Dexamethasone encapsulated coaxial electrospun PCL/PEO hollow microfibers for inflammation regulation

Abstract

One of the major challenges in the field of biomaterials is to develop strategies to regulate the innate host inflammatory response. Coaxial electrospun PCL/PEO hollow fibers containing dexamethasone were evaluated for a local delivery of dexamethasone to reduce adverse inflammation responses often resulting from biomaterial implantation. Core–shell PCL/PEO hollow fibers with a highly porous surface were successfully developed using coaxial electrospinning. The release of dexamethasone exhibited a burst release during the first 0.5–3 h followed by a sustained release over more than 12 days. In vitro study showed that the dexamethasone encapsulated coaxial PCL/PEO hollow fibers significantly reduced cell proliferation of LPS-stimulated macrophages and meanwhile significantly decreased the mRNA expression levels of the pro-inflammatory cytokines TNF-α and IL-1β. The dexamethasone encapsulated coaxial PCL/PEO hollow microfibers are promising biomaterials for implantation to combat the acute inflammation responses which take place during first 24–48 h after biomaterial implantation and meanwhile to regulate possible hostile chronic inflammations, thus increasing the efficacy and longevity of the implants in vivo.

---

Introduction

One of the major challenges in the field of biomaterials is to develop strategies to combat the innate host inflammatory response.¹ Though a mild innate immune response is part of the healing process, implantation of synthetic biomaterials usually initiates a foreign body response (FBR), which is characterized by macrophage infiltration, foreign giant cell formation and fibrotic encapsulation of the implant leading to chronic pain and device rejection and failure.^1,2 Glucocorticoids, such as dexamethasone (Dex), are well-known anti-inflammatory drugs with the ability to minimize the FBR at the implant site. Since long-term systemic use of this drugs leads to unwanted side-effects, localized and sustained release delivery of anti-inflammatory drugs has been investigated.³ Therefore, development of biomaterials loaded with anti-inflammatory drugs for a local and sustained release offers an attractive alternative to diminish the foreign body reaction and therefore to overcome the limited in vivo functionality and longevity of implantable devices. Several studies have been carried on the design of biomaterials containing dexamethasone to modulate the host innate response. Dex delivery constructs using PLGA microspheres,^4,5 nanofiber gels,¹ porous scaffolds produced by gas-foaming/salt-leaching methods⁶ or electrospun fibers composed of poly(L-lactic)acid and poly(ε-caprolactone)⁷ have been developed and the effects of Dex released has been evaluated either in vitro and in vivo.

Electrospun fibers have recently attracted intense attention in the field of tissue engineering because of both its resemblance to the extracellular matrix and its ease of add-on functionality for providing cells with both nano-topographical elements and biological cues.^8–12 Drugs can be easily embedded in the fiber through one-step dissolution or dispersion in the polymer solution for electrospinning. Coaxial electrospinning, a relatively new technique of fabricating core–shell fibers have added versatility to this technique by adding an ideal additional barrier surrounding the encapsulated drugs (reservoir system) allowing tuning the release kinetics, as opposed to randomly distributing the drug throughout the fiber matrix (matrix system).^13–16 Furthermore, the coaxial electrospinning enables co-delivery of two bioactive molecules encapsulated separately at the core and the shell to allow both spatial gradients and temporal release of biochemical signals, to mimic the complex bio-signalling systems in tissue microenvironment.^17,18

PCL and PEO are biocompatible polymers that have been widely used in tissue engineering and drug delivery. PEO is, furthermore, known to suppress protein adhesion¹⁹ and applied as anti-inflammatory polymeric coatings for implantable biomaterials and devices.²⁰ The interesting combination of biocompatible, biodegradable PCL with protein-resistant PEO is thus promising in regulating inflammation responses.²¹ In the present study we present coaxial PCL/PEO hollow fibers for a local and sustained anti-inflammatory drug release with promising further use in regenerative medicine. Dex release was monitored overtime. Biocompatibility of coaxial PCL/PEO hollow fibers on RAW 264.7 cells was evaluated by quantification of the LDH activity. Bioactivity of dexamethasone released from coaxial PCL/PEO hollow fibers was evaluated in vitro by determination of macrophage cell proliferation and by quantification of expression of inflammation related genes using real-time RT-PCR.

