Electrospun Acetalated Dextran Scaffolds for Temporal Release of

May 31, 2013 - Copyright © 2013 American Chemical Society. *E-mail [email protected]; Tel 614-688-3797; Fax 614-292-7766 (K.M.A.)...
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Electrospun Acetalated Dextran Scaffolds for Temporal Release of Therapeutics Hassan M. Borteh,† Matthew D. Gallovic,‡ Sadhana Sharma,† Kevin J. Peine,§ Simeng Miao,†,⊥ Deanna J. Brackman,† Katie Gregg,‡ Yanyi Xu,∥ Xiaolei Guo,∥ Jianjun Guan,∥ Eric M. Bachelder,† and Kristy M. Ainslie*,†,‡ †

Division of Pharmaceutics, College of Pharmacy, ‡William G. Lowrie Department of Chemical and Biomolecular Engineering, College of Engineering, §Molecular, Cellular, and Developmental Biology Program, and ∥Department of Materials Science and Engineering, The Ohio State University, Columbus, Ohio 43210, United States ⊥ Vanderbilt University School of Engineering, Nashville, Tennessee 37235, United States S Supporting Information *

ABSTRACT: Electrospun acetalated dextran (Ac-DEX) scaffolds were fabricated to encapsulate resiquimod, an immunomodulatory toll-like-receptor (TLR) agonist. Ac-DEX has been used to fabricate scaffolds for sustained and temporal delivery of therapeutics because it has tunable degradation rates that are dependent on its synthesis reaction time or the molecular weight of dextran. Additionally, as opposed to commonly electrospun polyesters that shift the local pH upon degradation, the degradation products of Ac-DEX are pH-neutral: dextran, an alcohol, and the metabolic byproduct acetone. Formulations of Ac-DEX with two different degradation rates were used in this study. The effects of electrospinning conditions on the scaffold size and morphology were examined as well as fibroblast adhesion as imaged with fluorescence microcopy and scanning electron microscopy. Macrophage (MΦ) viability further indicates that the scaffolds are cytocompatible. Also, the controlled release profiles of resiquimod from loaded scaffolds and nitric oxide (NO) production by MΦ incubated with these scaffolds show the potential for Ac-DEX scaffolds to be used to temporally and efficiently deliver therapeutics. Overall, we present a novel scaffold that can have tunable and unique drug release rates for tissue engineering, drug delivery, immunomodulation, and wound healing applications.

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sustained delivery of small molecules and proteins.10 As a result of the limitations of existing polymers, there is a need for the development of new polymers for tissue engineering and drug delivery. One alternative polymer is acetalated dextran (Ac-DEX). AcDEX is a newly developed biodegradable polymer derived from dextran, which is a polysaccharide of glucose and FDA approved as a blood expander. In Ac-DEX’s structure, acetal groups replace the parent hydroxyl groups on the glucose backbone.12 The acetal groups can be either acyclic or cyclic, with the number of cyclic acetals increasing as reaction time progresses. Since acyclic acetals undergo hydrolysis at a faster rate than cyclic acetals, tunable degradation rates that depend on the synthesis reaction time of the polymer or the molecular weight of dextran can be achieved.13,14 Additionally, Ac-DEX does not produce acidic byproducts like those created by polyesters and therefore aids in the maintenance of the tissue’s homeostatic microenvironment, whether it is implanted or used topically. The degradation products of Ac-DEX are acetone, an

INTRODUCTION Unique temporal release of proteins, small molecules, and other therapeutics can be advantageous for systems such as stem cell therapies, immunotherapy, wound healing, and tissue engineering.1−3 Polyesters such as poly(lactic-co-glycolic acid) (PLGA), polycaprolactone (PCL), and poly(lactic acid) (PLA) have all been used as electrospun scaffolds for these applications;4,5 however, the development of a scaffold that can release cargo temporally using the same material has been limited because these polyesters have fixed degradation rates that are on the order of months to years. In addition, their byproducts are acidic and lower the pH of the microenvironment, which can drastically affect the bioactivity of sensitive cargo or diminish cellular function.6−9 In addition to polyesters, scaffolds are also fabricated from naturally occurring polymers such as collagen, but these have their own limitations as well. It is difficult to have a relatively stable un-cross-linked collagen scaffold in culture due to a very high degradation rate that quickly disrupts the integrity of the scaffold.10,11 Moreover, chemically crosslinked collagen scaffolds do not degrade in physiological pH conditions, and glutaraldehyde or other cross-linking chemicals may cause cytotoxicity to cells.10,11 Therefore, collagen has to be blended with other synthetic polymers like PCL for © XXXX American Chemical Society

Received: February 14, 2013 Revised: May 24, 2013

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as well as present a new scaffold that can be broadly applied as a topical or injectable therapy.

