Dexamethasone (DXM)-Coated Poly(lactic-co-glycolic acid) (PLGA) Microneedles as an Improved Drug Delivery System for Intracochlear Biodegradable Devices

Uniform drug delivery techniques are challenging to develop for the inner ear due to the complexity of the cochlear anatomy. A promising solution is the use of biodegradable polymers because the continuous release of therapeutics without introducing toxic compounds is desirable. Using a microneedle approach lends the polymeric microneedle the capability to be placed inside of the scala tympani, releasing drugs overtime. Poly(lactic-co-glycolic acid) (PLGA) microneedles is prepared by dissolving dimethyl sulfoxide with either a) Rhodamine B, to study the drug release profile in vitro, b) FM1-43, to study the drug release profile in vivo, or c) dexamethasone (DXM), to protect hair cell (HC) loss in vivo. The Rhodamine B studies show that the dye begins release from the microneedles within 30 min. The ototoxicity assessment of the DXM-coated microneedles in vitro shows a significant reduction of HC losses when compared to control microneedles in an ototoxic environment. In vivo data show reduced hearing threshold for animals treated with DXM-infused microneedles, providing a proof of concept of the methodology developed. Drug-infused polymeric microneedles provide a promising method to deliver DXM to the inner ear over controlled periods of time protecting hair cells, thus minimizing hearing loss (HL).


Introduction
The design and development of reliable, biocompatible, and effective drug delivery systems for both systemic and localized applications is paramount given the emergence of novel biopharmaceuticals, biologics, and other therapeutic agents that require tailored administration characteristics different from those afforded by traditional systems. One of the FDA approved molecular mechanism of HC death in the organ of Corti (OC) is not yet fully elucidated. However, the cascade of events to create the electrical signal is known to be terminated, resulting in HL. [9] To address these issues, studies have repeatedly shown that dexamethasone (DXM), a corticosteroid, combats HC death by altering the expression levels of apoptosis related genes. [10][11][12][13][14][15] While systemic forms of drug delivery of DXM are available, undesirable side effects are common and the effective dose may not reach the ear when given systemically. [16] Other currently available treatment methodologies and strategies, apart from new drug development, include intratympanic (IT) injections of drugs, cochlear implants, stem cell therapies, manipulation of gene expression, viral vectors, thermosensitive gels, nanoparticles, and placement of drug delivery materials on the surface of the round window membrane (RWM), amongst others. [17][18][19][20] However, the development of these techniques for the delivery of DXM into the inner ear has been hindered by the complex cochlear anatomy, which limits molecular transportation of drugs. [17,21] In addition to the inherent complications caused by the cochlear anatomy, the short-lived time period that the perilymph solution experiences in the cochlea creates additional convolutions. Specifically, the blood labyrinth barrier separates the inner ear from systemic circulation with tight junctions, protecting the integrity of inner ear in a similar fashion to the blood brain barrier, and functions as a biochemical barrier by employing efflux pump systems to remove foreign agents from inner ear circulation, including DXM. [22] To overcome these challenges, a novel microfabricated drug delivery system was designed and developed that was implanted into the ear to deliver a uniform dose of the drug of choice over an extended period in a cost-effective manner. PLGA, a well-known biodegradable polymer, and DXM were employed as the materials for the fabrication of the microneedles. The polymeric microneedle implant has enough mechanical strength to pierce the RWM and be placed inside the cochlea, a procedure of which is well documented in the literature using cochlear implants. [23] Placing the microneedle directly inside of the scala tympani allows for the appropriate amount of drug to be released overtime directly to the affected area. Additionally, the extended-release profile of the polymer allows necessary doses of drug to remain in the perilymph solution despite perilymph clearance. Herein described is the method of microneedle development including fabrication of a customized mold. Furthermore, in vitro testing is described using OC explants exposed to said microneedles as well as different in vivo experimentation designed to provide a proof of concept as well as to study the preliminary intracochlear drug release profile through the assessment of HC viability and functionality. The developed technology not only addresses the hindrances associated with the convoluted anatomy and dynamic environment of the cochlea, but also proposes a new methodology that is applicable for many other intricate areas of the body that require targeted, localized, and uniform drug delivery.

