4 Development of 3D Wet-spun Polymeric Scaffolds Loaded with Antimicrobial Agents for Bone
Engineering
Dario Puppi1, Dinuccio Dinucci1, Cristina Bartoli1, Carlos Mota1, Chiara Migone1, Francesca Dini2, Giovanni Barsotti2, Fabio Carlucci2 and Federica Chiellini1
1 Laboratory of Bioactive Polymeric Materials for Biomedical and Environmental Applications (BIOLab), via Vecchia Livornese 1291, 56010 San Piero a Grado (Pi). Department of Chemistry and Industrial Chemistry, University of Pisa, Italy.
2 Department of Veterinary Clinic, University of Pisa, Via Livornese 56010 S. Piero a Grado (PI) Italy.
Abstract
Three-dimensional wet-spun microfibrous meshes of a star poly(ε-caprolactone) (*PCL) were developed as potential scaffolds endowed with antimicrobial activity. The in vitro release kinetics of the meshes, under physiological conditions, was initially fast and then a sustained controlled release for more than one month was observed. Cell cultures of a murine preosteoblast cell line showed good cell viability and adhesion on the wet-spun *PCL fibre scaffolds. These promising results indicate a potential application of the developed meshes as engineered bone scaffolds with antimicrobial activity.
Keywords: scaffolds, tissue engineering, drug release, wet-spinning, star polymer, microfibrous meshes
4.1 Introduction
One of the most common concepts underlying the engineering of biological tissues is to combine a
biodegradable matrix, commonly referred to as a scaffold, with cells and/or biologically active
molecules, to produce a living construct to promote repair and regeneration of tissue. The scaffold
guides the development of the required tissue by supporting cell colonization, migration, growth
and differentiation, and acting as vehicle for the delivery of bioactive agents [Puppi et al., 2010a;
Stock and Vacanti, 2001]. Several different polymeric materials and processing techniques have been proposed for the development of tissue engineering (TE) scaffolds with a wide range of macro-shapes and micro/nanostructures [Puppi et al., 2010a]. The recent development of different processing methods for the production of fibrous meshes has made it possible to generate scaffolds with variable fibres diameter and assembly, as well as, pore size, shape and distribution [Domingos et al., 2009; Gomes et al., 2006b; Gomes et al., 2003; Puppi et al., 2010c; Puppi et al., 2011;
Tuzlakoglu et al., 2004; Tuzlakoglu et al., 2005]. Meshes of micro- nano-sized fibres exhibit high surface area to volume ratio and high and interconnected porosity that favour cell adhesion and migration, as well as, mass transport governing the exchange of cell nutrients and wastes, the material degradation rate and the release kinetics of loaded bioactive factors. Electrospun ultrafine fibre meshes for TE are being pursued based on their ability to resemble the highly porous nanosized structure of native extracellular matrices. However, the small pore size together with the high fibre packing density limits cell infiltration into these meshes [Heydarkhan-Hagvall et al., 2008; Leong et al., 2010; Zhang et al., 2005]. Studies have shown that 3D structures composed of an assembly of randomly oriented fibres with diameters in the range of hundreds of microns, produced by melt spinning followed by fibre bonding or wet-spinning, provided good in vitro cell adhesion and proliferation with good penetration of cells into the inner part of the scaffold [Oliveira et al., 2007; Tuzlakoglu et al., 2010]. This type of scaffold is used in bioreactor cell cultures as they provide large, organized cell communities throughout the tissue engineered construct [Gomes et al., 2006a; Gomes et al., 2006b; Gomes et al., 2003]. These highly interconnected porous microfibre meshes allow the development of hydrodynamic microenvironments closely resembling natural interstitial fluid conditions in vivo with minimal diffusion constraints [Gomes et al., 2003].
Wet-spinning is a non-solvent induced phase inversion technique that is used to form continuous polymer microfibre by solidifying the polymer solution as it is injected into a coagulation bath composed of a poor solvent or a non-solvent/solvent mixture with respect to the polymer. A number of wet-spun fibres have been made of natural polymers, such as chitosan [Tuzlakoglu et al., 2004; Tuzlakoglu and Reis, 2008] and regenerated silk fibroin [Lee et al., 2007; Um et al., 2004;
Yao et al., 2001], that cannot be formed by conventional spinning techniques. Wet-spun natural-
material fibres, such as chitosan [Denkbas et al., 2000] and poly(L-lactic acid) (PLLA) fibres [Gao
et al., 2007], are also being studied as carriers for drug release. Polymer meshes composed of
poly(ε-caprolactone) fibres [Puppi et al., 2010a; Williamson and Coombes, 2004], braided
PLLA/chitosan fibres [Zhang et al., 2007] or starch-based fibres [Tuzlakoglu et al., 2010] are being
studied as scaffolds for TE.