alcohol, and water-soluble dextran, which do not change the pH of the local microenvironment.12 These advantages make AcDEX a suitable polymer to fabricate three-dimensional scaffolds for the temporal release of therapeutics. There are several methods by which scaffolds can be fabricated including solvent casting, gas foaming, and porogen leaching. In the solvent casting method, a polymeric solution is added into a mold, and the solvent evaporates over time to form a polymeric membrane. Because an organic solvent is used in this process, it can result in denaturation of proteins during fabrication and residual solvent may have an adverse effect on cells. Another commonly used technique is gas foaming, which uses high-pressure carbon dioxide to generate pores which are commonly not interconnected. Interconnected pores can be important with cells infiltration into the scaffold, thus limiting the use of these scaffolds in vivo. A porogen leaching method uses dispersion of a porogen (e.g., salt, sugar, wax) in the polymeric solution. The solvent is evaporated, and the porogen is leached in a solvent that is usually water. As with gas foaming, limited control over pores interconnectivity and size may limit its applications, in addition to residual concentrations of porogens. Other methods such as freeze-drying, self-assembly, and phase separation have one or more processing limitation that restrict their broad application.15 On the other hand, electrospinning is a simple, inexpensive, and versatile technique that can produce scaffolds with interconnected pores and threedimensional structure with elements ranging in size from nanometers to micrometers.16 In addition, scaffolds with desired size and morphology can be fabricated by varying electrospinning parameters.17 In this method, a continuous fiber of electrostatically charged polymer travels toward a highvoltage collector via a syringe capillary that contains a reservoir of a conductive polymer solution. Electrospinning has been used to encapsulate small molecules and proteins into scaffolds that result in a sustained release of the compound.17,18 Although Ac-DEX scaffolds could encapsulate hydrophilic (e.g., proteins) or hydrophobic compounds, in this work, an Ac-DEX scaffold was formulated via electrospinning to encapsulate the hydrophobic immunomodulatory drug resiquimod. Resiquimod (4-amino-α,α-dimethyl-2-ethoxymethyl-1Himidazo[4,5-c]quinoline-1-ethanol) is part of a class of toll-like receptor (TLR) 7/8 stimulating molecules called imidazoquinolinamines.19,20 As a result of its ability to stimulate T-helper 1 (Th1) responses, resiquimod has been used as a topical cream for treatment of cutaneous diseases caused by Leishmania spp., type 2 herpes simplex virus (HSV-2), and human papilloma virus (HPV).19,21,22 However, the compound has to be applied frequently to maintain efficacious levels and cannot be delivered parenterally as a result of its hydrophobicity. Therefore, new delivery methods such as topical polymer scaffolds should be explored to decrease drug dosing frequency and increase the efficiency with which the drug is delivered. To illustrate the application of this novel polymer in a scaffold of this type, system parameters such as flow rate and polymer concentration were varied using biofriendly ethanol as the polymer solvent. The cytocompatibility of the scaffold was assessed by the ability of fibroblasts to adhere to the scaffolds and the viability of RAW macrophages (MΦ). Temporal release of resiquimod from AcDEX scaffolds with two different degradation rates was evaluated. Furthermore, the capability of resiquimod-loaded scaffolds to stimulate nitric oxide (NO) production by MΦs was investigated. It is the goal of the work here to illustrate a feasible topical therapy for the treatment of cutaneous disease

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EXPERIMENTAL SECTION

Materials. All chemicals, unless mentioned otherwise, were purchased from Sigma-Aldrich (St. Louis, MO) and used as received. RAW 264.7 macrophages were purchased from ATCC (Manassas, VA). CellTracker Green (5-chloromethylfluorescein diacetate, CMFDA) and propidium iodide (PI) were purchased from Life Technologies (Grand Island, NY). The Griess Reagent System was purchased from Promega (Madison, WI). Synthesis of Acetalated Dextran. Ac-DEX was synthesized and characterized based on the previously described method.12 Briefly, lyophilized dextran (molecular weight = 71 kDa) and pyridinium ptoluenesulfonate were dissolved in dimethyl sulfoxide (DMSO) and reacted with 2-methoxypropene under anhydrous conditions for 5 min (fast degrading polymer) and 6 h (slow degrading polymer). The reaction was quenched after each desired time with triethylamine (TEA). Basic water (pH 9.0) was made by adding 2% v/v of TEA to nanopure water (Millipore Milli-Q Integral Water Purification System, Billerica, MA). The polymer was precipitated by slowly adding the reaction mixture to stirring basic water. The precipitated polymer was filtered through paper (8 μm pore size) and lyophilized overnight. The lyophilized powder was dissolved in 200 proof ethanol and centrifuged at 10 000 rpm (Allegra 25R centrifuge, Beckman Coulter, Indianapolis, IN) for 10 min to further purify the polymer. The supernatant was transferred to a different container and resuspended in basic nanopure water (pH 9.0) and lyophilized until dry. Based on a previously developed method, the extent of cyclic acetal coverage for the fast and slow degrading scaffolds was calculated via nuclear magnetic resonance (NMR).12 Electrospinning Setup and System Parameters. Solutions of fast and slow degrading Ac-DEX in ethanol were prepared at concentrations of 0.20, 0.25, and 0.40 g/mL. Each solution was loaded into a syringe attached with a 21 gauge stainless steel capillary. A bias of 30 kV (−15 kV to the aluminum foil collection surface and +15 kV to the needle) was applied over a 20 cm needle-to-surface distance. In addition to observing the effects of the concentration on fiber size and morphology, a fixed 0.40 g/mL polymer solution was electrospun at 2, 4, and 6 mL/h flow rates to study the effect of flow rate on the size of fibers. 0.1% w/w of resiquimod was dissolved with fast or slow degrading Ac-DEX in ethanol. The flow rate and concentration were set to 4 mL/h and 0.40 g/mL, respectively, for the fabrication of unloaded or loaded scaffolds being used in all in vitro experiments. Fluorescent Imaging of Cell Adhesion. Polydimethylsiloxane (PDMS) was made by mixing the base polymer with 1% w/w of the curing agent. The cured PDMS was cut into small pieces (1 cm × 1 cm). The fast and slow degrading scaffolds without resiquimod (blank scaffolds) were cut into smaller squares (1 cm × 1 cm) and mounted along the edges of the PDMS pieces by uncured PDMS glue (Figure 1). The PDMS glue was applied such that it did not interfere with the scaffold. NIH 3T3 fibroblast cells were cultured in Thermo Scientific HyClone DMEM/High Glucose (Logan, UT) media with 10% fetal bovine serum and 1% penicillin/streptomycin. The cells were allowed to become confluent in a 75 cm2 flask and then were detached from

Figure 1. A schematic of PDMS/scaffold construct used for cellular assays. Top view (top image) and side view (bottom image) of the construct. B