Microneedle design for In Vitro experiments
The initial microneedle design for the in vitro experiments was on a larger scale for preliminary investigation. Different com- positions of glycolic acid and lactic acid were investigated as the crystallinity, and therefore the mechanical strength of the PLGA, varies depending on the monomer ratios, from amorphous to fully crystalline. Polylactic acid contains methyl side chain groups that increase the disorder of the polymer chain depending on the ratio of lactic to glycolic acid. Additionally, by altering the ratio of lactic acid to glycolic acid, the degradation time changes due to the additional methyl group on the lactic acid affecting the hydrolysis of the polymer, making it less hydrophilic and slower to degrade. [24] Thus, the 50:50 PLGA copolymer was selected to manufacture the microneedles as this ratio of lactic acid to glycolic acid degrades the quickest, which was desirable for the encapsulated drug to reach and protect the HC of the cochlea accordingly. To fabricate the microneedles, a positive mold was 3D printed using a custom designed mold via CAD software. A negative mold was cast from the positive 3D printed mold using polydimethylsiloxane (PDMS) due to its low cost, high flexibility, and ease of use ( Figure 1D). [25,26] The microneedles were then prepared by pipetting a homogenous solution of PLGA and DMSO into the mold and the solvent was allowed to evaporate overnight. The microneedle design incorporated a circular component on each end of the structure to allow for easier handling during the development process, which is subsequently removed to produce five microneedles per casting structure.

Biocompatibility Studies
Biocompatibility of the PLGA copolymer was investigated via analysis of 150 μm segments from the basal, middle, and apical turns of the OC's basilar membrane as these areas are important for low-, middle-, and high-frequency hearing, respectively (Figure 2). Studies were then performed to investigate potential toxicity of DMSO, ethyl acetate (EtOAc), and acetone as the solvents for the dissolving of PLGA for physical manipulation. Microneedles were prepared using these solvents and the experiment was carried out as described in the Experimental Section. The percent hair cell loss was calculated for 150 μm segments of the basal, middle, and apical turns of the OCs exposed to DMSO and acetone, both of which showed minimal HC loss when compared with EtOAc (Figure 3). Since the preliminary in vitro studies using EtOAc showed hair cell loss, only DMSO and acetone were investigated with additional OC explants (n = 3). DMSO creates a homogenous solution when dissolved with other drugs relevant for future studies including other experimental drugs, while acetone does not. Thus, DMSO was selected as the solvent of choice for the microneedle preparation for this application.

Drug Release Profile
After demonstrating that the material and solvent of the microneedles are safe and nontoxic, the microneedles were then cast using a fluorescent compound commonly used to charac-terize drug release profiles, Rhodamine B, in place of DXM to study the hypothetical drug migration in vitro in an artificial perilymph solution. [27,28] While DXM and Rhodamine B are inherently different compounds, this study was performed as an initial proof of concept to demonstrate that PLGA can encapsulate a foreign compound uniformly and subsequently release said molecule upon hydrolysis in an aqueous environment. The release profile demonstrated that the PLGA copolymer begins release of Rhodamine B within 30 min of contact with the artificial perilymph solution. The initial release of the drug was rapid and began to slowly trend toward a more uniform release overtime (Figure 4). To reduce the lag time from microneedle placement to initial drug release, the subsequent set of microneedles for in vitro experiments included an additional experimental group with a coating of pure DXM on the surface of the microneedle to deliver an immediate bolus of drug upon insertion. This coating served as a bridge between microneedle insertion and DXM release.
Following the demonstration of the immediate release of Rhodamine B from the microneedles for preliminary investigation, the long-term release profile of the drug-infused microneedles was assessed via liquid chromatography with tandem mass spectrometry (LCMS/MS). A drug-infused microneedle was placed in an artificial perilymph solution at 37°C and the concentration of drug in solution was measured every week for 1 month (n = 4) ( Figure 5). The concentration of drug released after 1 month was Figure 3. % HC Loss of OC explants exposed to microneedles prepared using different solvents demonstrating the biocompatibility of DMSO and acetone and the ototoxicity of EtOAc. Green: number of HCs in the basal; orange: middle; red: apical turns of the organ of Corti.  approximately one half of the total drug encapsulated in the microneedle, demonstrating the continued release of the drug from the microneedle over an extended period.

