Poly(3,4-ethylenedioxythiophene) nanoparticle and poly(E-caprolactone) electrospun scaffold characterization for skeletal muscle regeneration Kristin D. McKeon-Fischer,* Daniel P. Browe,* Ronke M. Olabisi, Joseph W. Freeman Department of Biomedical Engineering, Rutgers University, Piscataway, New Jersey 08854 Received 18 November 2014; revised 30 March 2015; accepted 7 April 2015 Published online 25 June 2015 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35481 Abstract: Injuries to peripheral nerves and/or skeletal muscle can cause scar tissue formation and loss of function. The focus of this article is the creation of a conductive, biocompatible scaffold with appropriate mechanical properties to regenerate skeletal muscle. Poly(3,4-ethylenedioxythiophene) (PEDOT) nanoparticles (Np) were electrospun with poly(E-caprolactone) (PCL) to form conductive scaffolds. During electrospinning, ribboning, larger fiber diameters, and unaligned scaffolds were observed with increasing PEDOT amounts. To address this, PEDOT Np were sonicated prior to electrospinning, which resulted in decreased conductivity and increased mechanical properties. Multi-walled carbon nanotubes (MWCNT) were added to the 1:2 solution in an effort to increase conductivity. However, the addition of MWCNT had

little effect on scaffold conductivity, and the elastic modulus and yield stress of the scaffold increased as a result. Rat muscle cells attached and were active on the 1–10, 1–2, 3–4, and 1–1 PCL-PEDOT scaffolds; however, the 3–4 scaffolds had the lowest level of metabolic activity. Although the scaffolds were cytocompatible, further development of the fabrication method is necessary to produce more highly aligned scaffolds capable of promoting skeletal muscle cell alignment C and eventual regeneration. V 2015 Wiley Periodicals, Inc. J Biomed Mater Res Part A: 103A: 3633–3641, 2015.

Key Words: poly(3,4-ethylenedioxythiophene), poly(E-caprolactone), conductive nanoparticles, electrospinning, fibrous scaffolds

How to cite this article: McKeon-Fischer KD, Browe DP, Olabisi RM, Freeman JW. 2015. Poly(3,4-ethylenedioxythiophene) nanoparticle and poly(E-caprolactone) electrospun scaffold characterization for skeletal muscle regeneration. J Biomed Mater Res Part A 2015:103A:3633–3641.

INTRODUCTION

Traumatic injuries can damage both skeletal muscle and peripheral nerves, resulting in loss of function at neuromuscular junctions.1–7 Both skeletal muscle and peripheral nerves have a limited capacity to regenerate proportional to the amount of tissue loss.3,5,8 After skeletal muscle injury, satellite cells become activated, proliferate, and fuse to form multinucleated myotubes to repair the damaged site. However, the satellite cell population is finite and incapable of restoring the tissue after large volumetric loss, resulting in scar tissue formation and decreased muscle function.8,9 Significant loss of muscle and muscle function may require surgical reconstruction or even amputation leading to loss of patient quality of life and lower self-esteem.7,8,10 No suitable restoration method exists for large volumetric skeletal muscle loss.6,7,9 After peripheral nerve injury, there can be partial or complete loss of motor, sensory, and autonomic functions.4 When peripheral nerve damage occurs, the axonal end attempts to reunite with its synaptic target and results in restored function as long as the gap is small.3,5,11 During surgical repair, these ends are directly sutured together to aid in the regeneration process.11 As the axonal regeneration rate for humans is only 2–5 mm/day, a large gap (> 3 cm in humans) will take

months to heal.5,11 The regeneration process must also compete with scar tissue formation within the gap, and if outpaced may lead to permanent loss of function.5,11 In such cases, an autograft or allograft is used; however, complications associated with these procedures include donor site morbidity, loss of function at the donor site, and size mismatch.11,12 An alternative approach involves the use of synthetic grafts in the form of tubes to bridge the two nerve ends together; unfortunately, these outcomes have been variable.11,12 Tissue engineering provides a way to create, repair, or replace damaged tissues and organs.7,13–16 The tissue engineering paradigm employs a combination of cells, biomaterial scaffolds, and chemical stimuli to regenerate lost tissue.13–15 Scaffold material properties affect cell migration, adhesion, growth, and differentiation.17,18 To support conductive tissues, conductive elements can be incorporated into scaffolds to provide a cellular environment that is similar to the native tissue. Scaffold conductivity has been shown to play a role in excitable tissue growth: both skeletal muscle and nerve tissues have shown enhanced tissue formation when an electrical stimulus is applied.3,19 The focus of this study is to investigate the use of conductive scaffolds for regenerating skeletal muscle. Our

