Combining Electrospun Scaffolds with Electrosprayed Hydrogels

Jul 23, 2008 - Singapore, and Institute of Health and Biomedical Innovation, Queensland ... was coelectrospinning medical grade poly(ε-caprolactone)/...
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Biomacromolecules 2008, 9, 2097–2103

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Combining Electrospun Scaffolds with Electrosprayed Hydrogels Leads to Three-Dimensional Cellularization of Hybrid Constructs Andrew K. Ekaputra,† Glenn D. Prestwich,‡ Simon M. Cool,| and Dietmar W. Hutmacher§,* Graduate Program in Bioengineering, National University of Singapore, Singapore, Department of Medicinal Chemistry, University of Utah, Salt Lake City, Utah, Institute of Molecular and Cell Biology, Singapore, and Institute of Health and Biomedical Innovation, Queensland University of Technology, Brisbane, Queensland, Australia Received May 22, 2008; Revised Manuscript Received June 30, 2008

A common problem in the design of tissue engineered scaffolds using electrospun scaffolds is the poor cellular infiltration into the structure. To tackle this issue, three approaches to scaffold design using electrospinning were investigated: selective leaching of a water-soluble fiber phase (poly ethylene oxide (PEO) or gelatin), the use of micron-sized fibers as the scaffold, and a combination of micron-sized fibers with codeposition of a hyaluronic acid-derivative hydrogel, Heprasil. These designs were achieved by modifying a conventional electrospinning system with two charged capillaries and a rotating mandrel collector. Three types of scaffolds were fabricated: medical grade poly(ε-caprolactone)/collagen (mPCL/Col) cospun with PEO or gelatin, mPCL/Col meshes with micron-sized fibers, and mPCL/Col microfibers cosprayed with Heprasil. All three scaffold types supported attachment and proliferation of human fetal osteoblasts. However, selective leaching only marginally improved cellular infiltration when compared to meshes obtained by conventional electrospinning. Better cell penetration was seen in mPCL/Col microfibers, and this effect was more pronounced when Heprasil regions were present in the structure. Thus, such techniques could be further exploited for the design of cell permeable fibrous meshes for tissue engineering applications.

Introduction Scaffold-based tissue engineering involves the combination of cells, bioactive factors, and structural scaffolding materials to promote repair and regeneration of tissues.1 Although tissue engineered constructs must be tailored specifically to the unique site of injury, certain intrinsic properties are expected from a scaffold. Suitable mechanical properties and nontoxic degradation products are inherent requirements for biomaterials to be used as scaffolds. Furthermore, the scaffold’s biochemical properties and architecture should support cell attachment, migration, growth, and ultimate tissue maturation.2 An open and interconnected pore system within the three-dimensional (3D) structure is critical in ensuring proper nutrient and waste transport, tissue in-growth, vascularization, and eventually integration of the constructs with the host. Electrospinning provides a user-friendly approach to fabricate macro- to nanoscaled fibers that are within the size range of the extracellular matrix (ECM).3 These type of scaffolds are widely being exploited in tissue engineering due to their inherently high porosities and surface area-to-volume ratios as well as a wide variety of topographical features to encourage cellular adhesion, surface migration, and proliferation.4,5 Furthermore, the versatility of electrospinning allows tailoring of the surface chemistry, fabrication of aligned structures, and incorporation of bioactive molecules.6,7 The challenge comes, however, from the difficulty in creating good integration between * To whom correspondence should be addressed. Phone: +61 7 3138 6077. Fax: +61 7 3138 6030. E-mail: [email protected]. † National University of Singapore. ‡ University of Utah. | Institute of Molecular and Cell Biology. § Queensland University of Technology.