The employment of antibiotic-loaded implants, such as hydroxyapatite bone cement [Buranapanitkit et al., 2004], poly(L-lactide-co-glycolide) (PLGA)/bioactive glass composite matrix [Makinen et al., 2005], and PLGA microspheres [Ambrose et al., 2004], is being explored for the treatment of bone diseases caused by microbial infections to improve the efficacy of antibiotic therapy in the poorly perfused bone bed. In addition, biodegradable TE scaffolds with antimicrobial activity have been investigated for one-stage surgical treatment of chronic osteomyelitis [Baro et al., 2002; Castro et al., 2003; Koort et al., 2005; Makinen et al., 2005].
Besides providing high local concentrations of bactericidal agents for prolonged periods, the scaffolds, at the same time, fill in the space caused by surgical debridement of the dead bone, as well as, guide and promote tissue regeneration at the defect site.
Fluoroquinolones are synthetic antimicrobial drugs relatively easy to synthesize and modify, that are extensively used in human and veterinary medicine as a broad spectrum of antibacterial action against Gram-positive, Gram-negative and atypical bacteria [Drilica and Hooper, 2003; Rubinstein, 2001]. Their antimicrobial spectrum includes the bacterial pathogens frequently encountered in bone and joints microbial infections as well as for treating osteomyelitis and other bone infections [Darley and MacGowan, 2004; Greenberg et al., 2000]. Enrofloxacin (EF) was the first antimicrobial fluoroquinolone used in veterinary medicine for infections in the urogenital and respiratory tracts [de Lucas et al., 2008]. In vitro, both EF and its metabolite ciprofloxacin are active against osteomyelitis pathogens [Lautzenhiser et al., 2001; Meinen et al., 1995].
Levofloxacin (LF), a third generation fluoroquinolone, has a broad spectrum of antimicrobial activity; in clinical trials, it effectively treated infections of the respiratory and genitourinary tract, skin and skin structures [Croom and Goa, 2003]. LF is also effective in the treatment of osteomyelitis [Greenberg et al., 2000; Lazzarini et al., 2004].
Star polymers are linear polymer chains attached to a smaller central moiety. Because of their small size, spherical structure and limited interaction between molecules, star polymers have different properties compared to the linear polymers with equivalent molecular weight. They are less crystalline and have lower melting and solution viscosity but are easier to process and degrade faster [Burchard, 1999; Celik et al., 2009; McKee et al., 2005; McKee et al., 2004].
In this study, 3D meshes were produced by wet-spinning and evaluated as scaffolds for the
regeneration of bone tissue in the presence of infections. A three-arm branched poly(ε-
caprolactone) (*PCL), for TE scaffolds, was electrospun [Puppi et al., 2010b]. The wet-spinning
parameters for 3D meshes loaded with either EF or LF, employed as fluoroquinolone models, were
investigated and the scaffolds characterized in terms of loading efficiency, in vitro release kinetics,
and compatibility towards the MC-3T3 preosteoblast cell line.
4.2 Materials and Methods 4.2.1 Materials
Three-arm star branched poly(ε-caprolactone) (Mw = 189,000 g/mol, Mn = 80,000 g/mol, PDI = 2.36, *PCL
189) was supplied by Prof. Ramani Narayan (Michigan Biotechnology Institute, Lansing, MI, USA). Enrofloxacin (EF), Levofloxacin (LF) and all the solvents were purchased from Sigma Aldrich (Italy) and used as received.
4.2.2 Fabrication of wet-spun meshes
The *PCL (15 g) was dissolved in 100 ml of acetone at 40°C under gentle stirring for 2 h to form a homogeneous solution. To produce the composite fibres, 0.1 and 0.2% w/v EF (or LF) were dissolved, with stirring for 1 h, in the above polymer solution.