Safety and Otoprotective Properties
To investigate the safety and otoprotective properties of the microneedles, OC explants were exposed to different experimental conditions (n = 6 explants/group). Gentamicin (GM) was used to mimic an ototoxic environment leading to about 50% of HC loss without treatment (this ideal concentration of GM was determined in titration trials performed before the actual experiments). Group 1: GM only (positive control); group 2: GM and DXM in solution (75 μg mL −1 ); group 3: GM and DXMcoated and infused microneedle (DXM coating); group 4: GM and DXM-infused microneedle (no DXM coating); and group 5: only media without GM, DXM or microneedles (negative control). HC counts were determined via confocal microscopic examination of phalloidin-FITC stained OC explants and plotted for each of the five groups. DXM-coated microneedles showed a significant reduction of both inner hair cell (IHC) and outer hair cell (OHC) losses when compared to the non-coated prototype (p < 0.0001). A similar efficacy in protecting from HC loss was found when compared to DXM solution after 72 h of culture ( Figure 6). This was confirmed visually via confocal microscopy images (Figure 7). Thus, DXM-coated and infused microneedles protected both IHCs and OHCs, from the base to the apex of the cochlea, which encompasses all frequencies of hearing that can potentially cause ototoxicity, thus demonstrating the functionality of the developed technology in vitro.

Fabrication of Miniature Microneedles for In Vivo Models
To prepare the microneedles for insertion into the cochlea, a miniaturized mold was engineered via an SU-8 photolithography process in the Dr. JT Macdonald Foundation Biomedical Nanotechnology Institute at the University of Miami (BioN-IUM) Nanofabrication facility (Figure 8). The microneedle was designed with a length of 3.5 mm to sit in the base of the scala tympani (Figure 9). The average full-length dimensions of the rat scala tympani is 11 mm. This small parameter allows the microneedle to be inserted into the cochlea without fracture of the osseous spiral lamina, thus reducing modiolar injury. [29] The FM1-43FX microneedles were prepared by codissolving PLGA copolymer and FM1-43FX, a fixable fluorescent dye that behaves as a permanent blocker of the mechanotransducer channel, in DMSO ( Figure 8D). The final concentration of the FM1-43FX dye was 1% (w/w) of the PLGA. For microneedles carrying DXM, 25 mg of 50:50 PLGA copolymer and 2.5 mg of DXM were dissolved in 100 μL of DMSO over 3 h at RT, and this homogenous solution was then cast into the PDMS mold. As each microneedle weighs ≈300 μg, each microneedle contains 50 μg of DXM and 250 μg of PLGA, indicating the microneedle drug loading efficiency is 16.6%.

Physical Characterization of Polymeric Microneedles
The DXM polymeric microneedles were characterized for their mechanical properties and imaged to investigate the surface and integrity of the samples. Scanning electron microscopy (SEM) analysis revealed that the microneedles have a smooth outer surface, which allows for the microneedles to effortlessly pierce the RWM causing minimal insertion trauma. [30] The SEM imaging also confirmed that the microneedles have a length of 3.5 mm and diameter of 400 μm as designed, yielding a tip area of 1.26 × 10 −7 m 2 (Supplementary Information). The nanohardness (H) was measured to be 40.35 ± 1.54 MPa, reduced elastic modulus  (Er) was 1.43 ± 0.06 GPa and the contact depth (hc) was 479.73 ± 10.04 nm (N = 5). The H threshold required to puncture the RWM of adult guinea pigs has been found to be 1.19 mN, which translates to 9.47 × 10 −3 MPa over the area of the designed microneedle tip, with the force required to puncture a human RWM being four to five times this value. [31,32] The measured H value of 40.35 ± 1.54 MPa demonstrated that the microneedles have over ≈4000× the mechanical strength required to puncture rodent RWMs, as well as ample force for use in human studies for future work. [33]