Results and discussion

One of the major challenges of biomaterial implantation is to combat the body reaction from innate inflammatory response. In order to modulate the host immune responses, several studies have focused on the design of biomaterials containing anti-inflammatory drugs. Coaxial electrospinning fibers has been recently explored for delivery of biomolecules for tissue engineering purposes due to its promise for preserving drug activity and the subsequently controlled delivery. Furthermore, electrospun fibers mimic the fibrillar structure of extracellular matrix (ECM) offers topographical cues for cells attachment that may enhance the drug efficacy.²² Here, Dex was encapsulated into the core of the PCL/PEO hollow fibers by one-step coaxial electrospinning, and the inflammation regulation was evaluated with LPS-stimulated macrophages.

Drug encapsulation

In order to ensure the Dex released from coaxial electrospun fibers exhibit anti-inflammatory properties, our first question was addressed to determine the appropriate Dex amount range to be encapsulated into the fibers. Both safety and anti-inflammatory properties of different dosages of Dex (10⁻⁴ M, 10⁻⁶ M and 10⁻⁸ M of Dex) on LPS-stimulated macrophages cells culture on tissue culture plastic (TCP) was evaluated by our research group (Fig. S1†). All Dex dosages tested were safe for the cells; among them, Dex at 1 × 10⁻⁴ M and at 1 × 10⁻⁶ M showed similarly a higher down-regulation of several inflammation related marker genes (Fig. S2†). Coaxial PCL/PEO hollow fibers were loaded with 0.13 w/w% Dex, which with a 100% release would result to an equal amount of 6.2 × 10⁻⁵ M added in media. Therefore, Dex at 6.2 × 10⁻⁵ M was selected to carry out this study.

Fig. 1 represents the SEM images of the fibers fabricated by one-step coaxial electrospinning. All the fibers were of hollow structure, and the cylindrical shapes of the fibers were well kept without any collapse (Fig. 1a and d). The diameter of the PCL/PEO hollow fibers was 8.255 ± 0.838 μm before, and 8.7 ± 0.813 after encapsulation with Dex, and the thickness of the wall was around 1.6 μm. Highly porous and rough structures were found on the fiber surfaces (Fig. 1c and f). The morphology of the hollow fibers was well preserved and no obvious change could be observed after loaded with 0.13% w/w dexamethasone in the PEO core solution.

Fig. 1 Microscopic structure of Coaxial PCL/PEO hollow fibers. (a), (b), (c) are the SEM images of the electrospun hollow fibers: (a) cross section of the mat, insert a′: the cross section of single fiber; (b) and (c) top-view of the fibers with different magnification. (d), (e), (f) are the SEM images of hollow fibers with 0.13% w/w dexamethasone loaded in the PEO core solution: (d) cross section of the mat, insert d′: the cross section of single fiber; (e) and (f) top-view of the fibers with different magnification.

Drug release

In agreement with the previously reported electrospun fibrous drug release system,^7,23–25 the developed coaxial PCL/PEO fibers exhibited a typical two-stage release pattern (Fig. 2b and c). In the initial stage, a burst 34% release of dexamethasone was observed during the first 0.5 h of incubation. Further 19% of Dex was slowly released up to 3 h, followed by a sustained cumulative release of 14% of Dex over time up to 12 days.

Fig. 2 (a) Diffusion of liquid solution of fluorescein in the coaxial PCL/PEO hollow fibers. Cumulative release profile (b) and data points (c) of dexamethasone from PCL/PEO fibers up to 12 days. Values represent mean ± SD.

The release profile is correlated with the interesting morphology of the coaxial PCL/PEO hollow fibers, where the mechanism of hollow fiber formation has been discussed.²⁶ However, the shell composition has not been investigated. Here Nile red and BSA-FITC were loaded into the PCL and the PEO solution, respectively, prior to electrospinning and the resulted fibers were imaged under laser scanning confocal microscopy. As seen in Fig. S1,† clearly PCL/Nile red was observed laying out around PEO/BSA-FITC (Fig. S1a–c†). However, there was also overlapping between them, (Fig. S1d–f†) suggesting the entanglement of PCL and PEO during electrospinning. Considering the small molecular size, and solubility of Dex in both PCL and PEO solutions, during electrospinning Dex could be distributed either at the PEO inner layer, at the PCL/PEO interfaces or at the PCL outer layer. Due to the hydrophilicity of PEO, when the fibers were immersed into the aqueous buffer, the liquid could quickly wet the inner wall of the hollow fibers through any cutting openings (Fig. 2a). The Dex on the inner PEO surfaces released immediately, which resulted in the burst release in the first 0.5 hour. Since the PEO has a large molecular weight of M_w 900 000, Dex encapsulated inside PEO would be released during the swelling of PEO.²⁵ Therefore, between 0.5 h and 3 h, there was still a fast release of Dex of 19%. On the other hand, SEM analysis of the microstructure of PCL/PEO fibers, both without and with Dex, revealed a highly porous surface with a pore diameter of around 600 nm at the outer layer of the nanofibers, which might be aroused by the fast evaporation rate of DCM (boiling point 40 °C) in the shell PCL solution and also the high humidity during electrospinning process serving as a coagulation bath for PCL precipitation.^27,28 Considering PCL is hydrophobic and slowly degradable, Dex entangled with PCL was released gradually by diffusion through these pores. A sustained release was observed up to 12 days, while 32% of Dex was still not released.