Meshes were wet-spun by loading the polymer solution into a syringe equipped with a pump (NE- 1000, New Era Pump Systems, NY) to control the feed rate (F = 2.25 ml·h
-1). The solution was forced through the syringe 0.4 mm inner diameter needle into an ethanol coagulation bath. The spun meshes were left in the coagulation bath for 24 h, then removed, placed in a vacuum chamber for 48 h and stored in a desiccator. Three replicates of each developed wet-spun mesh were prepared.
4.2.3 Morphology analysis
Samples were analysed by scanning electron microscopy (SEM, model JEOL JSM 300, Japan) under backscattered electron imaging. The average fibre diameter was determined using ImageJ 1.36b software on micrographs with a 50X magnification. The fibre diameters were calculated from over 30 measurements per specimen taken randomly from selected fields.
4.2.4 Drug loading
Drug loading was determined by analysing the antibiotic content in the coagulation bath and in the meshes. The coagulation bath was vacuum-desiccated while the mesh was dissolved in acetone and vacuum-desiccated. A 10 ml of the HPLC mobile phase solution was added to each of the mesh samples and stirred for 1 h. The drug content in the suspensions was measured by reverse-phase HPLC coupled to an ultraviolet (UV) detector, with the mobile phase (49.95% water, 49.95%
acetonitrile and 0.1% trifluoroacetic acid) at a flow rate of 1 ml·min
-1. The LF was quantified at
295 nm (standard curve 0.39-50 µg·ml
-1, R2=0.9997) and EF at 278 nm (standard curve 0.39-50 µg·ml
-1, R2=0.9996). All analyses were performed in triplicate.
4.2.5 In vitro drug release
Cylindrical samples (16 mm in diameter and 15 mm in height) weighing around 320 mg, were tested. The samples were placed in 10 ml of PBS (pH 7.4 at 37°C) and rotationally agitated (100 rpm). At fixed times, 1 ml of release medium was analysed in replicate by HPLC, as described above, and replaced with fresh PBS.
4.2.6 In vitro biological evaluation
Cell culture onto *PCL scaffolds: Mouse calvaria-derived preosteoblast cells MC3T3-E1, obtained from the American Type Culture Collection (ATCC CRL-2593), were cultured as monolayers in Alpha Minimum Essential Medium with ribonucleosides, deoxyribonucleosides, 2 mM L- glutamine, 1 mM sodium pyruvate, supplemented with 10% fetal bovine serum and 100 U/ ml:100 µg/ml penicillin: streptomycin (GIBCO, Invitrogen Corporation, Italy) at 37 °C in a 5% CO
2enriched atmosphere.
Samples from *PCL scaffolds were sterilized in a 70% ethanol/water solution for 24 h, washed with PBS (0.01 M, pH 7.4) and exposed to UV light for 40 min. Cells (1 x 10
4) were seeded directly onto each sample in 12 well plates and allowed to proliferate. Growth medium was changed every 48 h. At days 7 and 14, samples were processed for microscopy and viability analysis. Experiments were performed in triplicate.
4.2.6.1 WST-1 cell proliferation assay
Cell viability and proliferation were measured by using the WST-1 assay (Roche Molecular Biochemicals), which is based on the conversion of tetrazolium salt WST-1 into soluble formazan.
After 7 days of culture, WST-1 reagent (diluted 1:10) was added to the cultures and incubated for 4 h at 37 °C. The formazan dye absorbance was measured at 450 nm, with the reference wavelength at 620 nm with a Microplate Reader (Biorad).
4.2.6.2 Cytochemical stain
At specific culture endpoints, the mesh scaffolds were rinsed with PBS and fixed with 4%
paraformaldehyde for 45 min. Each step of the sample processing was carried out at room
temperature and under mild stirring on an orbital plate to enhance uniform assess of the fixative and
other solutions throughout the mesh samples. Samples were rinsed with PBS and incubated with 1% w/v toluidine-blue for 30 min, then washed with PBS for microscopic analysis.