*These authors contributed equally to this work. Correspondence to: J. W. Freeman; e-mail: [email protected]

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laboratory has incorporated several different conductive elements within electrospun scaffolds to aid in cell alignment and fusion.19–26 Most recently, our work has focused on creating a biocompatible artificial muscle that actuates when electrically stimulated which we have termed a “reverse” ionic polymer metal composite (IPMC). An IPMC consists of metal plated on both sides of a polymer layer, and the structure is capable of actuating when placed in an ionic solution and an electric field is applied.27 Our reverse IPMC is a coaxially electrospun scaffold with poly(E-caprolactone) and multi-walled carbon nanotubes (MWCNT) in the inner core of each fiber and a poly(acrylic acid)/poly(vinyl alcohol) hydrogel outer fiber shell. The composite scaffold had an elastic modulus and yield strength that exceeded those of skeletal muscle but was still capable of actuation when stimulated.23–25 However, the high modulus and yield strength may have limited the amount of actuation. The objective of this study was to investigate a conductive element that would increase scaffold conductivity while tuning the mechanical properties to better match those of native skeletal muscle. The conductive polymer poly(3,4-ethylenedioxythiophene) (PEDOT) in nanoparticle (Np) form was selected as the conductive agent for its high stability and conductivity.28,29 While previous articles have investigated conductive scaffolds for skeletal muscle tissue engineering, this article attempts to utilize PEDOT Np as a conductive element while tuning the elastic modulus and yield stress of the scaffold to match those of native muscle tissue. Several different concentrations of PEDOT were combined with PCL and electrospun to form scaffolds that were then characterized for their material properties and cytocompatibility with rat myoblasts. We hypothesized that the PCL-PEDOT Np scaffolds would have similar elastic moduli and yield stresses to native skeletal muscle and that the addition of PEDOT Np would enhance scaffold conductivity, making the scaffolds more suited for skeletal muscle regeneration. MATERIALS AND METHODS

Solutions PEDOT Np (Sigma-Aldrich) suspended in water were placed in an oven ranging in temperature from 35–54 C to evaporate the water. The nanoparticles were then dispersed in dimethylformamide (DMF) to make 1:10, 1:2, 3:4, and 1:1 PEDOT:DMF (v/v) solutions. PCL and dichloromethane (DCM) were then added to make a final solution containing 25% PCL (w/v). In two separate solutions, sonication was utilized to better disperse the 1:1 PEDOT:DMF solution prior to adding the PCL. The 1:1 PEDOT:DMF solution was sonicated using a Branson digital sonifier 450 (Danburg, CT) using pulses with 2 s on and 3 s off at an amplitude of 20% for 2 min. The entire 2 min sonication session was repeated three times. After sonication, PCL was then added to make either a 25% or 20% (w/v) solution. Another solution was produced containing 1:2 PEDOT:DMF (v/v), 0.025% MWCNT (w/v), and 25% PCL (w/v). This final solution was sonicated using the same procedure described above. Acronyms for each of the solutions are listed in Table I.

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TABLE I. PCL, PEDOT, and MWCNT Containing Solution Components, the Acronym for Each Solution, and the Positive Voltage Range Electrospinning Parameter

Solution

Acronym

Positive Voltage Range (kV)

25% PCL-1/10 PEDOT 25% PCL-1/2 PEDOT 25% PCL-3/4 PEDOT 25% PCL-1/1 PEDOT 20% PCL-1/1 PEDOT_Sonicated 25% PCL-1/1 PEDOT_Sonicated 25% PCL-1/2 PEDOT-0.025% MWCNT