In Vivo Analysis of the Polymeric Microneedles
To study the impact the drug-infused microneedles have on combating inner ear HC loss, two in vivo studies were performed.
For study 1, FM1-43FX dye infused microneedles were implanted into the scala tympani of male C57BL/6 mice (n = 3) via a retroauricular approach for 3-7 days (Figure 10). At this time, cochleae were harvested and processed for microscopic visualization to determine the distribution of the dye along the cochlea's HCs. Ototoxicity was intentionally not used, since intact HCs are necessary for entry of the FM1-43FX dye. For study 2, Norway-Brown rats were used instead of mice due to the larger average dimensions of the basal turn of the rat scala tympani when compared with mice. Study 2 intended to determine the protective effects of the drug-infused microneedle against HL induced by introduction of a foreign object to the ear. For this study, animals were randomly divided into two groups (n = 3 per group): group 1: no ototoxicity with PLGA microneedle as a positive control, and group 2: no ototoxicity with DXM-infused microneedle. Subsequently, animals underwent hearing tests (see below) to assess functional  hearing outcomes and cochleae were harvested and processed for histological analysis after 1 week.

In Vivo Analysis Demonstrating Drug Release and Diffusion within the Cochlea
The miniaturized FM1-43FX-infused microneedles were implanted in the scala tympani of adult C57BL/6 mice for 1 week which confirmed slow release of the dye from the microneedles. A red signal was identified in the HCs at different levels of the cochlea suggesting that the dye had passed through the mechanotransduction apparatus, confirming the diffusion of the dye through the perilymph from the basal turn all the way to the apex of the cochlea (Figure 11).

In Vivo Feasibility Demonstration of the Functionality of the Polymeric Microneedles
Following successful demonstration of the migration of the FM1-43 dye to the HCs in vivo, further in vivo investigation was performed following the same procedure to demonstrate the feasibility of the proposed platform with DXM using microneedles with and without the drug.
Auditory brainstem response (ABR) data was collected to determine the hearing threshold of the animals throughout the experiment. Specifically, baseline ABR was recorded on day 0, the microneedle was inserted on day 1, and ABR data was collected 1 day, 3 days, and 1 week post insertion. This timeline was selected as previous studies using the same animal model and approach have shown significant differences between threshold ABR data and ABR data collected 7 days post cochlear implantation into the scala tympani through the RWM. [34] ABR data demonstrated that hearing thresholds of animals were lower on days 2-7 for each frequency analyzed in ears implanted with a DXM-coated and DXM-infused PLGA microneedle when compared to the group which received a control PLGA microneedle (Figure 12). An ABR hearing threshold of only 5-6 dB is the point of intensity in which an adult can detect the presence of a stimulus. [35] Thus, the ABR data from Figure 11 demonstrate reduced hearing thresholds across all frequencies of more than 10 dB on day 7, a reduction of which is large enough to affect hearing. This lower threshold reveals that that DXM is not only being released from the microneedle but is also reaching the appropriate HCs to effectively allow the animals to hear at lower levels consistently across all frequencies. The ABR readings are anticipated to return to baseline as the confocal microscopy of both groups of in vivo animals showed no hair cell death in the basal, middle, and apical turns of the cochlea indicating no HL associated with cochlear insertion trauma (Figure 13). This in vivo data demonstrates the functionality of the microneedle releasing DXM in real time over an extended period without introducing any undesirable side effects while improving hearing thresholds as a proof of concept for the developed platform. Extended in vivo analysis is currently being conducted to determine the long term otoprotective effects of the DXM microneedle with a larger sample size over an extended period of time.