4.2.6.3 Confocal laser scanning microscopy (CLSM) of the cell morphology
Cell culture morphology and proliferative distribution on scaffolds were investigated by Confocal Laser Scanning Microscopy (CLSM). After paraformaldehyde fixation, the samples were washed with PBS and incubated with a 2% PBS solution of Triton X-100 (Sigma-Aldrich, Italy) for 15 min to permeabilize cells. PBS solutions of 4’-6’-diamidino-2-phenylindole (DAPI; Invitrogen, Italy) and phalloidin-Alexa488 (Invitrogen, Italy) were added and incubated for 1 hour. After dye incubation, the samples were washed with PBS and scaffolds were cut so as to observe the cross- sections. A Nikon Eclipse TE2000 inverted microscope equipped with a EZ-C1 confocal laser and D-Eclipse C1-si spectral optical system (Nikon, Japan) was used to analyse the samples. A 405 nm laser diode (405 nm emission) and Argon ion laser (488 nm emission) were used to excite DAPI and phalloidin fluorophores, respectively. The images were captured with Nikon EZ-C1 software using identical settings for each sample. The merged images were processed with ACT-2U software (Nikon, Japan).
4.2.6.4 Live/Dead assay
Cell culture viability was assessed by the Live/Dead assay dyeing (Invitrogen, Italy). Scaffold samples were washed with PBS to remove extracellular esterase activity prior to incubating with a calcein-AM 2uM and EthD-1 4uM PBS solution for 30 min at room temperature following the producer’s protocol. The incubating solution was added to each well to completely cover the scaffolds. The excited scaffold fluorophores at 488 and 514nm, respectively, were observed by CLSM.
4.2.7 Statistical analysis
Quantitative data were presented as the mean ± standard deviation (SD). Data sets consisting of
average fibre diameter or drug loading efficiency were screened by one-way ANOVA and a Tukey
test was used for post hoc analysis; significance was defined at p<0,05.
4.3 Results and Discussion
Polymer scaffolds that are able to release antibiotics in a controlled manner represent an innovative therapeutic strategy for the regeneration of bone defects in the presence of infections, such as osteomyelitis. In this study, the fabrication of 3D wet-spun *PCL scaffolds loaded individually with EF or LF were investigated and optimised. The polymeric meshes loaded with fluoroquinolones were characterized for their morphology, actual drug loading and in vitro drug release kinetics. In addition, the cytocompatibility of *PCL meshes was evaluated using the MC3T3-E1 murine preosteoblast cell line by means of biochemical assays and confocal laser scanning microscopy analysis.
4.3.1 Meshes production and morphological analysis
By applying the wet-spinning conditions optimised in previous work (15% w/v *PCL concentration, acetone as solvent and ethanol as the non-solvent) [Puppi et al., 2010a], 3D polymer meshes with a non-woven fibre structure were produced by random movement of the coagulation tank during the injection process (Figure 1a). To produce the polymer meshes loaded with either EF or LF, the polymer and the antibiotic (0.1 or 0.2% w/v) were processed as described above. By adding 2.5 ml of polymer solution to a cylindrical coagulation tank, cylindrical scaffolds, 16 mm in diameter and 15 mm in height were obtained (Figure 1b).
Figure 1 - a) Scheme of wet-spinning process for the production of 3D polymer scaffolds. A polymer solution filament is extruded directly into a coagulation bath, leading to the formation of a continuous polymer fibre which is randomly collected into a non-woven, porous mesh; b) 3D mesh of *PCL obtained by wet-spinning.
The SEM analysis, using backscattering electron imaging, revealed that the meshes had a non- woven fibrous structure (Figures 2a-i) with fibre diameters in the range 100-300 µm (Table 1) and a porous fibre surface with pore sizes of few micrometres. The addition of the antibiotics to the polymer solution affected the final morphology of the wet-spun fibres. By comparing the high magnification micrographs of fibre surface (Figures 2c, 2f and 2i), an increase in the pore size was observed; this effect was more evident in the EF-loaded fibres. Statistical analysis of the fibre diameters (Table 1) revealed significant differences between the loaded and unloaded fibres.
Figure 2 - Representative EBS micrographs of fibrous structure (a, d, g), single fibre (b, e, h) and fibre surface (c, f, i) in *PCL-based meshes. Unloaded *PCL mesh (a-c); *PCL mesh loaded with EF (d-f); *PCL mesh loaded with LF (g-i).
Table 1 - Fibre dimension analysis.