the electrospun scaffolds with cells and ECM. More often than not, cells cultured on electrospun fibers remain largely on the surface, that is, lacking structural penetration as often reported.8–13 Although pores in an electrospun fiber mesh are interconnected, densely deposited fibers that result in pores with small dimensions prohibit cell penetration along the structure’s thickness. Such limitations could potentially constrain the development and application of electrospun fibers as tissue engineering scaffolds especially for 3D tissues/organs. It is, therefore, of great importance to formulate a method allowing the fabrication of cell permeable scaffolds using the electrospinning technique. The objective of this study was to design and fabricate scaffolds based on electrospinning, which will allow cellular infiltration into the scaffolds’ architecture. The feasibility of three methods was studied to achieve this goal. The first method was coelectrospinning medical grade poly(ε-caprolactone)/ collagen (mPCL/Col) as the main fiber with water-soluble polymers PEO and gelatin. The rationale behind this was that selective removal of solid material from the mesh would increase void volume in the structure. The second method was electrospinning a micron-sized mPCL/Col (µmPCL/Col). This was done to increase the fiber-to-fiber distance and, hence, increase the pore dimension. The third method used was a codeposition of µmPCL/Col with Heprasil, a synthetic ECM based on chemically modified hyaluronic acid (HA) and chemically modified heparin14 developed for 3D cell culture and tissue engineering.15 This resulted in a hybrid fibrous mPCL/Col mesh with regions rich with the HA hydrogel matrix. HA has shown considerable potential in tissue engineering to facilitate angiogenesis, preservation of cellular phenotype, and controlled release of bioactive molecules.16–18 Inclusion of the glycosami-

10.1021/bm800565u CCC: $40.75  2008 American Chemical Society Published on Web 07/23/2008

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Figure 1. Three different electrospinning setups were used in this study: conventional flat plate collection (A), two-capillary coelectrospinning system (B), and two-capillary electrospinning-electrospraying system (C). In all three systems, mPCL/Col was used as the main fiber material. PEO or gelatin was used as water-soluble fiber material for selective leaching approach (B). Heprasil hydrogel was embedded in the mPCL/Col mesh using a simultaneous electrospraying-electrospinning setup (C). To ensure homogeneous collection of PEO, gelatin, and Heprasil, a rotating mandrel collector was used at a speed of 100 RPM.

noglycan hydrogel created pockets of enzymatically degradable matrix within the dense assembly of fibers into which cells could migrate.

Materials and Methods Materials. All reagents and chemicals were purchased from Sigma (St. Louis, MO) unless otherwise stated. Cell culture reagents and supplements were all obtained from Gibco (Grand Island, NY) unless otherwise stated. Medical grade PCL was purchased from Absorbable Polymers International (Birmingham, AL). Medical grade type I bovine collagen was purchased from Symatese Biomateriaux (Chaponost, France). Heprasil was obtained as a kind gift from Glycosan Biosystems (Salt Lake City, UT). A high voltage generator was purchased from Gamma High Voltage Research (Ormond Beach, FL). Syringe pumps for electrospinning were purchased from KD Scientific (Holliston, MA). Fabrication of Scaffolds. Electrospinning solution was made by preparing a 20% w/v solution of mPCL/Col (80/20 by mass) in 1,1,1,3,3,3-fluoro 2-propanol (HFIP). This solution was then electrospun using a 19-G blunt needle connected to a negative 9 kV power supply. Material feed rate and needle-to-mandrel distance of 4.0 mL/hr and 6 cm was used, respectively. This process yielded mPCL/Col fibers with a diameter of 1-2 µm (µmPCL/Col). Heprasil was prepared as per manufacturer’s protocol. Briefly, lyophilized Heprasil and its thiolreactive cross-linking agent, Extralink were reconstituted into their recommended working concentrations and were mixed together to form the hydrogel solution with a 4:1 volume ratio of Heprasil to Extralink. For electrospraying, the solution was further mixed in a 1:1 volume ratio with 3% gelatin solution to modify its viscosity and gelling time. The hydrogel solution was electrosprayed within 30 min after mixing of the components, before extensive cross-linking occurred. Electrospraying was undertaken by connecting a 21-G blunt needle to a positive 7 kV power supply with a hydrogel solution feed rate of 1.6 mL/hr and a distance of 6 cm. The µmPCL/Col-Hep scaffolds were made by simultaneous collection of electrospun mPCL/Col and electrosprayed Heprasil solution on a collection mandrel connected to a ground rotating at a speed of 100 RPM. The collected mesh was stored in a desiccator overnight to let complete evaporation of organic solvent. Figure 1 shows a schematic representation of the three electrospinning setups used. Conventionally, electrospun meshes were made by preparing 12.5% w/v solution of mPCL/Col (80/20 by mass) in HFIP. This solution was then electrospun onto an earthed flat metal collector with positive 10 kV voltage at 0.75 mL/hr feed rate at a 10 cm tip-to-collector distance. Selectively leached mPCL/Col scaffolds were made by coelectrospinning 12.5% w/v mPCL/Col solution with either 20% w/v poly(ethylene oxide) for mPCL/Col-PEO or 10% w/v gelatin for mPCL/Col-Gel in HFIP using a rotating mandrel connected to a ground (100 RPM) for collection. Electrospinning voltage used was negative 9 kV for mPCL/ Col and positive 7 kV for both PEO and gelatin. Tip-to-collector