Conclusion
There is need for novel drug delivery systems to treat complex conditions in the human body, especially those related to otologic disorders. An ideal drug delivery system for such applications should need minimal number of drug doses, injections, or insertions, depending on the approach, as well as have minimal side effects. While there are no FDA approved drugs developed specifically to treat inner ear disease, approved drugs are used that provide systematic relief. [36] Previous and current methods used to treat HL include systemic administration, IT injections, cochlear implants, nanoparticle formulations and various other methodologies that place drug delivery materials on the RWM directly for osmotic delivery. Systemic delivery often fails to deliver the required amount of drug to the inner ear due to the blood labyrinth barrier, which simultaneously requires higher doses of these drugs to cross. Unwanted side effects such as Figure 11. FM1-43 blended microneedles implanted in adult mice for 1 week to confirm dye release and distribution through the perilymph. A,B) Control adult mice cochleae; C,D) mice adult cochleae exposed to a dye blended microneedle. A-D) 40× images using confocal microscopy; E) 63× magnification of sample. The dye is being slowly released from the microneedles and reaches HCs passing through the mechanotransduction apparatus of the HC, confirming the diffusion of the dye through the perilymph.
irritability, hypertension, and organ damage are introduced after long term use of these drugs at high levels. [6] IT injections were introduced to the field to combat the downfalls of systemic delivery systems. However, IT injections deliver significantly smaller amounts of drug to the inner ear due to minimal diffusion through the RWM and drainage of the drugs out of the ear via the Eustachian tube. [12] Additionally, the pharmacokinetic profile of the drugs administered intratympanically to the inner ear are highly dependent on the physiochemical property of the drug itself. [37] Drug delivery materials were subsequently developed which were placed on or near the RWM, yet when these materials are displaced, the drug delivery is compromised. Furthermore, drugs with high molecular weight and low lipid solubility cannot traverse the RWM readily, even when directly applied onto the RWM. [38] A method to circumvent this issue was the introduction of polymeric based nanoparticles that are able to permeate the RWM. [39] However, the drug retention time in the inner ear has been limited. [40] Most recently, it has been demonstrated that DXM-coated silicone rods are able to release DXM over extended periods of time into the inner ear in vivo. [41] These data support our current claims that releasing DXM long term into the cochlea is beneficial, on the downside the drug coated silicone rods employed by Liebau et al. are not biodegradable and remain in the cavity, which may lead to foreign body response, immunotoxicity, and other long-term undesirable side effects. In our work, we demonstrated a solution to the various problems of the prior work described in the literature by incorporating the corticosteroid DXM, the anti-inflammatory drug of choice for treatment of conditions leading to HL, in a controlled delivery system that addresses the drawbacks associated with current delivery techniques such as minimal drug retention over extended periods of time as well as biodegradability, leaving no foreign bodies behind in the cochlea once the microneedle has fully dissolved. The biocompatible FDA approved polymer, PLGA, was chosen as the vehicle of choice due to its tunable degradation and mechanical characteristics. The 50:50 PLGA copolymer was selected, specifically, as this copolymer has equal amounts of lactic acid to glycolic acid, which has previously been shown to degrade the fastest. [42] DXM was chosen as the drug of choice to minimize HL due to extensive prior investigation. However, the microneedle development was designed to be compatible with other common drugs used for HL including mannitol and L-N-acetylcysteine (L-NAC) to allow for a wider future applicability. [43] The PLGA microneedle is designed to pierce the RWM, which spontaneously heals itself, in order to enter the cochlea. Due to the mechanical strength of the polymer, the microneedle can be introduced into the cochlea without the use of any additional tools and without the introduction of any toxic agents as was demonstrated through the in vivo ototoxicity studies. Once the biocompatibility of the PLGA polymer was confirmed, further in vitro investigation demonstrated that a necessary coating of DXM was required to protect 85%, 87%, and 100% of the OHCs in the basal, middle, and apical turns of the cochlea, respectively, and 100% of the IHCs. In summary, the proposed work addresses an unmet need in the field of continuous drug delivery and more specifically in Otolaryngology, by employing microengineering that will decrease and/or prevent HL and, thus, enhance the quality of life of many patients who suffer from otologic disorders with the  potential to lead to HL. This study provides useful insight into a promising tool that could potentially revolutionize the treatment of different debilitating conditions, especially in otology.