Biocompatibility

In order to determine the effect of coaxial PCL/PEO hollow fibers on cell viability, the LDH activity present in the culture media was determined after 26 h of cell culture. As seen in Fig. 3, the adverse effects of LPS decreased when cells were treated with Dex, showing a significant decrease in the LDH activity in LPS-stimulate cells treated with Dex directly in the culture media (PCL/PEO + mDex) compared to LPS-stimulated cells without Dex treatment (PCL/PEO).

Fig. 3 LDH activity measured from culture media collected 26 hours after exposure of LPS-induced RAW 264.7 cells to coaxial PCL/PEO hollow fibers. High control (100% cytotoxicity) was cell culture media from cells cultured on TCP and treated with 1% Triton X-100. Low control (0% cytotoxicity) was cell culture media from cells seeded on TCP. Values represent the mean ± SEM. Differences between groups were assessed by Student t-test: p < 0.05 (*) versus positive control of inflammation.

Cell proliferation

Once a sustained release of Dex from coaxial electrospun fibers was confirmed, we next evaluated the biological response of LPS-stimulated macrophages cells to the fibers. MTS assay was used to evaluate cell proliferation (Fig. 4). While a significant increase of 24% was observed in cell proliferation from day 1 to day 3 for the cells seeded on coaxial PCL/PEO hollow fibers, no proliferation were observed when Dex was added in the medium on coaxial PCL/PEO hollow fibers or Dex was loaded in coaxial PCL/PEO hollow fibers. In addition, a significant decrease in macrophage cell proliferation was observed either after day 1 (p = 0.041) and day 3 (p = 0.01) when the cells were seeded on coaxial PCL/PEO hollow fibers containing Dex, compared to coaxial PCL/PEO hollow fibers without Dex. Similar anti-proliferation effects from Dex were also observed when Dex was added in the media on coaxial PCL/PEO hollow fibers.

Fig. 4 Proliferation of LPS-induced RAW 264.7 cells after 1 and 3 days of culture on coaxial PCL/PEO hollow fibers. LPS-induced RAW 264.7 cells cultured on coaxial PCL/PEO hollow fibers without Dex (named as PCL/PEO) and treated with Dex added in the culture media (named as PCL/PEO + mDex) served as positive and as negative control groups of inflammation, respectively. Cell proliferation on fibers was measured by MTS assay. Data were expressed as percentage of positive control cells at day 1, which was set to 100%. Values represent mean ± SEM. Differences between groups were assessed by Student t-test: p ≤ 0.05 (*) versus positive control. Gene expression of cell adhesion and cell proliferation markers.

RAW 264.7 cells cultured on coaxial PCL/PEO fibers containing Dex exhibited a reduced cell proliferation, which was in accordance to the role of glucocorticoids on reducing the release of inflammatory mediated cells at the injured sites. In agreement, several studies have reported that glucocorticoids decreases proliferation of lymphocytes,⁶ fibroblasts and macrophages.^29,30 In addition, no differences were observed between coaxial PCL/PEO hollow-Dex and Dex added directly in the culture media, suggesting that even though a 38% of the theoretical total amount of Dex loaded was still not released at the end of the in vitro study (day 3), a significant reduction in cell growth of inflammated cells was observed.

Further, the results from LDH assay (Fig. 3) ruled out the possibility that the decrease in cell proliferation could be due to a toxic effect of Dex. In fact the treatment with Dex decreased the adverse effects of LPS.