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Figure 2. SEM images displaying morphology of electrospun fast degrading Ac-DEX as a function of electrospinning conditions: (A) 0.20, (B) 0.25, and (C) 0.40 g/mL at a flow rate of 4 mL/h; (D) 2, (E) 4, and (F) 6 mL/h using a concentration of 0.40 g/mL. Scale bars = 30 μm. the flask with trypsin, centrifuged, and then seeded on the PDMSframed Ac-DEX scaffolds. The PDMS frames were used to aid in keeping the scaffold fully submerged in the cell culture media. The samples (3 × 104 cells/sample) were placed into 6-well tissue culture plates for 24 h. After this time, viability of the cells adhered to the scaffolds was studied by staining the cells with CellTracker Green and PI to count the number of live and dead cells, respectively. CellTracker Green was dissolved in DMSO to a concentration of 10 mM. The solution was diluted to a final concentration of 20 μM by adding DMEM high glucose serum-free media. For PI staining, a 1 mg/mL of PI solution in DI water was prepared. The media was removed, and the cells (triplicate samples of each group) were incubated with 250 μL of CellTracker Green solution and 0.5 μL of PI solution for 30 min at 37 °C. The cells were washed with PBS twice, and live cell imaging was conducted using a Nikon Eclipse 800 (Melville, NY) fluorescence microscope at excitation wavelengths of 492 and 535 nm for CellTracker Green and PI, respectively. The live/dead cells were counted using ImageJ (National Institutes of Health, Bethesda, MD) software. Empty PDMS tabs that were the same size as the framed scaffolds and glass coverslips were used as controls. Scanning Electron Microscopic Imaging of Cell Integration. For imaging the cells within the scaffolds, the cells were seeded and incubated for 24 h as described above. The media was aspirated from the wells. The scaffolds were washed twice with phosphate buffer saline (PBS) (pH 7.4). The cells were fixed using a 2% glutaraldehyde solution in molecular grade water (G-Biosciences, St. Louis, MO) for 10 min. The fixture was aspirated, and the scaffolds were washed with PBS two more times. The samples were allowed to air-dry overnight at room temperature. Samples of the seeded scaffolds were cut into small pieces and mounted on scanning electron microscopy (SEM) stubs. All samples were coated with gold−palladium using a sputter coater prior to SEM imaging. The cells within the scaffolds were characterized using a FEI Nova 400 NanoSEM (FEI Inc., Hillsboro, OR). Encapsulation Efficiency of Resiquimod. Triplicate samples of fast or slow degrading scaffolds were dissolved overnight at 37 °C in a mixture of 990 μL of PBS and 5 μL of 50% v/v formic acid in water. The solutions were neutralized by adding 5 μL of 13.25 M of sodium hydroxide to each sample the following day. After neutralization, 150 μL of the samples was pipetted into a 96-well microplate. The amount of loaded resiquimod was quantified using its autofluorescence (excitation: 260 nm/emission: 360 nm) on a Spectra Max Gemini XS microplate reader (Molecular Devices, Sunnyvale, CA).

Cell Viability Assay. The viability of MΦ cells after incubation with Ac-DEX scaffolds was examined by a 3-(4,5-dimethylthiazol-2-yl)2,5-diphenyltetrazolium bromide (MTT) assay. 50 μL of a 5 mg/mL of the MTT solution was added to separate wells of cells cultured with the scaffolds and the PDMS and coverslip controls. The MTT solution was incubated with the cells for 30 min. The media from the wells was gently aspirated, 4 mL of isopropanol was added to each well, and the alcohol was pipetted up and down gently to dissolve the purple crystals formed during the assay. Then, 150 μL of the purple solution from each well was transferred to a 96-well microplate. The absorbance of the solution was recorded at 560 and 670 nm (background absorbance). The MTT data were normalized to the viability of the cells incubated without any scaffolds. Release Profile of Resiquimod. Triplicates of the fast and slow degrading scaffolds with the same total amount of resiquimod were placed into SnakeSkin (Thermo Scientific, Rockford, IL) cylindrical membranes and attached to the bottom of 6-well tissue culture plates via cured PDMS. 7 mL of PBS (pH 7.4) was added to each well. Blank scaffolds were used as controls. The samples were agitated on a shaker plate at 37 °C. At predetermined time points, 100 μL of the supernatant was transferred to a 96-well microplate and stored at −20 °C. After all the samples were collected, they were warmed to room temperature. The fluorescence of the released resiquimod was again recorded by a Spectra Max Gemini XS microplate reader. The percentage of resiquimod released was normalized to the fluorescence reading for maximum release of the fast scaffolds since the scaffold had fully degraded by the final time point. Production of Nitric Oxide by Cells. MΦs were seeded in 12well plates at a density of 5 × 104 cells/well. The cells were allowed to adhere to the wells for 24 h, and then 4 mL of fresh media was then added to each well. Fast or slow degrading scaffolds with the same total mass of encapsulated resiquimod were placed in the wells on top of the cells and kept submerged in the media using the PDMS tabs as described above. Blank scaffolds and cells without any scaffolds were used as negative controls, and 100 ng/mL of free unencapsulted soluble lipopolysaccharide (LPS) was used as the positive control. The samples were incubated for 48 h. 50 μL of the supernatants from each well was transferred to a 96-well microplate for the Griess assay, which measures the amount of nitric oxide. The assay was performed based on the manufacturer’s protocol.

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RESULTS AND DISCUSSION Ac-DEX Scaffold Size and Morphology as a Function of Electrospinning Parameters. Electrospinning is a C

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Figure 3. Adhesion of fibroblasts on electrospun acetalated dextran (Ac-DEX) scaffolds with fast and slow degradation rates. Cells were seeded for 24 h and then stained with CellTracker Green (5-chloromethylfluorescein diacetate) and propidium iodide (PI) for fluorescence based live/dead cellular analysis. (A) Live cells on fast degrading Ac-DEX. (B) Dead cells on fast degrading Ac-DEX. (C) Merged image of (A) and (B) (all scale bars = 100 μm). SEM images of fibroblasts adhesion on fast (C) and slow (E) degrading polymers (scale bar = 10 μm). (F) Percentage of live and dead fibroblast cells attached on the substrates (normalized to total cells). Data are presented as mean plus standard deviation.