1:10 1:2 3:4 1:1 20_1:1_S

13–17 12–13 16–19 9–15 17–19

25_1:1_S

17

25_1:2_S_MWCNT

15–19

Electrospinning Solutions were loaded into a 5-mL syringe with a 20-gauge blunt stainless steel needle and placed in a syringe pump set at an extrusion rate of 5 mL/h. A distance of 18 cm was set between the needle tip and the 3000 6 200 rpm rotating mandrel with a 5-cm diameter used to collect each scaffold. A 2-kV negatively charged plate was placed behind the rotating mandrel to aid in attracting the positively charged solutions. The positive voltage ranges for each solution are located in Table I. The positive voltage was adjusted for each solution to produce the most stable Taylor cone during the electrospinning process, which should produce the most consistent fiber morphology in the scaffold. Fiber diameter and fiber angle measurements Scanning electron microscopy (SEM) surface images were taken of each scaffold, and the resulting images were analyzed for fiber morphology, fiber diameter, and fiber angle with Image J software (version 1.47v).30 To determine average fiber diameter, a total of 36 fibers from three fields in two separate samples were analyzed for each scaffold. Fiber angle was quantified using 210 fibers from three fields in two separate samples with the angle measured relative to the vertical axis of the images. In addition, each fiber angle value was normalized to an average value of 0 for each scaffold. Conductivity As previously reported, four 1 cm 3 1 cm pieces were cut from each of the seven different scaffolds and were vacuum soaked in phosphate buffered saline (PBS) for 30 min with repressurization occurring every 10 min.21–24 A constant voltage of 20.0V was applied by an E3646A dual output DC power supply (Agilent Technologies, Santa Clara, CA). The current output was also measured with this device. Conductivity was calculated using the equation, r 5 ‘/(R 3 A), where ‘ is the length of the scaffold, R is the electrical resistance, and A is the cross-sectional area of the scaffold.31

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Mechanical properties Seven 1 cm x 4 cm scaffold strips were evaluated for tensile mechanical properties using an Instron 5869 device (Norwood, MA). Samples were vacuum soaked as described above prior to being placed in a 37 C PBS bath and tested in tension at a rate of 10% strain/min. All mechanical tests were performed in the direction parallel to the primary direction of the fiber alignment. The elastic modulus and yield stress were determined from the resulting stress versus train graphs. Cellular study Primary rat skeletal muscle cells were harvested as previously described.23,24 Briefly, skeletal muscle was obtained from the hind limbs of Sprague Dawley rats. Fascia and tendon were removed before the skeletal muscle was sectioned into smaller pieces. These pieces were placed in 150 cm2 tissue culture flasks with Dulbecco’s-modified Eagle’s medium (DMEM) containing 10% fetal bovine serum (FBS) and 1% penicillin/streptomycin. The media was changed two times per week and after approximately 2 weeks, the muscle tissue was removed. Rat muscle cells attached to the bottom of the flask were removed using trypsin, spun down to form a cell pellet, and resuspended in DMEM before being plated in a new 150 cm2 tissue culture flask. These cells were expanded to confluence and frozen at passage 2. Samples from the 1:10, 1:2, 3:4, and 1:1 scaffolds were cut to fit inside a 24-well plate. These were soaked in 70% ethanol for 30 min and allowed to dry before being glued to the bottom of each well using a silicon adhesive (Factor II, Lakeside, AZ) that had been exposed to UV light for sterilization purposes. Each side of the well plate was then exposed to UV light for 1 h and allowed to dry overnight. Scaffolds were washed with sterile PBS prior to the addition of media and placed into a humidified incubator for 2 days (37 C, 5% CO2). Scaffolds were then seeded with ;100,000 rat muscle cells in suspension (passage 4) and returned to the incubator for 1 h before adding media. The media was changed three times per week over the 4-week study period. Cytocompatibility, metabolic activity, and cellular attachment were determined using a CellTiter 96 AQueousOne Solution Cell Proliferation (MTS) Assay (Promega, Madison, WI) and fluorescent staining. For the MTS assay, each scaffold was removed from its original well and placed into a new one. Then, the MTS assay liquid and DMEM media were added to the new wells in a 1:5 ratio before being placed into the incubator. After 3 h, each solution was diluted 1:4 using dIH2O and the absorbance read using an infinite M200 pro spectrophotometer (Tecan, Morrisville, NC). For fluorescent imaging, scaffolds were fixed with 3.7% paraformaldehyde solution prior to staining for actin (Oregon GreenV TM phalloidin; Invitrogen, Carlsbad, CA) and DAPI (NucBlue Fixed Cell Stain; Life Technologies, Grand Island, NY). R

Statistical analysis Data from the different groups were compared using Tukey’s HSD post hoc test. Analyses were performed using KaleidaGraph 4.1.0 (Synergy Software, Reading, PA)

statistical software, and data were considered statistically significant when p < 0.05.