Experimental Section
Microneedle Fabrication: A custom positive mold was designed using Rhino 3D software (ver. 6; Rhinoceros Inc., Seattle, WA) and manufactured by employing a 3D printer (LulzBot TAZ5; Aleph Objects, Inc., Loveland, CO) using acrylonitrile butadiene styrene (ABS) (Filabot ABS 3mm Orange; Filabot, Inc., Barre, VT), a thermoplastic polymer, as the substrate ( Figure 1A-C). The ABS mold was treated with acetone vapor in a thin layer chromatography (TLC) chamber for 30 min to smooth the surface of the plastic. To create the PDMS mold, a silicone elastomer base was mixed in a 1:10 ratio with silicone elastomer curing agent using a Sylgard 184 Silicone Elastomer Kit (Dow Corning Co., Midland, MI). Once homogenous, the mixture was poured into the positive 3D-printed mold, placed under vacuum (BestValueVac, Inc., Naperville, IL) until all air bubbles were removed and left to cure overnight at room temperature (RT). The mold was placed at 65°C for 2 h the following day to eliminate any sticky residue and then removed from the positive 3D-printed mold manually with a spatula. The microneedles were prepared by pipetting a homogenous solution of 1:4 (w/v) 50:50 PLGA copolymer (Sigma Aldrich, St Louis, MO) dissolved in DMSO (VWR International Co., Radnor, PA) into the custom PDMS mold. The mold was left overnight at 65°C to allow the DMSO to evaporate, and the procedure was repeated until the mold was filled with PLGA copolymer. The microneedles were brought to RT and removed from the mold manually using the end circular structure, and the five hardened microneedles were cut from the excess material ( Figure 1E). For microneedles carrying DXM, 25 mg of 50:50 PLGA copolymer and 2.5 mg of DXM were dissolved in 100 μL of DMSO over 3 h at RT, and this homogenous solution was then cast into the PDMS mold and incubated overnight at 65°C to allow DMSO to evaporate before use. For the coated microneedles group, the microneedles were prepared by dipping them into a 10% DXM solution using DMSO as the solvent and allowed to dry for an additional 24 h in a desiccator before usage.
Animals for Data Collection: All animal protocols were approved by the Institutional Animal Care and Use Committee (IACUC) of the University of Miami (protocol #17-225) and were in full compliance with published National Institute of Health (NIH) guidelines for the care and use of laboratory animals (guide for the care and use of laboratory animals, 8th edition, The National Academies Press, Washington, DC, 2011). All animals were purchased from Charles River Laboratories, Wilmington, MA, USA and kept on a 12 h light/dark cycle and fed a standard diet upon arrival. For the in vitro studies, OC explants were obtained from 3-day-old Sprague-Dawley rat pups. For in vivo studies, male C57BL/6 mice weighing ≈20-30 g were used to assess the drug release profile from the microneedles and its distribution throughout the cochlea in short term experiments. Male Norway Brown rats weighing ≈250 g were used for long-term in vivo experiments to assess safety and efficacy of the microneedles.
Biocompatibility Studies: OC explants were dissected from 3-day-old Sprague-Dawley rat pups by euthanizing the animals and extracting whole temporal bones (n = 12). The whole OC explants were dissected en bloc, and one OC explant per well was placed into 24-well culture plates in 400 μL of complete serum-free media consisting of Dulbecco's modified Eagle's medium (DMEM) supplemented with glucose (final conc. 6 g L −1 ), N-1 supplement (1%), and penicillin G (500 U mL −1 ) for 72 h. First, experimental OCs were cocultured with a pure PLGA bead (n = 6), ≈3 mm in circumference, and placed ≈0.5-1 mm from the OC explant, control samples (no biopolymers) were cultured in absence of PLGA (n = 6). Second, OCs were cocultured with or without microneedles made by dissolving PLGA beads in either DMSO, acetone, or EtOAc as described in Section 2.1 (n = 6/group).
Drug Release Profile Study with Rhodamine B: A fluorescent compound, Rhodamine B, was used in place of DXM to study the drug release profile in an artificial perilymph solution. The Rhodamine B microneedles were prepared by dissolving PLGA copolymer and Rhodamine B in DMSO and the solution was cast into the PDMS mold. The final concentration of the Rhodamine B dye was 1% (w/w) of the PLGA. The prepared needles were then placed in a quartz cuvette containing an artificial perilymph solution as described by Salt et al. [44] The absorbance of the perilymph solution was measured every 30 min using a spectrophotometer until the absorbance of the solution at 552 nm was constant.