fabrication process that is affected by several parameters, many of which can be systematically examined and optimized. The effect of polymer flow rate and concentration on the Ac-DEX fiber size and morphology is shown in Figure 2. At the lowest concentration of polymer (0.20 g/mL), a mixture of particles and fibers is formed due to the solution’s relatively low viscosity and polymer entanglement (i.e., degree of polymer chain overlap). The beading of Ac-DEX fibers observed at lower concentrations is consistent with other polymers. Zong et al. showed biodegradable polymers at low concentrations maintained a viscosity that was still high enough to prevent complete jet breakup into individual droplets. This results in a bead-like structure connected by thin fibers.23 When the concentration of Ac-DEX is increased to 0.25 g/mL, similar beads are sparsely present within a fibrous network. Finally, at 0.40 g/mL, continuous Ac-DEX fibers that display a ribbon-like morphology are formed (without beads). Munir et al. proposed a model that provided a direct relationship between the polymer concentration and polymer entanglement number for polyvinylpyrrolidone (PVP), another biodegradable polymer.24 As the PVP concentration was increased at a fixed molecular weight of polymer, three successive types of structure were formed: particles, bead fibers, and continuous fibers. Ac-DEX behaves in a similar fashion. Previously, Ac-DEX was prepared at an initial concentration of 0.05 g/mL in ethanol and

microparticles were fabricated by electrospray, a method similar to electrospinning.25 In the present study, as the Ac-DEX concentration in ethanol is increased, the polymer entanglement number reaches and exceeds one threshold value that results in bead-like fibers and surpasses another level that results in continuous fibers (Figure 2A−C). This provides a simple means by which Ac-DEX scaffolds with multiple structures and function can be fabricated. It is assumed that the ribbon-like structure is the result of using ethanol, which has a low boiling point (76 °C). The solvent on the surface of the fiber evaporates more quickly than the solvent contained within the interior of the fiber.26 During the time that the fibers travel toward the collector and reside on the plate, the solvent in the interior evaporates over this longer period of time, which causes the external layer to then collapse and form a ribbon-like structure. These structures have been reported with other polymers such as PLGA, PCL, and collagen as a result of using solvents with a low boiling point.27,28 AcDEX fiber size also depends on the flow rate, which is consistent with previously reported observations of other polymers.29,30 As the flow rate increases (at a constant polymer concentration), the fiber width also increases (Figure 2D−F). Table S1 in the Supporting Information supports this conclusion as well as illustrates that an increase in polymer concentration yields an increase in fiber width. D

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Adhesion and Viability of Cells Seeded on the Scaffolds. Although microscopically the extracellular matrix (ECM) is a highly organized matrix of proteins and polysaccharides, more macroscopic observations depict a random bed of fibers, which is mimicked with the electrospun Ac-DEX scaffolds.31 Cells such as fibroblasts use the fibrous structure backbone of the ECM as a platform to initiate such functions for tissue regeneration and wound healing.31 Although the size of the ECM fibers is typically on the order of tens or hundreds of nanometers, the randomly oriented AcDEX scaffolds demonstrate the ability for fibroblasts to adhere to microstructures (Figure 3). In this study, fibroblast cells were seeded on the surface of the scaffolds, stained, and imaged by epi-fluorescent microscopy. Figure 3A−C illustrates the cells morphology on the scaffold surface. From the micrograph it can be seen that the cells are spread on the surface, illustrating the cytocompatibility of the surface.32 Furthermore, the micrographs highlight that the cells are fairly homogenously dispersed on the surface and are not clustered, which is indicative of cellular integration and compatibility.33 A closer inspection of the cell−fiber interaction is shown in SEM images of fibroblast cells that were seeded on the fast and slow degrading scaffolds (Figure 3D,E). Chen et al. have shown adhesion of fibroblasts only to the surface of electrospun PCL scaffolds,34 whereas the fibroblasts seeded on Ac-DEX scaffolds penetrate the surface and appear to integrate themselves within the three-dimensional network of the fibers. Moreover, as opposed to the successful adhesion of fibroblasts to Ac-DEX, the use of PCL has shown a reduction in fibroblast adhesion and proliferation as a result of its hydrophobicity.35 Cell integration is more representative of the in vivo environment and can mimic cell adherence to the ECM.36,37 Although a full examination of the fibroblast phenotype was not performed, the SEM images indicate the cells do spread and adhere to the fibers. Further work is necessary to determine specifically which adhesion molecules are involved in the cell− fiber interactions observed here. The results of the image analysis of the live/dead staining are shown in Figure 3F; the percentage of live/dead cells on the fast and slow degrading scaffolds is not significantly different than the glass and PDMS controls. This indicates that Ac-DEX forms a cytocompatible scaffold. Additionally, since Ac-DEX’s degradation byproducts are pH-neutral, scaffold degradation products would not adversely affect cell viability and proliferation, unlike commonly used polyesters.7 Moreover, cytocompatibility with the Ac-DEX system could be as a result of ethanol being used as the solvent to electrospin the polymer solution.38 Residues from harsh solvents commonly used for electrospinning PLGA and PCL, such as 1,1,1,3,3,3-hexafluoro2-propanol (HFIP), N,N-dimethylformamide (DMF), tetrahydrofuran (THF), and chloroform,39 have been shown to be cytotoxic or carcinogenic and have much lower LD50s than biofriendly ethanol.40−45 Although the amount of residual solvent that remains in the fibers is likely not significant enough to cause widespread cytotoxicity in vitro, there has been discussion that it could become more of a concern in the clinical setting when larger volumes of scaffolds might be used.46 Researchers commonly treat electrospun fibers in vacuo to remove any residual solvent, but this process gives rise to the possibility of damaging a protein therapeutic.47 Using a less harsh solvent such as ethanol reduces the need for postprocessing vacuum treatments.

In order to further study the cytocompatibility of Ac-DEX scaffolds, MΦs were incubated with blank and resiquimodloaded Ac-DEX fast and slow degrading scaffolds. An MTT assay to determine the viability of these cells showed no significant difference between MΦs cultured with the resiquimod-loaded scaffolds compared to the cells cultured with the blank scaffold, soluble LPS, or cells alone (Figure 4) (p

Figure 4. Viability of RAW macrophages incubated with blank scaffolds, scaffolds encapsulated with resiquimod, and controls (soluble lipopolysaccharide [LPS] and cells only). Data are presented as mean ± standard deviation.