RESULTS

Figure 1 displays representative SEM surface images for each of the seven electrospun scaffolds. Among the fibers, ribboning and fiber fusion is evident. Average fiber diameters and ranges are shown in Table II. The 20_1:1_S fiber diameter was significantly smaller than all other scaffolds except for the other sonicated solution, 25_1:1_S, which was itself significantly smaller in fiber diameter than the 1:10 fibers. All other fiber diameter averages were not significantly different from one another. Figure 2 shows histograms for the fiber orientation for each of the seven scaffolds with normal curves overlaid. The standard deviations of the fiber angle distributions are shown in Table II. The 1:10 scaffold had the lowest standard deviation of the fiber angles at 15.7 and the 3:4 scaffold had the highest standard deviation of the fiber angles at 42.5 . Among the 1:10, 1:2, 3:4, and 1:1 scaffolds, increased fiber alignment (lower standard deviation) was observed with lower amounts of PEDOT while larger PEDOT amounts resulted in less fiber alignment. For the sonicated solutions, the 20_1:1_S scaffold had a higher degree of fiber alignment (standard deviation 5 27.0 ) than the 25_1:1_S scaffold (standard deviation 5 45.4 ). Finally, the 25_1:2_S_MWCNT scaffold had the lowest degree of fiber alignment among all of the scaffolds with a standard deviation of 56.7 . Conductivity values (average 6 standard deviation) are reported in Table III. The 1:10 scaffold had the lowest conductivity, and the conductivity of the scaffolds increased with the amount of PEDOT. The conductivity of the 3:4 scaffold was significantly higher than that of the 1:10 scaffold and the conductivity of the 1:1 scaffold was larger than that of all other scaffolds except for the 3:4 scaffold. The 20_1:1_S and 25_1:1_S scaffolds that were sonicated prior to electrospinning had a lower conductivity than the 1:1 scaffold. The conductivity of the 25_1:2_S_MWCNT scaffold was not significantly different from the 1:2 scaffold. There was less variation in the conductivity of the scaffolds that were sonicated prior to electrospinning than those scaffolds that were not sonicated. The elastic modulus and yield stress were determined for all seven scaffolds (Table IV). The elastic modulus and yield stress for the 1:10 scaffolds were significantly greater than the 3:4 and 1:1 scaffolds. The 20_1:1_S scaffold elastic modulus was significantly larger than the 1:2, 3:4, and 1:1 scaffolds. Additionally, the yield stress for the 20_1:1_S scaffold was significantly higher than the other six scaffolds. Yield stress for the 25_1:1_S scaffold was significantly greater than that of the 3:4 scaffolds. The addition of MWCNT increased the mechanical properties significantly compared to the 3:4 and 1:1 scaffolds. As sonication negatively affected the conductivity of the scaffold and the addition of MWCNT did not significantly improve the conductivity, the cell study was performed only

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FIGURE 1. SEM images for the (a) 1:10, (b) 1:2, (c) 3:4, (d) 1:1, (e) 20_1:1_S, (f) 25_1:1_S, and (g) 25_1:2_S_MWCNT scaffolds. Scale bars represent 10 lm.

with the 1:10, 1:2, 3:4, and 1:1 scaffolds with tissue culture plastic (TCP) used as a control (Fig. 3). Rat muscle cells grown on the scaffolds on days 7, 14, and 21 had significantly lower metabolic activity than the cells cultured on TCP at every time point. On day 28, rat muscle cells cul-

tured on 1:10 and 1:2 scaffolds had significantly higher metabolic activity than all scaffold groups on days 7, 14, and 21. Day 7 scaffolds, day 14 scaffolds except for the 1:2 value, and day 21 scaffolds except for the 1:10 value had significantly lower metabolic activity than the 1:1 scaffolds

TABLE II. Average Fiber Diameter With Standard Deviation and Range for Each of the Scaffolds Scaffold 1:10 1:2 3:4 1:1 20_1:1_S 25_1:1_S 25_1:2_S_MWCNT

Average Fiber Diameter (lm)

Fiber Diameter Range (lm)

Fiber Angle Standard Deviation ( )

2.78 6 1.49 2.35 6 1.22 2.36 6 1.67 2.37 6 1.73 1.22 6 0.80a 1.47 6 0.76b 2.19 6 1.31

0.955–6.151 0.901–5.135 0.654–7.709 0.728–8.121 0.441–3.794 0.626–3.543 0.522–5.215

15.7 22.7 42.5 37.1 27.0 45.4 56.7

Data were considered statistically significant when p < 0.05. a Significance from all other scaffolds except the 251-1_S scaffold. b Significance from the 1 to 10 scaffold.