Extended Drug Release Profile Study with Dexamethasone: One DXMinfused microneedle structure per sample was placed into 100 μL of an artificial perilymph solution at 37°C. A sample was collected after 1, 2, 3, and 4 weeks for investigation of DXM concentration via LCMS-MS according to Bird and co-workers. [45] The obtained data was analyzed to determine the DXM concentration release per microneedle structure. This calculated data was then fitted to a variable slope non-linear sigmoidal dose-response curve using GraphPad Prism (GraphPad Software, La Jolla California USA) to generate the calibration curve for the DXM concentration release per microneedle.
In Vitro Investigation: OC explants were dissected from 3-day-old Sprague-Dawley rat pups by inducing hypothermia and then extracting whole temporal bones (n = 30 explants). One OC explant per well was placed into 24-well culture plates in 400 μL of complete serum-free media with or without 5 × 10 −6 m of Gentamicin (GM) to mimic an ototoxic environment leading to about 50% of HC loss without treatment (this ideal concentration of GM was determined in titration trials performed before the actual experiments). OC explants were exposed to different experimental conditions (n = 6 explants/group; group 1: GM only; group 2: GM and DXM in solution (75 μg mL −1 ); group 3: GM and DXM releasing needle (no DXM coating); group 4: GM and DXM releasing needle (DXM coating); and group 5: only media without GM, DXM or needles). For groups 2-4, GM was added to the media containing OCs 30 min before adding the drug or needle. This allowed sufficient time for the GM to diffuse to the HCs. The microneedle was placed ≈0.5-1 mm from the OC explant and the plates were cultured for 72 h.
Microneedle Preparation for In Vivo Investigation: To prepare the positive mold for the miniature microneedles, a silicon wafer was cleaned with ethanol, prebaked to remove residual moisture, coated with 100 μm thick SU-8 photoresist using CEE 200X Precision spin coater (Brewer Science, Inc., Rolla, MO) and soft baked to remove any bubbles created during the coating process. UV light was applied via an OAI Model 804 MBA Optical Mask Aligner (OAI, San Jose, CA) through a custom-made soda lime mask onto the wafer and the wafer was post-baked, developed, and hard baked again to harden the final elevated pattern (Figure 2A-C). PDMS was then cast over the wafer containing the desired pattern to create the corresponding negative mold.
Physical Characterization of Polymeric Microneedles: The DXM polymeric microneedles were characterized for their mechanical properties and imaged using a JOEL IT800 Ultrahigh Resolution Field Emission Scanning Electron Microscope (JOEL USA, Inc., Peabody, MA) to investigate the surface and integrity of the samples. SEM analysis revealed the microneedles have a smooth outer surface, which allows the microneedles to effortlessly pierce the RWM with minimal insertion trauma. [30] The SEM imaging also confirmed the microneedles to have a length of 3.5 mm and diameter of 400 μm as designed, yielding a tip area of 1.26 × 10 −7 m 2 (Supplementary Information). The nanohardness, H, and reduced elastic modulus, E r , were obtained by performing nanoindentation tests using a Hysitron TriboIndenter Nanomechanical Test System (Bruker, Billerica, MA). A load of 250 μN was applied to the sample at 22°C in a chamber with 35% relative humidity using a Diamond Berkovich indenter tip. The area of the residual indentation was measured and the nanohardness, H, was defined as the maximum load, P max , divided by the residual indentation area, A r as seen in Equation (1) Throughout the indentation process, the depth of penetration was recorded in relation to the load at that time. A load-displacement curve was generated, the slope of which indicated the stiffness, S, of the contact. The reduced elastic modulus, E r , was then calculated using Equation (2), where A p (h c ) is the area of the indentation at the contact depth, h c , and is a geometrical constant The nanohardness (H) was measured to be 40.35 ± 1.54 MPa, reduced elastic modulus (E r ) was 1.43 ± 0.06 GPa and the contact depth (h c ) was 479.73 ± 10.04 nm (N = 5). The H threshold required to puncture the RWM of adult guinea pigs has been found to be 1.19 mN, which translates to 9.47 × 10 −3 MPa over the area of the designed microneedle tip, with the force required to puncture a human RWM being four to five times this value. [31,32] The measured H value of 40.35 ± 1.54 MPa demonstrated the microneedles have over ≈4000× the mechanical strength required to puncture rodent RWMs, as well as ample force for use in human studies for future work. [33]