Gene expression

Glucocorticoids accomplish its anti-inflammatory effects mainly by modulating the formation and secretion of inflammatory mediators and cytokines to an inflammatory stimulus.³¹ Thus we further investigated whether the amount of Dex released from the PCL/PEO fibers would be enough to mitigate the inflammatory response of RAW 264.7 to LPS.

Fig. 5 shows expression of TNF-α and IL-1β, two hallmark genes of inflammation. As shown in this figure, compared to PCL/PEO fibers, PCL/PEO-Dex fibers decreased significantly the expression of IL-1β, both after day 1 and day 3, and of TNF-α only after day 3. After 1 day of culture, down-regulation of TNF-α was significantly lower on coaxial PCL/PEO hollow-Dex fibers, compared to PCL/PEO fibers treated with Dex in the media. PCL/PEO-Dex fibers decreased 22.1% in the expression of TNF-α compared to coaxial PCL/PEO hollow fibers, although the data did not reach to be statistically significant (p = 0.135).

Fig. 5 Expression of inflammation related genes in LPS-stimulated RAW 264.7 cells cultured for 1 and 3 days onto coaxial PCL/PEO hollow fibers without (named as PCL/PEO) and with Dex added in the media (named as PCL/PEO + mDex). Data represent relative mRNA levels of target genes normalized with reference genes, expressed as a percentage of positive control of inflammation (named as PCL/PEO), which were set to 100%. Values represent the mean ± SEM. Differences between groups were assessed by Student t-test or Mann Whitney test depending on their normal distribution: p < 0.05 (*) versus positive control (named as PCL/PEO), (#) versus negative control (named as PCL/PEO + mDex).

These results are consistent with the Dex release profile, where only 59.3% and a 61% of Dex was released after day 1 and day 3 of culture, respectively. Differences on the effect of Dex on TNF-α and IL-1β at day 1 might be due to differences in the temporal sequence of expression of these markers. In addition, it has been reported that the effect of glucocorticoids depends on the interaction of glucocorticoid diffused through the cell membrane to its cytosolic receptor which subsequently undergoes the repression of pro-inflammatory cytokines (TNF-α, IL1β, IL6) and several inflammatory mediators (iNOS, ICAM-1, COX-2).^31–33 Therefore, the results from this study suggest that: (i) encapsulation process through a coaxial electrospinning do not alter the structure and function of Dex, thus might allow an effective binding of Dex to its receptor; and (ii) the amount of Dex released was enough to exhibit its anti-inflammatory effects. Taking all together, these results suggest that even though only 62% of Dex was released from the coaxial PCL/PEO hollow fibers after 3 days of culture; it was active and sufficient to suppress the inflammatory response. In addition, it should be pointed out that despite its beneficial anti-inflammatory properties, Dex has very potent side effects including immunosuppression or hypertension.³³ Thus the observed anti-inflammatory effectiveness of locally released Dex suggests that the PCL/PEO fibers could be used as a suitable local delivery system with a minimized side effects associated to glucocorticoids administration.

Conclusions

In conclusion, our results demonstrate that the coaxial PCL/PEO hollow fibers are suitable for local and controlled release of dexamethasone to reduce growth of inflammation-related cells and decrease the expression of pro-inflammatory cytokines TNF-α and IL-1β. These coaxial polymeric fibers may be used in temporal drug delivery and tissue engineering applications. This two-stage drug release offers an attractive drug delivery system to combat the acute inflammation response which take place during first 24–48 h after biomaterial implantation while the long-term, slow release of drug would be able to regulate possible hostile chronicle inflammations, thus increasing the efficacy and longevity of the implants in vivo.

Experimental

Electrospinning

The hollow PCL/PEO fibers were fabricated by the coaxial electrospinning process described by Dror et al.²⁶ The polymer solution for shell was prepared by dissolving 12% w/w PCL (M_w = 70 000–90 000, Sigma-Aldrich) in 80/20 (v/v) DCM/DMF, and the polymer solution for core was prepared by dissolving 4% w/w PEO (M_w = 900 000, Sigma-Aldrich) in 40/60 (v/v) EtOH/H₂O. For the hollow fibers loaded with Dex (named as PCL/PEO-Dex), Dex was dissolved in the core solution of PEO at 0.13% w/w of the total amount of PCL and PEO in the hollow fibers. All the solutions were prepared at room temperature and stirred until homogeneous solutions formed. The electrospinning was performed with a homemade coaxial needle. Both the inner and outer syringes were fixed horizontally on the syringe pumps, and a high voltage power supply was applied between the coaxial needle and a grounded rotary drum collector. The electrospinning process was carried out under the following conditions: applied voltage 12.5 kV, feeding rate of shell solution 4 mL h⁻¹, feeding rate of core solution 0.5 mL h⁻¹, and distance between the tip of needle and collector 19 cm. The experiments were carried out at room temperature, and the relative humidity was 37–45%. The obtained fibers were dried under vacuum (Labconco Freezone TriadTM) overnight to remove the excess solvents before further use. The fibers were then punched out at 12 mm Ø (∼1.1 cm²).