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FIGURE 2. Histograms of the fiber orientation for the (a) 1:10, (b) 1:2, (c) 3:4, (d) 1:1, (e) 20_1:1_S, (f) 25_1:1_S, and (g) 25_1:2_S_MWCNT scaffold. Each bar represents a 10 span and normal curves are overlaid on each.

on day 28. In addition, the day 21 TCP value, which was the largest during the study, was significantly larger than all other scaffolds at all time points. It was interesting to note that metabolic activity on the 3:4 scaffolds remained significantly low throughout the entire study. In order to confirm the presence of rat muscle cells on the scaffolds, cytoskeletal actin and nuclei were fluorescently stained and imaged. Figure 4 shows images of the fluorescently stained rat muscle cells for the 1:10, 1:2,

3:4, and 1:1 scaffolds from day 28. All four of the scaffolds supported cellular attachment; however, the fewest cells were observed on the 3:4 scaffolds. The smaller number of attached rat muscle cells explains the low metabolic activity recorded with the MTS assay (Fig. 3). The 1:10 scaffolds had smaller patches of cells compared to the 1:2 and 1:1 scaffolds where the cells attached on the majority of the scaffold. In addition, varying amounts of cell fusion with actin interaction were observed on all

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TABLE III. Average Conductivity Values for Each of the Scaffolds (p < 0.05) Scaffold

Conductivity (S/cm)

1:10 1:2 3:4 1:1 20_1:1_S 25_1:1_S 25_1:2_S_MWCNT

0.020 6 0.009 0.040 6 0.014 0.053 6 0.012a 0.067 6 0.015b 0.031 6 0.006 0.040 6 0.001 0.035 6 0.004

a

Significance from the 1:10 scaffold. Significance from all other scaffolds except the 3–4 scaffold.

b

four of the scaffolds with the 3:4 scaffold having the smallest amount. DISCUSSION

Several different concentrations of PEDOT Nps were combined with PCL and/or MWCNT then electrospun to form conductive scaffolds for skeletal muscle tissue engineering. The use of sonication was explored as a means of improving scaffold fiber alignment through evenly dispersing the PEDOT Np in solution. In addition, adding a second conductivity element, MWCNT, to the solution was investigated as a means of improving fiber alignment while further enhancing scaffold conductivity. These scaffolds were characterized for their fiber diameter and alignment, conductivity, mechanical properties, and in vitro cytocompatibility. The material properties of the different types of scaffolds were compared, and the feasibility and effectiveness of using PEDOT Nps in an electrospun scaffold for skeletal muscle regeneration was evaluated. While electrospinning the various PEDOT solutions, there was some difficulty in obtaining a stable polymer stream, and there is evidence that the PEDOT Np may be negatively affecting the electrospinning process.23–25 There are a few theories that may explain why the PEDOT-PCL solutions were unable to successfully electrospin. The first is

FIGURE 3. Rat primary skeletal muscle cell metabolic activity on the four PCL-PEDOT scaffolds and TCP as a control (n 5 4). Each sample was diluted 1:4 with dIH2O prior to absorbance measurement at 490 nm. Averages with standard deviations are shown (p < 0.05). The symbol “*” indicates significance from all TCP time-points, 1:10_day28, and 1:2_day28. The symbol “#” indicates significance from 1:1_day28. The symbol “%” indicates significance from day 21 TCP.

that the PEDOT Np are attracted to one another and aggregate (observed within the manufacturer’s bottle). The second is that the partial positive areas of PEDOT are reacting with the applied voltage and disrupting the electrospinning process.28,29,32 Finally, the addition of the PEDOT increases the conductivity of the overall solution, which may increase

TABLE IV. Tensile Elastic Modulus and Yield Stress Values for the Seven Scaffold Groups (p < 0.05)

1:10 1:2 3:4 1:1 20_1:1_S 25_1:1_S 25_1:2_S_MWCNT Rat soleus muscle Pig medial muscle

Elastic Modulus (kPa)

Yield Stress (kPa)

754.9 6 212.7a 500.9 6 77.8 290.9 6 26.1 435.5 6 107.7 837.3 6 198.4b 598.4 6 148.3 741.1 6 334.7e 48.11 6 16.1923 151.83 6 50.3023

837.0 6 176.5a 571.3 6 105.3 323.3 6 51.8 469.4 6 133.8 1402.7 6 351.5c 702.4 6 145.8d 905.4 6 327.6e 17.84 6 5.8323 54.00 6 21.1723

The tensile elastic moduli and yield stresses of muscle tissue from Sprague-Dawley rats and pigs were previous reported.23 a Significance from the 3:4 and 1:1 scaffolds. b Significance from the 1:2, 3:4, and 1:1 scaffolds. c Significance from all other scaffolds. d Significance from the 3:4 scaffold. e Significance the 3:4 and 1:1 scaffolds.