distance was kept at 6 cm for both capillaries, while the material feed rate was 0.75 mL/hr for mPCL/Col and 1.25 mL/hr for both PEO and gelatin. The collected meshes were kept in a desiccator overnight to allow the complete evaporation of organic solvent. PEO was removed from the mPCL/Col-PEO scaffolds by water immersion for 1 h at room temperature with continuous agitation. Gelatin was left in the mPCL/ Col-Gel structure until cell culture experiments. Scaffold Characterization and Cell Culture. Morphology of the fibrous meshes was examined by gold coating a small sample of the mesh previously cut into 1 × 1 cm2 specimens using a glow discharge sputter-coater (BALTEC, Liechtenstein) followed by viewing under a JEOL 5600 electron microscope (JEOL, Japan) with 10 kV accelerating voltage. Fiber diameter measurements were performed on the electron micrographs using ImageJ, open source image analysis software from the National Institutes of Health (NIH).19 The coelectrospinning processes were confirmed by incorporating 0.05% w/v rhodamine dye into mPCL/Col solution and 0.05% w/v fluorescein isothiocyanate (FITC) dye into PEO solution. The codeposited meshes were cut into 1 × 1 cm2 pieces followed by viewing under the laser scanning confocal microscope (Olympus, Tokyo, Japan). Regions of Heprasil embedded in the scaffolds were viewed by incorporating AlexaFluor 488-labeled bovine serum albumin (AF-BSA; Invitrogen, Carlsbad, CA) into the hydrogel during electrospraying and viewing under the laser scanning confocal microscope. Loading concentrations were set at 0.75 µg/cm2 of mesh produced. The effect of the mandrel diameter used as the collection platform to the efficiency of hydrogel incorporation into the electrospun mesh via electrospinning was studied by fabricating three different meshes using mandrels of 0.8, 1.4, and 1.7 cm in diameter. All three meshes had AF-BSA incorporated into the hydrogel. Other electrospinning parameters were kept constant. The three meshes were cut into 1 × 1 cm2 pieces and were incubated in a 3% bovine testes hyaluronidase solution at 37 °C for 24 h to digest the hydrogel. The supernatant solutions were sampled at the end of the digestion period and fluorescent intensities at 520 nm were read via a FLUOstar Optima spectrophotometer (BMG Labtech, Offenburg, Germany). Collection efficiency was calculated by dividing the spectrophotometer concentration reading by the pre-electrospraying loading and multiplying by 100%. Five samples from each mandrel size were processed and a triplicate reading from each sample was taken. Human fetal osteoblasts (hFOBs; ATCC, Manassas, VA) were used as the model cell line for cellular attachment and infiltration assays. The cells were culture expanded in Dulbecco’s Modified Eagle Media (DMEM) supplemented with 10% fetal calf serum and 1% antibiotics in a humidified, 5% CO2 incubator at 37 °C. For cell seeding, the meshes were cut into discs of 8 mm in diameter using a biopsy punch (B.Braun Medical, Bethlehem, PA). Prior to cell seeding, meshes were sterilized via ethanol immersion and UV irradiation. Each mesh was then placed in a 100 µm cell strainer (BD, Franklin Lakes, NJ) in 60 mm Petri dishes (TPP; Trasadingen, Switzerland). Media conditioning