Characterization of fibers morphology by scanning electron microscopy

The morphologies of the electrospun fibers were examined with a high-resolution scanning electron microscopy (SEM) (FEI, Nova 600 NanoSEM). The fibers were placed directly into the SEM chamber without any metal sputtering or coating. All the images were captured with an acceleration voltage of 5 kV and using secondary electrons detector.

Dexamethasone release from coaxial PCL/PEO hollow-Dex fibers

To obtain the release profile of dexamethasone from hollow fibers, hollow fibers containing 0.13% w/w Dex in the PEO core layer were placed on the bottom of a sterile standard 48-well plate. The wells with fibers were then filled with 400 μL of warm phosphate buffer saline (PBS). In order to mimic cell culture conditions, samples were maintained at 37 °C in a humidified atmosphere for up to 12 days. At designed time points (0.5, 1, 2, 3, 5, 7, 24, 48, 72, 144, and 288 hours), 40 μL solutions was collected and refilled with an equal amount of warmed fresh PBS buffer. To determine the amount of Dex released, absorbance at 242 nm were measured by using a UV-vis spectrophotometer (UV-1800, Shimazu). The experiment was performed in sixplicates (n = 6). Relative absorbance units were correlated with the amount of dexamethasone released using a linear standard curve. The percentage of the cumulative drug released was then calculated based on the initial weight of the dexamethasone incorporated in the electrospun scaffold.

Diffusion through coaxial fibers

A drop of aqueous solution of fluorescein (0.1 mM) was added on the top of the coaxial fiber mesh and images under fluorescence microscope were taken every 5 seconds.

Cell culture of RAW 264.7 cells

The murine RAW 264.7 macrophage-like cell line (ATCC, Manassas, USA) was selected as in vitro model of inflammation. Cells were routinely cultured in DMEM-Glutamax (4.5 g L-D-glucose, (−) pyruvate) (GIBCO, Grand Island, NY) supplemented with 10% FBS (Biowhittaker, Walkersville, MD) and antibiotics (100 IU and 100 IU streptomycin antibiotics) (Gibco, Grand Island, NY) at 37 °C in a humidified atmosphere of 5% CO₂. Cells were routinely subcultured 1 : 10 before reaching confluence by scrapping. All experiments were performed at passage 17.

To test anti-inflammatory properties, RAW 264.7 cells were seeded onto coaxial PCL/PEO hollow fibers containing Dex (named as PCL/PEO-Dex) at a density of 1.6 × 10⁵ cells per cm and subsequently treated with 0.1 μg mL⁻¹ lipopolysaccharide (LPS, E. Coli 055:B5, Schnelldorf, Germany) to induce cell inflammation. In parallel, LPS-stimulated macrophages cells seeded on coaxial PCL/PEO hollow fibers without (named as PCL/PEO) or with treatment of Dex (6.2 × 10⁻⁵ M) directly added in the culture media (named as PCL/PEO + mDex) served as a positive and negative control of cell inflammation, respectively. Cells were maintained in standard cell culture conditions (37 °C in a humidified atmosphere of 5% CO₂) for 1 and 3 days.

Lactate dehydrogenase activity

Lactate dehydrogenase (LDH) activity in the culture media was used as an index of cell death. The activity of the cytosolic enzyme was determined spectrophotometrically after 30 min incubation at 25 °C of 50 μL of culture and 50 μL of the reaction mixture, by measuring the rate of oxidation of NADH at 490 nm in the presence of pyruvate, according to the manufacturer's kit instructions (Roche Diagnostics, Mannheim, Germany). Results were presented relative to the LDH activity in the medium of cells cultured in growing media (low control, 0% of cell death) and of cells treated with Triton X-100 1% (high control, 100% of death), using the equation:

Cytotoxicity (%) = ((exp. value − low control)/(high control − low control)) × 100.