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FIGURE 4. Fluorescently stained images showing actin (green) and nuclei (blue) on day 28 for the (a) 1:10, (b) 1:2, (c) 3:4, and (d) 1:1 scaffolds. Scale bars represent 100 lm. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]

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ORIGINAL ARTICLE

its attraction to the charged target. This may cause the solution to leave the syringe faster, leading to spatter, inconsistent fiber morphology, and a more random fiber orientation. The average fiber diameters for the scaffolds from the nonsonicated solutions were between 2 and 3 lm, but sonication of the PEDOT Np before the addition of PCL caused the fiber diameter to decrease and display greater uniformity (smaller standard deviations). The addition of the MWCNT to the 1:2 scaffolds had no effect on fiber diameter. Feng et al.33 electrospun several 1:10, 3:10, and 5:10 PEDOT:poly(lactide-co-glycolide) (PLGA) solutions that had average fiber diameters of 10, 3.7, and 2.9 lm, respectively, which are larger than the values we measured. Interestingly, they found that as the amount of PEDOT increased, the fiber diameter decreased; however, we did not observe this effect. Kiristi et al. combined PVA and PEDOT with either unmodified chitosan or plasma-modified chitosan. The fiber diameter values recorded by Kiristi et al.34 are much smaller than ours, ranging from 170 to 246 nm. We believe that the PEDOT partial positive areas affect the electrospinning process and hence, the fiber diameter. Scaffolds made from the sonicated solutions had a lower conductivity, which led to smaller diameter fibers. Increasing concentrations of PEDOT caused fiber alignment to decrease among the 1:10, 1:2, 3:4, and 1:1 scaffolds (Fig. 2 and Table II). This may be caused by an increase in the repulsion between fibers with higher concentrations of PEDOT Np due to the higher concentrations of partial positive charges on the Np. The partial positive charges may have also reacted with the positive voltage used in the electrospinning set-up, leading to a larger variation in fiber angle.28,29,31 Another theory for the lack of alignment concerns the use of polymers with enhanced conductivity. As discussed earlier, solutions with increased conductivity may be more attracted to the charged target and move to the target faster. This increase in velocity may not allow them to be drawn into an aligned state by the speed of the mandrel. If the mandrel speed were increased to account for the faster moving polymer, the amount of alignment, and possibly fiber diameter, may have been increased. Sonication of the solution was utilized to alleviate the fiber alignment issue, but it did not appear to have a large effect on fiber orientation. However, lowering the concentration of PCL from 25% to 20% in the 25_1:1_S scaffolds to the 20_1:1_S scaffolds decreased the standard deviation of the fiber angle from 45.4 to 27.0 . Since decreasing the concentration of PCL increases the relative concentration of PEDOT, it would seem that this contradicts our earlier finding that increasing the concentration of PEDOT causes decreased fiber alignment. Perhaps the increased viscosity of the 25% solution caused a less uniform distribution of the nanoparticles than in the lower viscosity 20% solution. Further investigation into these results is needed. The addition of MWCNT to the PEDOT solutions caused the largest observed variation in fiber orientation. An unfavorable interaction between the MWCNT Np and PEDOT Np may be occurring that affected the results, but more experimental study is needed to determine this effect.