> 0.05). Polyesters such as PLGA and PCL that are commonly electrospun into biomimetic fibers hydrolyze into acidic byproducts that can reduce the pH of a microenvironment down to 3.0 or lower.6,7,48,49 It is therefore not surprising that a study by Sung et al. showed decreased mouse aortic smooth muscle cell viability and growth after exposure to degrading PLGA scaffolds that was partially due to the rapidly accumulating acidic byproducts.7 However, unlike PLGA, pHneutral degradation byproducts of fast and slow degrading AcDEX polymers result in little cytotoxicity. The pH neutrality of Ac-DEX’s degradation byproducts, the polymer’s ability to be electrospun using biofriendly ethanol, the three-dimensional integration of the fibroblasts into the scaffold, and the viability of MΦs makes electrospun Ac-DEX an optimal candidate for tissue engineering, immunotherapy, and wound healing applications. Moreover, the micro/nanostructure of the AcDEX scaffolds likely mimics the ECM due to the fabrication technique50 and in addition to all of its biofriendly attributes lends to the interdigitation of the fibroblast observed with the micrographs presented (Figure 3A−E). Further studies with electrospun Ac-DEX fibers could likely result in the production of random or aligned nanofibers with a controlled porosity that even more closely mimic in vivo environments.37,51,52 Temporal Release of Bioactive Resiquimod from the Scaffolds. Displaying the temporal release and delivery of therapeutics to pertinent cells is a significant aspect of the novel Ac-DEX scaffold. Ac-DEX has a tunable, pH-dependent rate of hydrolysis12 that can vary from hours to months by changing either the cyclic acetal coverage of the polymer or the molecular weight of the dextran starting material.13,14 Resiquimod was encapsulated not only because it is a viable therapeutic for cutaneous infections but also to examine the tunable kinetics of drug release from fast and slow degrading Ac-DEX. Table 1 shows the encapsulation efficiency (EE) of scaffolds initially loaded with 0.1% w/w resiquimod. The fast degrading scaffold displays an EE approximately one and a half fold higher than E

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Table 1. Amount of Cyclic Acetal Coverage and Encapsulation Efficiency (EE) of 0.1% w/w Resiquimod in Fast and Slow Degrading Ac-DEX Scaffolds polymer

% cyclic acetal coverage

% EE

fast slow

38 63

98.9 63.8

that of the slow degrading scaffold. It is assumed that the lower EE in the slow degrading scaffold is caused by an increased cyclic acetal coverage, which results in a higher degree of hydrophobicity. And since both resiquimod and the fast degrading Ac-DEX have a mix of hydrophobic and hydrophilic moieties, it is reasonable to observe a higher EE in the fast degrading scaffold. Regardless, the 64% EE for the slow degrading polymer still results in a substantial load for this potent immunomodulatory drug. The release of resiquimod from the scaffolds was performed over 6 days at pH 7.4 (Figure 5). Since the degradation

Figure 6. Nitric oxide production after 48 h by RAW macrophages in presence of Ac-DEX scaffolds encapsulated with resiquimod. Data are presented as mean ± standard deviation. An asterisk represents a significant difference with respect to fast degrading scaffold encapsulated with resiquimod (p < 0.05).

degrading scaffold encapsulated with resiquimod and each of the other experimental groups (p < 0.05). It is important to note that there is no significant difference between the amount of NO produced by cells incubated with blank scaffolds and the cells only control. This confirms the polymer itself is not immunostimulatory, which would be crucial for use in certain applications where an immune response is undesired. According to the release data (Figure 5), after 48 h ∼8% of the drug is released from the slow degrading scaffold compared to 100% for the fast degrading scaffold. As with any degradable matrix where the drug is assumed to be uniformly dispersed in the polymer, drug release would occur as a result of diffusion, in the presence of a difference in chemical potential, and degradation. Our results indicate that the amount of the resiquimod released by the slow degrading scaffold in 48 h is not enough to induce NO production from MΦs. These results highlight that by encapsulating the same drug in different formulations of Ac-DEX (fast and slow degrading), temporal release of drugs and the subsequent cellular response can be controlled. Overall, the pH neutrality of Ac-DEX’s degradation byproducts, the polymer’s ability to be electrospun using a biofriendly solvent, the three-dimensional integration of the fibroblasts into the fibers, and the viability of MΦs cultured on the scaffold make electrospun Ac-DEX an excellent candidate not only for cutaneous therapy but also as a tissue scaffold implant. With implantation of a therapy, it is critical that homeostasis is maintained and the pH neutrality of Ac-DEX’s degradation products certainly lends itself to a more benign microenvironment than polyester-based scaffolds. Furthermore, we have illustrated that cells vital in wound healing, fibroblasts, and MΦs interact with the scaffold in a manner supportive of proper wound healing.53,54 As activation of TLR 2 and 4 have also been shown to create a proper wound healing response, lipopolysaccharides could be incorporated into the scaffold at low levels to further promote a wound healing response.55 Although it is well understood activation of the MyD88 pathway is critical in wound healing, it has not been fully elucidated if TLR 7/8 activation would promote proper wound healing. Future studies with our resiquimod loaded scaffolds could be performed to fully elucidate the role of TLR 7/8 activation in wound healing.

Figure 5. Release profile of resiquimod from fast and slow degrading Ac-DEX scaffolds in PBS buffer (pH 7.4). Data are presented as mean ± standard deviation.

products of Ac-DEX are pH-neutral, homeostasis would be maintained at this biological pH whether the scaffold is applied topically or implanted subcutaneously. The slow degrading scaffold has a slower rate of hydrolysis and more sustained drug release compared to the fast one due to its higher cyclic acetal coverage (Table 1), which was attained by simply extending the Ac-DEX synthesis reaction time. As shown in Figure 5, only 40% of the drug is released from the slow degrading scaffold over a period of 6 days compared to 100% release from the fast degrading scaffold after ∼60 h. This confirms that the release rate of the drug can be controlled temporally by using different forms of Ac-DEX. The ability to achieve a range of drug loadings in combination with the tunable kinetics of drug release further displays the potential benefits of electrospun AcDEX fibers. NO release is indicative of the successful delivery of resiquimod to MΦs since the drug is a TLR-agonist that induces this immune response. Figure 6 shows NO production by MΦs cultured for 48 h with either blank scaffolds or scaffolds encapsulated with resiquimod. There is a significant difference in the production of NO by MΦs between the fast F

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its tunable degradation, and the controlled temporal release of therapeutics make the Ac-DEX scaffold an excellent candidate for tissue engineering, wound healing, immunotherapy, and drug delivery applications.