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Figure 2. Morphology of the electrospun fibrous meshes was observed via scanning electron microscope. Conventional electrospun mesh showed fibers with diameters in the submicron range (A). Coelectrospinning performed using the modified system yielded meshes with two different fiber groups noticeable by their difference in size (B, mPCL/Col-PEO; C, mPCL/Col-Gel). Using the modified electrospinning system with an increased mPCL/Col solution concentration resulted in fibers with micron-sized diameters (D, E). Electrosprayed Heprasil in the structure of mPCL/Col-Hep (E) could not be observed under SEM due to the dehydration process during specimen preparation. The resulting mPCL/Col fiber diameters in all groups are shown graphically in F. Scale bars are 5 µm in A and B; 20 µm in C-E.

was done on all the meshes by pipetting 15 mL of cell culture media into each Petri dish and incubating them overnight in a humidified, 5% CO2 incubator at 37 °C. HFOBs were trypsinized, counted, and seeded onto each disc with a density of 50000 cells per disk. The cell-scaffold constructs were then cultured for a maximum of 10 days in working hFOBs media. On the third and tenth day after seeding, constructs were analyzed for cellular viability, morphology, and infiltration into the structure. Morphology and attachment of cells on the scaffolds was imaged using a scanning electron microscope (SEM). At the predetermined time points of 3 and 10 days, constructs were collected and fixed in 2.5% glutaraldehyde overnight at 4 °C. The samples were then dehydrated by ethanol gradient methods, followed by drying. For electron microscopy, the dried constructs were mounted on viewing stubs and gold sputtered using a glow discharge sputter coater. Viewing was performed under 10 kV accelerating voltage in a JEOL 5600 electron microscope. Viability of the cells was qualitatively assayed using a live-dead assay fluorescein diacetate-propidium iodide (FDA/ PI) staining. At the predetermined time points, constructs were collected and incubated in 2 µg/mL FDA solution for 15 min in a humidified 5% CO2 incubator at 37 °C. Following stringent washings with buffered saline, the samples were then incubated in 50 µg/mL PI solution for 2 min at room temperature. FDA/PI staining was imaged using an Olympus FV500 laser scanning confocal microscope. For the cellular infiltration assay, the constructs were collected at the predetermined time-points and were fixed-frozen in liquid N2 bath. The samples were embedded in cryo-freezing medium (Leica, Wetzlar, Germany), cut into 10 µm sections using a Leica cryo-microtome, and fixed onto poly(L-lysine)-coated microscope slides. The sections were then fixed in 10% formalin for 30 min at room temperature. Visualization of cells in the fibrous mesh was done by staining the fixed sections with hematoxylin and eosin (H&E) stain. Cell nuclei stain bluish-purple, while cytoplasm and ECM stain pink. Quantification of the extent of cellular infiltration was performed by using a modified method previously described by Pham et al.20 For this purpose, a set of 10 columns of 10 µm width and set 20 µm apart from each other were overlaid on the H&E histology images. The deepest distance (relative to the mesh’s surface) was measured and recorded in each of the

columns. Two samples per group per time point and at least three sections per sample were analyzed to obtain the quantitative infiltration depth data. Point-to-point distance measurements were performed using ImageJ. Statistical analysis of data was performed by using a twotailed paired student t test, with p < 0.01 taken as significant.