Although fiber alignment decreased with increasing concentration of PEDOT Np mixed with the PCL, electrospun PCL mixed with other components has not shown the same issue. For example, Choi et al.35 electrospun aligned PCL/ collagen nanofiber meshes and quantified fiber orientation using histograms as well. They utilized mandrel speeds up to 2350 RPM, which produced a highly aligned scaffold with a similar distribution to the 1–10 scaffolds from this study; although, no numerical standard deviation was reported. Thus, the addition of collagen to PCL did not have the same effect on fiber orientation that increasing PEDOT Np amounts did. It may be possible that conductive elements affect the electrospinning process; however, Chen et al.36 electrospun a mixture of PCL and conductive polyaniline with parallel block magnets to create highly aligned scaffolds. They were able to achieve scaffolds with [mt]90% of the fibers within 10 of the average fiber direction, which is a higher degree of alignment than any solution from this study. Also, the addition of PANi to the PCL did not disrupt the electrospinning process or affect the fiber diameter. Thus, the use of magnets may help increase the degree of alignment in our scaffolds. Finally, the standard deviation of native muscle fiber alignment has been found to be 3–7 , which is smaller than the standard deviation of the fiber alignment in the scaffold with the lowest concentration of PEDOT Np.37 Conductivity of the scaffolds increased with the concentration of PEDOT Np. However, sonication of the PEDOT Np prior to electrospinning caused the conductivity of the resulting scaffold to decrease. Dispersing the PEDOT particles within the PCL may have lowered the local concentration of PEDOT below the percolation threshold necessary for optimal conductivity. It was also found that the addition of MWCNT did not significantly increase the conductivity of the 1:2 scaffolds, but there was less variation between the measured values. Since MWCNT are acid functionalized with a carboxylic group (–COOH) and the PEDOT is partially positive, there may be some attraction occurring between the – COO2 groups and the PEDOT as well. In addition, the distribution of PEDOT Np within the fibers was likely changed as a result of the sonication process. In the future, transmission electron microscopy may be used to determine the distribution of PEDOT Np within the fibers from sonicated and nonsonicated solutions. Wang et al. fabricated PEDOT nanotubes alone or with lead telluride (PbTe) added as a semiconductor. The PEDOT nanotubes became dedoped and had a low conductivity value, 0.64 3 1023 S/cm, while the PEDOT-PbTe composite had a much larger value, 6.16 3 1023 S/cm.38 Both these values are smaller than those measured in this study; however, our samples most likely contained many more nanoparticles than the samples measured by Wang et al. In contrast, Granstr€ om and Ingan€as fabricated 10 nm PEDOT fibers and found the conductivity to be 780 S/cm.36 It is likely that as more PEDOT is in a continuous phase, there is an increase in the conductivity of the overall scaffold. This may explain why the Np in our PCL-PEDOT scaffolds had much lower conductivity. It may be necessary to increase

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the amount of PEDOT Np in order to achieve a higher conductivity to form a more continuous phase, but this may increase the scaffold mechanical properties and affect fiber alignment. The elastic moduli and yield stresses measured for the seven scaffold types are much closer to the values for both rat and pig skeletal muscle than our previously measured reverse IPMC scaffolds.23–25 However, some of these lower mechanical properties are most likely due to the random alignment of fibers within the scaffold. One of the sonicated scaffolds, 20_1:1_S, showed a significant increase in elastic modulus and yield stress while the 25_1:1_S scaffold was only significantly larger for yield stress. This significant difference could be due to the increase in fiber alignment, a better distribution of PEDOT, and/or the decrease in PCL as PEDOT is a much stiffer polymer. For example, Tahk et al.39 calculated a theoretical Young’s modulus value of 3.79 6 0.45 GPa for PEDOT which further supports the observed increase in mechanical properties. Jin et al. formed PEDOT coated electrospun PVA fibers and measured the tensile properties. The elastic moduli ranged from 27.3 MPa to 80.1 MPa while the yield stress ranged from 6.5 MPa to 11.7 MPa.40 All of these values are much greater than those measured for the PEDOT containing scaffolds in this study. Since the sonication process decreased the conductivity and the addition of MWCNT did not significantly increase the conductivity, only the 1:10, 1:2, 3:4, and 1:1 PCL-PEDOT scaffolds were tested for cytocompatibility with rat primary skeletal muscle cells. PEDOT has been known to have poor cytocompatibility which is why we used small amounts with the PCL solution.40 Over the 4-week period, metabolic activity increased on all four scaffolds; however, the metabolic activity of the rat muscle cells on the 3:4 scaffolds was significantly lower on day 28 than the other groups on day 28 (Fig. 3). Although metabolic activity was higher in the TCP group initially, the metabolic activity in most scaffold groups reached the level of the TCP group by day 28. Both the MTS assay and fluorescent images confirm that the PCL-PEDOT scaffolds are cytocompatible with rat skeletal muscle cells in vitro. These assays serve as preliminary tests which show the potential of this scaffold to be biocompatible with rat skeletal muscle tissue. Jin et al. cultured human cancer stem cells (hCSCs) on electrospun PVA-PEDOT core-shell fibers. The hCSCs cultured on the PVA-PEDOT fibers displayed similar morphology and attachment as those cultured on TCP.40 Feng et al.33 cultured dorsal root ganglia (DRG) on PEDOT microfibers and showed that the neurites followed the PEDOT fiber alignment and exhibited the classic spindle-shaped morphology. Since our PCL-PEDOT scaffolds were determined to be cytocompatible and Feng et al. showed that neurite growth was possible, it is possible that myoblasts and neurons seeded onto a PEDOT containing scaffold have potential to regenerate the neuromuscular junction.