Temporal control of drug or cell signal release rate is crucial for successful multidrug therapy, whether for cancer, HIV, antibacterial, or other treatments. For example, tissue regeneration and wound healing involve a series of steps, each of which possesses a unique requirement for the total time and rate of exposure of chemical signals and therapeutics.56−58 As seen in Figure 5, the tunability of Ac-DEX degradation is simple, in which different reaction times during its synthesis leads to varying rates of drug release. Encapsulating several therapeutics into an electrospun Ac-DEX scaffold would likely lead to sustained releases of the cargo at unique release rates. The results of this study suggest that Ac-DEX could be used in tissue engineering, immunotherapy, or wound healing applications where temporal release is desired, which is a very powerful and unique tool. Other researchers have discussed temporal release of multiple drugs encapsulated in the same fiber. Okuda et al. used a multilayered approach to encapsulate two model drugs in different fibrous layers.50 They did show temporal release of the two drugs, but their release data were performed in pH 8.6, conditions which are not exhibited in a significant biological environment. Dong et al. fabricated electrospun fibers that encapsulated protein-loaded nanoparticles.59 However, there was no indication of temporal release of this cargo and whether or not the protein remained active after delivery. Finally, Piras et al. have reported a dual-controlled drug release of two model molecules that showed a triggered or burst release over just a few hours in pH 7.4.60 This release at extracellular conditions may result in undesired side effects since all of the cargo was released within this short period of time. On the other hand, the time-controlled release of bioactive resiquimod at a relevant biological pH of 7.4 from two different formulations of Ac-DEX shows that each of the potential drawbacks of other scaffolds can be overcome by this system. The unique characteristics of the Ac-DEX polymer make it an excellent candidate for scaffolds to be used in the temporal release of therapeutics for applications such as tissue regeneration, wound healing, drug delivery, and immunotherapy.

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ASSOCIATED CONTENT

S Supporting Information *

Additional table. This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*E-mail [email protected]; Tel 614-688-3797; Fax 614-292-7766 (K.M.A.). Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors thank the staff at the Ohio State Campus Microscopy and Imaging Facility (CMIF) for their aid in acquiring SEM images.

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REFERENCES

(1) Jabbarzadeh, E.; Starnes, T.; Khan, Y. M.; Jiang, T.; Wirtel, A. J.; Deng, M.; Lv, Q.; Nair, L. S.; Doty, S. B.; Laurencin, C. T. Induction of angiogenesis in tissue-engineered scaffolds designed for bone repair: A combined gene therapy−cell transplantation approach. Proc. Natl. Acad. Sci. U. S. A. 2008, 105, 11099−11104. (2) Rabbany, S. Y.; Pastore, J.; Yamamoto, M.; Miller, T.; Rafii, S.; Aras, R.; Penn, M. Continuous delivery of stromal cell-derived factor-1 from alginate scaffolds accelerates wound healing. Cell Transplant. 2010, 19, 399−408. (3) Biondi, M.; Ungaro, F.; Quaglia, F.; Netti, P. A. Controlled drug delivery in tissue engineering. Adv. Drug Delivery Rev. 2008, 60, 229− 242. (4) Hong, Y.; Fujimoto, K.; Hashizume, R.; Guan, J.; Stankus, J. J.; Tobita, K.; Wagner, W. R. Generating elastic, biodegradable polyurethane/poly(lactide-co-glycolide) fibrous sheets with controlled antibiotic release via two-stream electrospinning. Biomacromolecules 2008, 9, 1200−1207. (5) Kenawy, E.-R.; Abdel-Hay, F. I.; El-Newehy, M. H.; Wnek, G. E. Processing of polymer nanofibers through electrospinning as drug delivery systems. Mater. Chem. Phys. 2009, 113, 296−302. (6) Lu, L.; Peter, S. J.; Lyman, M. D.; Lai, H. L.; Leite, S. M.; Tamada, J. A.; Uyama, S.; Vacanti, J. P.; Langer, R.; Mikos, A. G. In vitro and in vivo degradation of porous poly(dl-lactic-co-glycolic acid) foams. Biomaterials 2000, 21, 1837−1845. (7) Sung, H.-J.; Meredith, C.; Johnson, C.; Galis, Z. S. The effect of scaffold degradation rate on three-dimensional cell growth and angiogenesis. Biomaterials 2004, 25, 5735−5742. (8) Sachlos, E.; Czernuszka, J. T. Making tissue engineering scaffolds work. Review: The application of solid freeform fabrication technology to the production of tissue engineering scaffolds. Eur. Cells Mater. 2003, 5, 29−39 ; discussion 39−40. (9) Fu, K.; Pack, D.; Klibanov, A.; Langer, R. Visual evidence of acidic environment within degrading poly(lactic-co-glycolic acid) (plga) microspheres. Pharm. Res. 2000, 17, 100−106. (10) Lu, Q.; Ganesan, K.; Simionescu, D. T.; Vyavahare, N. R. Novel porous aortic elastin and collagen scaffolds for tissue engineering. Biomaterials 2004, 25, 5227−5237. (11) Lim, S. H.; Mao, H.-Q. Electrospun scaffolds for stem cell engineering. Adv. Drug Delivery Rev. 2009, 61, 1084−1096. (12) Bachelder, E. M.; Beaudette, T. T.; Broaders, K. E.; Dashe, J.; Frechet, J. M. J. Acetal-derivatized dextran: An acid-responsive biodegradable material for therapeutic applications. J. Am. Chem. Soc. 2008, 130, 10494−10495.