MTS assay

CellTiter 96® Aqueous One Solution Cell Proliferation Assay (G3580) (Promega) was used to evaluate the cell metabolic activity of cells cultured on the different fibers for 1 and 3 days. Briefly, at specific time points, cell culture media was taken out and refreshed by 100 μL media into each well. Subsequently, a mixture of 20 μL MTS reagent and 280 μL media was added into each well, and incubated for 1 hour at 37 °C in a humidified atmosphere (5% CO₂). The formazan product, which was bioreduced from MTS by living cells, was read at an absorbance of 490 nm using a spectrophotometer (VictorX5 multi-lable plate reader, 2030-0050, PerkinElmer, Inc).

Total RNA isolation and TNF-alpha gene expression by real-time RT-PCR

Total RNA was isolated using Trizol® (Invitrogen Life Technologies, Carlsbad, CA, USA), according to the manufacturer's protocol. Total RNA was quantified at 260 nm using a Nanodrop spectrophotometer (IMPLE AH Diagnostics, Helsinki, Finland). The same amount of RNA (400 ng) was reverse transcribed to cDNA using High Capacity RNA-to-cDNA kit (Applied Biosystems, Foster City, CA), according to the protocol of the supplier. Aliquots of each cDNA were frozen (−20 °C) until the PCR reactions were carried out.

Real-time PCR was performed in the Lightcycler 480® (Roche Diagnostics, Mannheim, Germany) using SYBR green detection. Real time PCR was done for three reference genes (18S rRNA and glyceraldehyde-3-phosphate dehydrogenase (GAPDH) and TBP) and 2 target genes (tumor necrosis factor alpha (TNF-α) and interleukin 1β (IL-1β)). The primer sequences are detailed in Table 1.

Table 1 Sequence of inflammation markers related genes

Gene	Primer sequence
18S	S 5′-GTAACCCGTTGAACCCCATT-3′
	A 5′-CCATCCAATCGGTAGTAGCG-3′
GAPDH	S 5′-ACCCAGAAGACTGTG-GATGG-3′
	A 5′-CACATTGGG-GGTAGGAACAC-3′
TBP	S 5′-AGAGAGCCACGGACAACTG-3′
	A 5′-ACTCTAGCATATTTTCTTGCTGCT-3′
TNF-α	S 5′-GTAGCCCACGTCGTAGCAAAC-3′
	A 5′-ATCGGCTGGCACCACTAGTT-3′
IL-1β	S 5′-GCCACCTTTTGACAGTGATGA-3′
	A 5′-GATGTGCTGCTGCGAGATTT-3′

---

Each reaction contained 7 μL Lightcycler 480 SYBR GREEN I Master (containing Fast Start Taq polymerase, reaction buffer, dNTPs mix, SYBR Green I dye and MgCl₂), 0.5 μM of each, the sense and the antisense specific primers and 3 μL of the cDNA dilution in a final volume of 10 μL. The amplification program consisted of a pre-incubation step for denaturation of the template cDNA (10 min 95 °C), followed by 45 cycles consisting of a denaturation step (10 s 95 °C), an annealing step (10 s 60 °C) and an extension step (10 s 72 °C). After each cycle, fluorescence was measured at 72 °C (λ_ex 470 nm, λ_em 530 nm). A negative control without cDNA template was run in each assay.

Real-time efficiencies were calculated from the given slopes in the LightCycler 480 software using serial dilutions, showing all the investigated transcripts high real-time PCR efficiency rates, and high linearity when different concentrations are used. PCR products were subjected to a melting curve analysis on the LightCycler and subsequently 2% agarose/TBE gel electrophoresis to confirm amplification specificity, T_m and amplicon size, respectively. Relative quantification after PCR was calculated by dividing the concentration of the target gene in each sample by the mean of the concentration of the three reference genes (housekeeping genes) in the same sample using the advanced relative quantification method provided by the LightCycler 480 analysis software version 1.5 (Roche Diagnostics, Mannheim, Germany).

Statistics

All data are presented as mean values ± standard error of the mean (SEM). To assume parametric or non-parametric distribution for the normality test, a Kolmogorov–Smirnov test was done; differences between groups were assessed by Student t-test or Mann-Whiney test depending on their normal distribution. The SPSS® program for Windows (Chicago, IL), version 17.0 was used. Results were considered statistically significant at the p-values < 0.05.