Results and Discussions Fabrication and Characterization of mPCL/Col Scaffolds and the Derivative Hybrid Meshes. As the interest in electrospinning as a tissue engineering scaffold fabrication route has shown an exponential increase over the last four years, the challenge of obtaining true cellular/tissue integration into the scaffold has becoming even more crucial. The issue arises from the dense deposition of fibers by the electrospinning process itself. Although the deposited fibers may be very small in size and porosity per unit volume is high, the actual pore size available for cells to infiltrate and migrate into the scaffolds’ interior is too small. Furthermore, the tortuous path inside will most likely deter cellular penetration more than a few fiber layers deep. Thus, as often reported, cells are almost exclusively situated on the surface of the electrospun scaffold with penetration of the lammelipodia and fillopodia, but without considerable cell infiltration. To address this issue, we devised two modified electrospinning techniques for the establishment of cell-penetrable fibrous meshes as well as the conventional setup for comparison purposes. Electron micrographs of the as-electrospun meshes and their respective average fiber dimensions are shown in Figure 2A-E and F, respectively. All fibers morphologically appear uniform albeit with slight differences in diameter. Conventionally electrospun mPCL/Col demonstrated diameters of 513 ( 83 nm (A). A second, larger fiber group appeared in the coelectrospun meshes (B and C). These were the PEO fibers (in B, mPCL/Col-PEO) and gelatin fibers (in C, mPCL/ColGel). The process of coelectrospinning mPCL/Col with either

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Figure 3. Visualization of the scaffolds was done by means of fluorescent dye incorporation during electrospinning. Coelectrospinning results were visualized by doping mPCL/Col with rhodamine (A) and PEO with FITC (B). This process yielded meshes with PEO embedded in the mPCL/Col structure (C). Leaching of PEO in water left the mPCL/Col fibers intact (D). To visualize the embedded Heprasil, AF-BSA was incorporated into the hydrogel during electrospraying which revealed regions of HA in mPCL/Col mesh (E, top view; F, mesh cross-sectional view). Efficiency of Heprasil incorporation with respect to size of the mandrel used for collection was studied by digesting, incubating the mesh in hyaluronidase solution, and measuring the amount of released fluorescent dye (G). Marked improvement of efficiency was achieved with a diameter of 1.7 cm. Bar is 200 µm in A-F.

gelatin or PEO did not significantly alter its average dimensions. µmPCL/Col shown in D was fabricated using a higher solution concentration resulting in fibers with a diameter of 1.29 ( 0.11 µm. Coelectrospraying Heprasil with the same fiber resulted in slightly larger fibers of diameter 1.61 ( 0.27 µm. The interfiber distances and pore sizes appear visibly larger in the µmPCL/ Col and µmPCL/Col-Hep. The embedded hydrogel, however, was not clearly visible under SEM. Collapse of the hydrogel was expected during the stringent dehydration processes and drying prior to electron microscopy. Both the coelectrospinning and the electrospinning-electrospraying techniques were successful in producing meshes with mixed fiber species and embedded Heprasil hydrogel, respectively. PEO fibers were distinguishable from the mPCL/Col counterpart due to their larger size, while the gelatin fibers had a similar size as the mPCL/Col fibers. Opposing polarity was used for both techniques to overcome the issue of charge repulsion, which may impair proper mixing of the fibers/gel. When the same electric field polarities were used, a transverse mandrel movement was necessary to ensure proper mixing.21 µmPCL/Col fibers were produced by increasing the solution concentration for electrospinning to 20% w/v (compared to 12.5% w/v used for mPCL/Col, mPCL/Col-PEO and mPCL/ Col-Gel). This resulted in an increase of viscosity, which in turn restricted the bending instability phenomenon, resulting in increased fiber diameter.22 This increase in fiber size was also accompanied by an apparent increase in the dimensions of the pores, as reported previously.20 To better visualize the codeposition of different fiber species (coelectrospinning) and the fiber-hydrogel system (electrospinning-electrospraying), the fibrous meshes were imaged using a laser scanning confocal microscope (Figure 3). Rhodamineincorporated mPCL/Col (Figure 3A) and FITC-incorporated PEO (Figure 3B) showed continuous and relatively uniform formation of fibers. Coelectrospun mPCL/Col-PEO (Figure 3C) revealed the presence of both fiber species intermixed in the structure although there was an overlap of FITC signal into the rhodamine filter, resulting in the appearance of yellow fibers instead of green. PEO was easily removed by water immersion, leaving intact mPCL/Col fibers behind (Figure 3D). AF-BSA loaded Heprasil hydrogel formed HA-rich regions in the mPCL/ Col fibrous architecture after electrospraying-electrospinning processes (Figure 3E,F). The electrospraying process produced random dispersion of the regions throughout the structure. Due

