at the highest RGD concentration. However the other genes tested were unimpacted by RGD concentration changes in scarred or unscarred fibroblasts. This knowledge is critical in developing bioactive materials that, when implanted into sites of tissue damage and scarring, direct cells to regenerate healthy tissues with normal ECM ratios and morphologies. Next, we moved on to developing our drug delivery system initially utilizing a nanoparticle based system. Our nanoparticle system was composed of a poly-lactic-co-glycolic acid (PLGA) core encapsulated within a shell of poly(N-isopropyacrylamide (pNIPAM) and had unique properties that were designed to help better facilitate more effective delivery of therapeutics including active targeting and the ability to respond to dynamic environmental stimuli. However, concerns arose that the small size of the nanoparticles would not allow for an effective amount of drug to be loaded into them and that they would more readily degrade and release loaded therapeutics due to their high surface to volume ratio. We, therefore, transitioned to using larger hydrogel templated PLGA microparticles instead. We also moved on to testing our ability to actually deliver an effective anti-fibrotic and anti-inflammatory agent to scarred vocal fold fibroblasts over the long term. To do this, the PLGA microparticles were loaded with the corticosteroid dexamethasone. Real-time PCR showed that only dexamethasone loaded microparticles proved effective at maintaining down regulation in the expression of COL3A1 and COL1A2 over the long term. While an ELISA showed that dexamethasone only decreased the deposition of the pro-inflammatory cytokine interleukin-6 over the long term, but had no impact on the other three pro-inflammatory cytokines tested. Future work will look at trying to continue to improve the drug delivery platform for use in vocal fold scarring by combining the pNIPAM-co-acrylic acid shell, from our core + shell nanoparticle work, with the PLGA microparticle delivery system. Such a combination would meld the advantages of the larger microparticle size and its ability to more readily deliver an effective dose of therapeutic with the many beneficial attributes, like targeting and responsiveness to environmental stimuli, of the pNIPAM-co-acrylic acid shell. Furthermore, with the addition of a shell, another level of control over therapeutic release would be available.
Sivasankar, Purdue University.
Speech therapy|Biomedical engineering
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