Acknowledgements

We gratefully acknowledge the Danish Council for Strategic Research for the funding to the ElectroMed Project at the iNANO Center, and the Lundbeck Foundation and the Carlsberg Foundation for their financial support. Marina Rubert and Yan-Fang Li contributed equally.

Notes and references

1. M. J. Webber, J. B. Matson, V. K. Tamboli and S. I. Stupp, Controlled release of dexamethasone from peptide nanofiber gels to modulate inflammatory response, Biomaterials, 2012, 33(28), 6823–6832.
2. J. M. Anderson, Biological responses to materials, Annu. Rev. Mater. Res., 2001, 31, 81–110.
3. J. Morais, F. Papadimitrakopoulos and D. Burgess, Biomaterials/tissue interactions: possible solutions to overcome foreign body response, AAPS J., 2010, 12(2), 188–196.
4. G. J. Dawes, L. E. Fratila-Apachitei, B. S. Necula, I. Apachitei, G. J. Witkamp and J. Duszczyk, Release of PLGA-encapsulated dexamethasone from microsphere loaded porous surfaces, J. Mater. Sci.: Mater. Med., 2010, 21(1), 215–221.
5. T. Hickey, D. Kreutzer, D. J. Burgess and F. Moussy, Dexamethasone/PLGA microspheres for continuous delivery of an anti-inflammatory drug for implantable medical devices, Biomaterials, 2002, 23(7), 1649–1656.
6. J. J. Yoon, J. H. Kim and T. G. Park, Dexamethasone-releasing biodegradable polymer scaffolds fabricated by a gas-foaming/salt-leaching method, Biomaterials, 2003, 24(13), 2323–2329.
7. N. M. Vacanti, H. Cheng, P. S. Hill, G. JoDT, T. T. Dang and M. Ma, et al., Localized delivery of dexamethasone from electrospun fibers reduces the foreign body response, Biomacromolecules, 2012, 13(10), 3031–3038.
8. P. D. Dalton, C. Vaquette, B. L. Farrugia, T. R. Dargaville, T. D. Brown and D. W. Hutmacher, Electrospinning and additive manufacturing: converging technologies, Biomater. Sci., 2013, 1(2), 171–185.
9. A. K. Ghag, J. E. Gough and S. Downes, The osteoblast and osteoclast responses to phosphonic acid containing poly(ε-caprolactone) electrospun scaffolds, Biomater. Sci., 2014, 2(2), 233–241.
10. D. Gugutkov, J. Gustavsson, M. P. Ginebra and G. Altankov, Fibrinogen nanofibers for guiding endothelial cell behavior, Biomater. Sci., 2013, 1(10), 1065–1073.
11. R. Murugan and S. Ramakrishna, Design strategies of tissue engineering scaffolds with controlled fiber orientation, Tissue Eng., 2007, 13(8), 1845–1866.
12. P.-O. Rujitanaroj, R. Aid-Launais, S. Y. Chew and C. Le Visage, Polysaccharide electrospun fibers with sulfated poly(fucose) promote endothelial cell migration and VEGF-mediated angiogenesis, Biomater. Sci., 2014, 2(6), 843–852.
13. H. Jiang, Y. Hu, Y. Li, P. Zhao, K. Zhu and W. Chen, A facile technique to prepare biodegradable coaxial electrospun nanofibers for controlled release of bioactive agents, J. Controlled Release, 2005, 105(2–3), 237–243.
14. A. K. Moghe and B. S. Gupta, Co axial electrospinning for nanofiber structures: preparation and applications, Polymer Reviews, 2008, 48(2), 353–377.
15. A. Saraf, L. S. Baggett, R. M. Raphael, F. K. Kasper and A. G. Mikos, Regulated non-viral gene delivery from coaxial electrospun fiber mesh scaffolds, J. Controlled Release, 2010, 143(1), 95–103.
16. S. Srouji, D. Ben-David, R. Lotan, E. Livne, R. Avrahami and E. Zussman, Slow-release human recombinant bone morphogenetic protein-2 embedded within electrospun scaffolds for regeneration of bone defect: in vitro and in vivo evaluation, Tissue Eng., Part A, 2010, 17(3–4), 269–277.
17. W. M. Saltzman and W. L. Olbricht, Building drug delivery into tissue engineering design, Nat. Rev. Drug Discovery, 2002, 1(3), 177–186.