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CONCLUSIONS Ac-DEX, a newly developed polymer, was used to fabricate electrospun scaffolds with fast and slow degradation that resulted in the temporal delivery of a therapeutic. A direct correlation between the electrospinning parameters such as flow rate and polymer concentration on fiber width and morphology was confirmed, and the fibers displayed a ribbonlike morphology. Fluorescent and SEM images of adhered fibroblasts show that the cells do not simply reside on the surface of the scaffolds but also incorporate themselves within the three-dimensional network of the matrix by adhering to the surface of individual fibers. Moreover, no significant difference was observed in the viability of MΦs incubated with the resiquimod-loaded scaffolds with respect to blank scaffolds or the cells only. Tunability of the scaffolds with different reaction times (fast and slow degradation) was confirmed by the release profile of the encapsulated resiquimod in pH 7.4. A burst release rate was achieved with the fast degrading scaffold as compared to a sustained release from the slow one. Finally, a differential NO production by macrophages using the fast versus slow degrading scaffolds was observed. The resiquimod released from the fast degrading scaffold after 48 h of incubation was enough to induce NO production in MΦ. Its cytocompatibility, the flexibility in its synthesis and fabrication, G

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(35) Zhang, Y. Z.; Venugopal, J.; Huang, Z. M.; Lim, C. T.; Ramakrishna, S. Characterization of the surface biocompatibility of the electrospun pcl-collagen nanofibers using fibroblasts. Biomacromolecules 2005, 6, 2583−2589. (36) Sun, S.; Titushkin, I.; Cho, M. Regulation of mesenchymal stem cell adhesion and orientation in 3d collagen scaffold by electrical stimulus. Bioelectrochemistry 2006, 69, 133−141. (37) Lee, J.; Cuddihy, M. J.; Kotov, N. A. Three-dimensional cell culture matrices: State of the art. Tissue Eng., Part B 2008, 14, 61−86. (38) Zha, Z. B.; Teng, W. B.; Markle, V.; Dai, Z. F.; Wu, X. Y. Fabrication of gelatin nanofibrous scaffolds using ethanol/phosphate buffer saline as a benign solvent. Biopolymers 2012, 97, 1026−1036. (39) Li, W.-J.; Cooper, J. A., Jr.; Mauck, R. L.; Tuan, R. S. Fabrication and characterization of six electrospun poly(α-hydroxy ester)-based fibrous scaffolds for tissue engineering applications. Acta Biomater. 2006, 2, 377−385. (40) Mengerink, Y.; van der Wal, S.; Claessens, H. A.; Cramers, C. A. Analysis of higher polyamide-6 oligomers on a silica-based reversedphase column with a gradient of formic acid as compared with hexafluoroisopropanol. J. Chromatogr. 2000, 871, 259−268. (41) Druckrey, H.; Preussmann, R.; Ivankovic, S.; Schmahl, D. [organotropic carcinogenic effects of 65 various n-nitroso- compounds on bd rats]. Z. Krebsforsch. 1967, 69, 103−201. (42) Patnaik, P. A Comprehensive Guide to the Hazardous Properties of Chemical Substances; Van Nostrand Reinhold: New York, 1992; pp 540−549. (43) Wang, G.; Bai, N. Structure-activity relationships for rat and mouse ld50 of miscellaneous alcohols. Chemosphere 1998, 36, 1475− 1483. (44) Nam, J.; Huang, Y.; Agarwal, S.; Lannutti, J. Materials selection and residual solvent retention in biodegradable electrospun fibers. J. Appl. Polym. Sci. 2008, 107, 1547−1554. (45) Fontes, S. T.; Fernandez, M. R.; Ogliari, F. A.; de Carvalho, R. V.; de Moraes, R. R.; Pinto, M. B.; Piva, E.. Tetrahydrofuran as solvent in dental adhesives: Cytotoxicity and dentin bond stability. Clin. Oral Invest. 2013, 17, 237−242. (46) Dalton, P. D.; Klinkhammer, K.; Salber, J.; Klee, D.; Moller, M. Direct in vitro electrospinning with polymer melts. Biomacromolecules 2006, 7, 686−690. (47) Carpenter, J. F.; Pikal, M. J.; Chang, B. S.; Randolph, T. W. Rational design of stable lyophilized protein formulations: Some practical advice. Pharm. Res. 1997, 14, 969−975. (48) Fu, K.; Pack, D. W.; Klibanov, A. M.; Langer, R. Visual evidence of acidic environment within degrading poly(lactic-co-glycolic acid) (plga) microspheres. Pharm. Res. 2000, 17, 100−106. (49) Liu, H.; Slamovich, E. B.; Webster, T. J. Less harmful acidic degradation of poly(lacticco-glycolic acid) bone tissue engineering scaffolds through titania nanoparticle addition. Int. J. Nanomed. 2006, 1, 541−545. (50) Okuda, T.; Tominaga, K.; Kidoaki, S. Time-programmed dual release formulation by multilayered drug-loaded nanofiber meshes. J. Controlled Release 2010, 143, 258−264. (51) Stevens, M. M.; George, J. H. Exploring and engineering the cell surface interface. Science 2005, 310, 1135−1138. (52) Karageorgiou, V.; Kaplan, D. Porosity of 3d biomaterial scaffolds and osteogenesis. Biomaterials 2005, 26, 5474−5491. (53) Diegelmann, R. F.; Evans, M. C. Wound healing: An overview of acute, fibrotic and delayed healing. Front. Biosci. 2004, 9, 283−289. (54) Clark, R. A. F.; Ghosh, K.; Tonnesen, M. G. Tissue engineering for cutaneous wounds. J. Invest. Dermatol. 0000, 127, 1018−1029. (55) Ioannou, S.; Voulgarelis, M. Toll-like receptors, tissue injury, and tumourigenesis. Mediators Inflammation 2010, DOI: 10.1155/ 2010/581837. (56) Saltzman, W. M.; Olbricht, W. L. Building drug delivery into tissue engineering design. Nat. Rev. Drug Discovery 2002, 1, 177−186. (57) Lutolf, M. P.; Hubbell, J. A. Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nat. Biotechnol. 2005, 23, 47−55.