Mesh Type1 Fibre diameter [µm]
*PCL 145.0 ± 23.0
EF-loaded *PCL 227.4 ± 53.3
LF-loaded *PCL 219.2 ± 55.1
In the wet-spinning process, once the filament of polymer solution was injected into the coagulation bath, the non-solvent diffused into the solution, whereas the solvent diffused into the bath. This counter-diffusion lowered the polymer solubility and induced a phase separation into a polymer rich and a polymer-poor phase. A dense, nonporous layer was usually formed at the solution/coagulant interface due to the instantaneous non-solvent diffusion into the polymer solution [Tsay and Mchugh, 1990, 1992]. In the case of delayed liquid-liquid demixing, a porous surface interfaced with the coagulant was formed [Barton et al., 1997; Wienk et al., 1996]. The presence of additives in the polymer solution affected both the thermodynamics and kinetics of the phase inversion process, which influenced the morphology of the fibres [Kim and Lee, 1998;
Mulder, 1996; Reuvers and Smolders, 1987; Strathmann et al., 1975]. It seems that the added fluoroquinolones to the polymer solution affected the polymer salvation, and consequently, the coagulation process. This would explain the increase in the fibre diameter and pore size of fluroquinolone-loaded meshes.
A variety of processing methods have been proposed to knit or physical bond the fibres prefabricated by wet-, dry- or melt-spinning to form a fibrous mesh. Previously, we proposed a single-step method for the production of *PCL non-woven structures by wet-spinning through a continuous random motion of the coagulation tank [Puppi et al., 2010a]. In this study, non-woven fibre assemblies were produced by adopting the same method and increasing the processing time.
Accordingly, 3D cylindrical scaffolds with dimensions suitable for the regeneration of large-size bone defects were obtained. However, the continuous random motion of the tank was not automatically driven, and the production of a 15 mm thick scaffold required about fifty minutes of continuous processing by a dedicated operator. Thus, the final fibrous texture depended on the ability of the operator. Current research is addressing the production efficiency and the reproducibility of the morphology by using a computer assisted automated motion system to the coagulation tank.
4.3.2 Drug loading
The polymer meshes were loaded with either EF or LF, produced with two different concentrations
(0.1% and 0.2% w/v) of the antibiotic in the polymer solution and analysed. Most of the antibiotic
was found in the coagulation bath; the amount loaded into the fibres was below 30% depending on
the type of antibiotic and the concentration of the solution (Table 2). The antibiotic in meshes
increased from 18.2 to 23.1% w/v in the case of EF and 22.5 to 26.3% in the case of LF.
Wet-spinning is a cost effective and versatile method to fabricate microfibrous polymeric structures loaded with bioactive agents; it does not require high temperatures or exposure to harsh solvents that would compromise the activity of bioactive agents [Crow et al., 2005]. However, the possibility of producing wet-spun fibres loaded with bioactive agents has only been reported in a few studies [Chang et al., 2008; Crow et al., 2005; Crow and Nelson, 2006; Denkbas et al., 2000;
Gao et al., 2007; Puppi et al., 2010a].
Table 2 - Drug loading efficiency.
Antibiotic type and concentration1 Loading efficiency [%]
EF (0.1% w/v) 18.2 ± 3.1
EF (0.2% w/v) 23.1 ± 1.7
LF (0.1% w/v) 22.5 ± 5.6
LF (0.2% w/v) 26.3 ± 7.4
1 Data are expressed as mean ± standard deviation.
In the present work, the relatively low loading efficiency is likely due to the low solubility of the two antibiotics in ethanol and/or to the slow solidification process during the fibre formation. Most of the antibiotic added to the polymer solution was able to pass from the solution phase to the coagulant, as confirmed by the analysis of the coagulation bath. Consequently, it seems that the drug loading is related to the choice of the solvent/non-solvent system used [Puppi et al., 2010a].
However, it should be considered that the solvents employed (acetone and ethanol) are classified by Food and Drug Administration as “Solvents in class 3 with Low Toxic Potential”. The difference in loading efficiency between EF- and LF-loaded meshes could also be related to the solidification process. The SEM analysis of EF-loaded fibres revealed a porous surface that could favour the mass transfer governing the fluoroquinolone solubilization in the non-solvent, thus impeding the drug loading.
4.3.3 In vitro drug release
The drug release of the meshes loaded with either EF or LF (0.2% w/v) was investigated for six
weeks, by placing wet-spun samples in PBS (pH 7.4, 37°C) under rotational shaking. The data
confirmed a cumulative release of the antibiotic loaded into the mesh (based on the loading
to five weeks (Figure 3). The two plots display a similar release profile with a fast release in the first 24 hours, with a release of ~50%. The release rate then gradually decreased, reaching a stable profile after twelve days. The total release was approximately 88% for EF and 79% for LF.