CONCLUSION

Four PCL-PEDOT solutions with various concentrations of PEDOT were electrospun to form fibrous scaffolds and char-

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acterized for skeletal muscle regeneration. Conductivity was found to increase as the concentration of PEDOT increased. In addition, scaffold mechanical properties decreased with increasing PEDOT amounts; however, as PEDOT has a higher modulus than PCL, this is most likely due to a decrease in fiber alignment. Sonication of the PEDOT particles was performed prior to electrospinning to determine the effect of particle dispersion on fiber diameter and alignment. Although more consistent fiber diameters were obtained, only the 20_1:1_S scaffold displayed better alignment. The sonication process also resulted in decreased conductivity for both scaffolds and increased the mechanical properties. The addition of MWCNT to the PEDOT solutions did not significantly increase conductivity, failed to improve fiber alignment, and increased the elastic modulus and yield stress. Nevertheless, skeletal muscle cells attached, displayed measurable metabolic activity, and displayed an elongated morphology on all scaffold groups. Thus, our hypothesis that the addition of PEDOT Np to PCL scaffolds would be cytocompatible, increase scaffold conductivity, and keep elastic modulus in the range of skeletal muscle was confirmed. However, the low degree of fiber alignment needs to be addressed in order to make a more suitable skeletal muscle scaffold. Either a different fabrication method or a conductive material other than PEDOT may be needed to form the necessary conductive, aligned fibers while maintaining mechanical properties similar to native skeletal muscle in order to successfully regenerate this tissue. REFERENCES  M, Valero-Cabre  A. Neural plasticity after periph1. Navarro X, Vivo eral nerve injury and regeneration. Prog Neurobiol 2007;82: 163–201. 2. Shear TD, Martyn JAJ. Physiology and biology of neuromuscular transmission in health and disease. J Crit Care 2009;24:5–10. 3. Campbell WW. Evaluation and management of peripheral nerve injury. Clin Neurophysiol 2008;119:1951–1965.  A, Navarro X. Regeneration and func4. Rodrıguez FJ, Valero-Cabre tional recovery following peripheral nerve injury. Drug Discovery Today 2004;1:177–185. € ker S, Wieland M, Dro € ger G, Kirschning A, 5. Bruns S, Stark Y, Ro Stahl F, Kasper C, Scheper T. Collagen biomaterial doped with colominic acid for cell culture applications with regard to peripheral nerve repair. J Biotechnol 2007;131:335–345. 6. Bian W, Bursac N. Engineered skeletal muscle tissue networks with controllable architecture. Biomaterials 2009;30:1401–1412. 7. Liao H, Zhou GQ. Development and progress of engineering of skeletal muscle tissue. Tissue Eng Part B Rev 2009;15:319–331. 8. Machingal MA, Corona BT, Walters TJ, Kesireddy V, Koval CN, Dannahower A, Zhao W, Yoo JJ, Christ GJ. A tissue-engineered muscle repair construct for functional restoration of an irrecoverable muscle injury in a murine model. Tissue Eng Part A 2011;17: 2291–2303. 9. Bach AD, Beier JP, Stern-Staeter J, Horch RE. Skeletal muscle tissue engineering. J Cell Mol Med 2004;8:413–422. 10. Clark ME, Bair MJ, Buckenmaier CC, Gironda RJ, Walker RL. Pain and combat injuries in soldiers returning from Operations Enduring Freedom and Iraqi Freedom: Implications for research and practice. J Rehabil Res Dev 2007;44:179–193. 11. Maquet V, Martin D, Malgrange B, Franzen R, Schoenen J,  ro ^ me R. Peripheral nerve regeneration using bioreMoonen G, Je sorbable macroporous polylactide scaffolds. J Biomed Mater Res 2000;52:639–651. 12. Dodla MC, Bellamkonda RV. Differences between the effect of anisotropic and isotropic laminin and nerve growth factor presenting

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