(13) Broaders, K. E.; Cohen, J. A.; Beaudette, T. T.; Bachelder, E. M.; Frechet, J. M. Acetalated dextran is a chemically and biologically tunable material for particulate immunotherapy. Proc. Natl. Acad. Sci. U. S. A. 2009, 106, 5497−5502. (14) Kauffman, K. J.; Kanthamneni, N.; Meenach, S. A.; Pierson, B. C.; Bachelder, E. M.; Ainslie, K. M. Optimization of rapamycin-loaded acetalated dextran microparticles for immunosuppression. Int. J. Pharm. 2012, 422, 356−363. (15) Subia, B.; Kundu, J.; Kundu, S. C. Biomaterial Scaffold Fabrication Techniques for Potential Tissue Engineering Applications, 1st ed.; InTeck: Rijeka, Croatia, 2010; p 524. (16) Frenot, A.; Chronakis, I. S. Polymer nanofibers assembled by electrospinning. Curr. Opin. Colloid Interface Sci. 2003, 8, 64−75. (17) Sill, T. J.; von Recum, H. A. Electrospinning: Applications in drug delivery and tissue engineering. Biomaterials 2008, 29, 1989− 2006. (18) Nie, H.; Wang, C.-H. Fabrication and characterization of plga/ hap composite scaffolds for delivery of bmp-2 plasmid DNA. J. Controlled Release 2007, 120, 111−121. (19) Dockrell, D. H.; Kinghorn, G. R. Imiquimod and resiquimod as novel immunomodulators. J. Antimicrob. Chemother. 2001, 48, 751− 755. (20) Bishop, G. A.; Ramirez, L. M.; Baccam, M.; Busch, L. K.; Pederson, L. K.; Tomai, M. A. The immune response modifier resiquimod mimics cd40-induced b cell activation. Cell. Immunol. 2001, 208, 9−17. (21) Garnier, T.; Croft, S. L. Topical treatment for cutaneous leishmaniasis. Curr. Opin. Invest. Drugs 2002, 3, 538−544. (22) Mark, K. E.; Corey, L.; Meng, T. C.; Magaret, A. S.; Huang, M. L.; Selke, S.; Slade, H. B.; Tyring, S. K.; Warren, T.; Sacks, S. L.; Leone, P.; Bergland, V. A.; Wald, A. Topical resiquimod 0.01% gel decreases herpes simplex virus type 2 genital shedding: A randomized, controlled trial. J. Infect. Dis. 2007, 195, 1324−1331. (23) Zong, X.; Kim, K.; Fang, D.; Ran, S.; Hsiao, B. S.; Chu, B. Structure and process relationship of electrospun bioabsorbable nanofiber membranes. Polymer 2002, 43, 4403−4412. (24) Munir, M. M.; Suryamas, A. B.; Iskandar, F.; Okuyama, K. Scaling law on particle-to-fiber formation during electrospinning. Polymer 2009, 50, 4935−4943. (25) Duong, A. D.; Sharma, S.; Peine, K. J.; Gupta, G.; Satoskar, A. R.; Bachelder, E. M.; Wyslouzil, B. E.; Ainslie, K. M. Electrospray encapsulation of toll-like receptor agonist resiquimod in polymer microparticles for the treatment of visceral leishmaniasis. Mol. Pharmaceutics 2013, 10, 1045−1055. (26) Koombhongse, S.; Liu, W.; Reneker, D. H. Flat polymer ribbons and other shapes by electrospinning. J. Polym. Sci., Part B: Polym. Phys. 2001, 39, 2598−2606. (27) Subbiah, T.; Bhat, G. S.; Tock, R. W.; Parameswaran, S.; Ramkumar, S. S. Electrospinning of nanofibers. J. Appl. Polym. Sci. 2005, 96, 557−569. (28) Orlova, Y.; Magome, N.; Liu, L.; Chen, Y.; Agladze, K. Electrospun nanofibers as a tool for architecture control in engineered cardiac tissue. Biomaterials 2011, 32, 5615−5624. (29) Deitzel, J. M.; Kleinmeyer, J.; Harris, D.; Beck Tan, N. C. The effect of processing variables on the morphology of electrospun nanofibers and textiles. Polymer 2001, 42, 261−272. (30) Demir, M. M.; Yilgor, I.; Yilgor, E.; Erman, B. Electrospinning of polyurethane fibers. Polymer 2002, 43, 3303−3309. (31) Alberts, B. Molecular Biology of the Cell, 5th ed.; Garland Science: New York, 2008; pp 1178−1195. (32) Ainslie, K.; Bachelder, E.; Sharma, G.; Grimes, C.; Pishko, M. Macrophage cell adhesion and inflammation cytokines on magnetostrictive nanowires. Nanotoxicology 2007, 1, 279−290. (33) Dunn, J. C.; Chan, W. Y.; Cristini, V.; Kim, J. S.; Lowengrub, J.; Singh, S.; Wu, B. M. Analysis of cell growth in three-dimensional scaffolds. Tissue Eng. 2006, 12, 705−716. (34) Chen, M.; Patra, P. K.; Warner, S. B.; Bhowmick, S. Role of fiber diameter in adhesion and proliferation of nih 3t3 fibroblast on electrospun polycaprolactone scaffolds. Tissue Eng. 2007, 13, 579−587. H

dx.doi.org/10.1021/la400541e | Langmuir XXXX, XXX, XXX−XXX

Langmuir

Article

(58) Tibbitt, M. W.; Anseth, K. S. Dynamic microenvironments: The fourth dimension. Sci. Transl. Med. 2012, 4, 160ps124. (59) Dong, B.; Smith, M. E.; Wnek, G. E. Encapsulation of multiple biological compounds within a single electrospun fiber. Small 2009, 5, 1508−1512. (60) Piras, A. M.; Chiellini, F.; Chiellini, E.; Nikkola, L.; Ashammakhi, N. New multicomponent bioerodible electrospun nanofibers for dual-controlled drug release. J. Bioact. Compat. Polym. 2008, 23, 423−443.

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