Figure 3 - Cumulative percent release of EF and LF from *PCL meshes during in vitro drug release studies (37°C, PBS 1X, pH 7.4). Error bars corresponding to ± standard deviation values calculated on three replicates for each time point.
The initial fast release, “burst effect”, was due to the dissolution of the drug adsorbed on the fibre outer surface. This was probably enhanced by the porosity of the fibre surface. Burst releases are often regarded as a negative consequence for long-term controlled release which researchers seek to avoid because it may lead near or above the toxic level of the drug [Huang and Brazel, 2001].
However, initial fast release of the antibiotic could potentially be clinically desirable to achieve high antibiotic levels immediately after a surgical treatment [Makinen et al., 2005].
4.3.4 In vitro biocompatibility
The biocompatibility of unloaded wet-spun *PCL meshes and the influence of the scaffold
architecture on cell behaviour were assessed through a preliminary investigation on cell viability,
adhesion and cytoskeleton morphology of the murine preosteoblast MC3T3-E1 cell line, carried out
by means of biochemical and microscopic analysis at days 7 and 14. Cell viability and proliferation
were evaluated by WST-1 as a direct contact toxicity assay as well as cellular viability on *PCL scaffolds. The cell count, assayed at days 7 and 14, increased (Figure 4a), indicating insignificant toxic effects on cellular viability. Quantitative data from WST-1 assay of the *PCL scaffolds (Figure 4b) confirmed an increase in the cell proliferation between the two time endpoints.
Figure 4 - Tetrazolium salt WST-1 assay. a) Direct contact toxicity assay; b) Assay on cell-seeded meshes.
Data are expressed as mean ± standard deviation.
Live/dead assay, a two-colour fluorescence assay (green stain for live cells and red for dead cells), was carried out microscopically. Analyses were performed on both the upper bottom sides, and on random cross-sections of the scaffolds. Viable cells were observed at both time endpoints. At day 7, the ratio between the two fluorescence dyes was consistent with a normal cellular adaptation to the scaffold where some cells failed to attach to the material and eventually died. At day 14, the *PCL constructs showed good cell viability with a high live/dead ratio (Figure 5a). The proliferative trend of the culture corroborated the WST-1 biochemical assay results.
Toluidine blue staining was performed to evaluate the cellular distribution on cultured scaffolds.
However, the mesh architecture of the scaffold exhibited a strong contrast that caused a low microscopic resolution of the cellular component. The microscopic analysis did ascertain that the cells were attached to the outer layer of *PCL fibres and that the cell population increased from day 7 to 14 with an almost continuous assembly of cells covering the surface of the fibres.
Cellular spreading and morphology on the *PCL constructs were evaluated by fluorescent CLSM
analysis (Figure 5c). At day 7, cells had adhered to the fibres and the morphology and distribution
were consistent with a viable status. The cells throughout the scaffold were not uniformly
distributed but clustered in colonies on discrete tracts on the fibres. Most of the cells showed a
spindle-like morphology but many extroflections of the plasma membranes were evident indicating
an ongoing cell spreading and adaptation to the environment. Samples at day 14 showed an
micrographs randomly taken throughout the constructs showed that many of the *PCL fibres were almost completely covered by a layer of cells exhibiting intercellular connections, as indicated by the intense green fluorescence emission of actin filaments from individual cells. In some cases, cellular bridging was observed at fibre-fibre contact points (arrows in Figure 5c) indicating a 3D rearrangement of the culture.
Figure 5 - Biological characterization of *PCL meshes/MC3T3-E1 constructs. a) Live/dead visual viability assay: viable (green fluorescence) and dead (red fluorescence) cells on scaffold. Scale bar corresponding to 200 µm, applicable to both micrographs. b) Cytochemical analysis of the constructs: Toluidine Blue staining shows the presence of cells on scaffold and their localization along fibres. Scale bar corresponding to 100 um, applicable to both micrographs. c) CLSM analysis of the constructs: actin filaments (green) and nuclei (blue) showing cellular morphology and distribution on scaffold fibres. Scale bar corresponding to 200 µm, applicable to both micrographs.