18. Z. Man, L. Yin, Z. Shao, X. Zhang, X. Hu and J. Zhu, et al., The effects of co-delivery of BMSC-affinity peptide and rhTGF-beta 1 from coaxial electrospun scaffolds on chondrogenic differentiation, Biomaterials, 2014, 35(9), 5250–5260.
19. A. Worz, B. Berchtold, K. Moosmann, O. Prucker and J. Ruhe, Protein-resistant polymer surfaces, J. Mater. Chem., 2012, 22(37), 19547–19561.
20. A. W. Bridges and A. J. García, Anti-inflammatory polymeric coatings for implantable biomaterials and devices, J. Diabetes Sci. Technol., 2008, 2(6), 984–994.
21. Y. F. Li, M. Rubert, H. Aslan, Y. Yu, K. A. Howard and M. Dong, et al., Ultraporous interweaving electrospun microfibers from PCL-PEO binary blends and their inflammatory responses, Nanoscale, 2014, 6(6), 3392–3402.
22. W. Ji, Y. Sun, F. Yang, J. J. van den Beucken, M. Fan and Z. Chen, et al., Bioactive electrospun scaffolds delivering growth factors and genes for tissue engineering applications, Pharm. Res., 2010, 28(6), 1259–1272.
23. W. Cui, X. Li, X. Zhu, G. Yu, S. Zhou and J. Weng, Investigation of drug release and matrix degradation of electrospun poly(DL-lactide) fibers with paracetanol inoculation, Biomacromolecules, 2006, 7(5), 1623–1629.
24. K. Kim, Y. K. Luu, C. Chang, D. Fang, B. S. Hsiao and B. Chu, et al., Incorporation and controlled release of a hydrophilic antibiotic using poly(lactide-co-glycolide)-based electrospun nanofibrous scaffolds, J. Controlled Release, 2004, 98(1), 47–56.
25. Y. Z. Zhang, X. Wang, Y. Feng, J. Li, C. T. Lim and S. Ramakrishna, Coaxial electrospinning of (fluorescein isothiocyanate-conjugated bovine serum albumin)-encapsulated poly(ε-caprolactone) nanofibers for sustained release, Biomacromolecules, 2006, 7(4), 1049–1057.
26. Y. Dror, W. Salalha, R. Avrahami, E. Zussman, A. L. Yarin and R. Dersch, et al., One-step production of polymeric microtubes by co-electrospinning, Small, 2007, 3(6), 1064–1073.
27. M. Bognitzki, W. Czado, T. Frese, A. Schaper, M. Hellwig and M. Steinhart, et al., Nanostructured fibers via electrospinning, Adv. Mater., 2001, 13(1), 70–72.
28. C. L. Casper, J. S. Stephens, N. G. Tassi, D. B. Chase and J. F. Rabolt, Controlling surface morphology of electrospun polystyrene fibers: effect of humidity and molecular weight in the electrospinning process, Macromolecules, 2003, 37(2), 573–578.
29. Y. Nakatani, T. Amano and H. Takeda, Corticosterone suppresses the proliferation of RAW 264.7 macrophage cells via glucocorticoid, but not mineralocorticoid, receptor, Biol. Pharm. Bull., 2013, 36(4), 592–601.
30. X. Wang, Y. Chen, Y. Wang, X. Zhu, Y. Ma and S. Zhang, et al., Role of RHOB in the antiproliferative effect of glucocorticoid receptor on macrophage RAW 264.7 cells, J. Endocrinol., 2009, 200(1), 35–43.
31. K. A. Smoak and J. A. Cidlowski, Mechanisms of glucocorticoid receptor signaling during inflammation, Mech. Ageing Dev., 2004, 125(10–11), 697–706.
32. W. Y. Almawi and O. K. Melemedjian, Molecular mechanisms of glucocorticoid antiproliferative effects: antagonism of transcription factor activity by glucocorticoid receptor, J. Leukocyte Biol., 2002, 71(1), 9–15.
33. T. Ito, I. P. Fraser, Y. Yeo, C. B. Highley, E. Bellas and D. S. Kohane, Anti-inflammatory function of an in situ cross-linkable conjugate hydrogel of hyaluronic acid and dexamethasone, Biomaterials, 2007, 28(10), 1778–1786.

---

Footnotes

† Electronic supplementary information (ESI) available: LDH assay and expression of inflammation related genes under the treatments of different concentration of Dex. See DOI: 10.1039/c4ra08358j

‡ These authors contributed equally to this work.

---