to the randomness of the process, coalescence of the hydrogel during deposition could have caused larger sized regions to form, resulting in a large distribution of size from a few tens of microns to several hundred microns. The efficiency of the hydrogel incorporation with respect to collection mandrel size was investigated by fabricating the mesh using three different diameters of mandrel. Figure 3G shows the percentage efficiency when 0.8, 1.4, and 1.7 cm diameter mandrels were used as collection platforms. A marginal increment in efficiency was observed when mandrel size was increased from 0.8 to 1.4 cm. However, a more pronounced increase was observed when mandrel diameter was further increased from 1.4 to 1.7 cm, which implies a reduction in material loss from 70% down to about 25%. Given the same capillary-to-collection distance, a larger mandrel would be able to capture the ejected hydrogel droplets more efficiently than a smaller one. The issue of material loss becomes increasingly critical when bioactive molecules such as growth factors are incorporated into the mesh via the hydrogel. Meshes produced using the largest mandrel were used for subsequent cellular attachment and infiltration experiments. In Vitro Cellular Attachment and Infiltration Assays on the Fibrous Meshes. Assessments for cellular attachment, material toxicity, and extent of cellular infiltrations were done by seeding a human fetal osteoblast (hFOB) cell line onto the meshes. Cellular attachment and distribution after 3 (D3) and 10 days (D10) of seeding on the scaffolds were visualized using electron microscopy, as shown in Figure 4. Morphologically, hFOBs were flatter and more spread when cultured on conventionally electrospun mPCL/Col (Figure 4A), cospun mPCL/ColGel (Figure 4B), and cospun mPCL/Col-PEO (Figure 4C). Cells were observed to possess adhesion points with multiple fibers on these meshes. They remained predominantly as monolayer sheets with fillopodia bridging the fibers instead of extending into the pores. Cellular infiltration was not observed on the mPCL/Col and mPCL/Col-Gel meshes, while sporadic subsurface migration of cells down to several layers of fibers was observed on mPCL/Col-PEO mesh. In contrast, cells on the surface of µmPCL/Col (Figure 4D) and µmPCL/Col-Hep (Figure 4E) were more rounded and spindle-shaped with adhesion to fewer fibers due to the larger fiber size. At the later time point, cells could be found in the interior of the µmPCL/ Col and µmPCL/Col-Hep. This suggested that the pore size of these meshes was large enough to allow vertical migration of

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Figure 4. Cellular attachment and morphology of hFOBs on the five scaffolds were observed by SEM at 3 (D3) and 10 (D10) days postseeding. Cells remain exclusively on the surface when cultured on mPCL/Col (A) and mPCL/Col-Gel (B). Only marginal cellular infiltration was observed in mPCL/Col-PEO group (C). Cellular infiltration into the interior was apparent on µmPCL/Col (D) and µmPCL/Col-Hep (E). Bar is 10 µm in all images.

Figure 5. Viability of the hFOBs on the different types of scaffolds was assayed qualitatively using FDA/PI at 3 (D3) and 10 (D10) days postseeding. Inclusion of gelatin (B), PEO (C), and Heprasil (E) did not exert apparent toxicity to the cells. Viability was qualitatively comparable to that of mPCL/Col (A) and µmPCL/Col (C). Bar is 200 µm in all images.

Figure 6. Cellular infiltration after 3 (D3) and 10 (D10) days post-seeding into the scaffolds was studied using cryo-sectioning followed by H&E staining. Cellular penetration was not observed in conventionally electrospun mPCL/Col (A) and mPCL/Col-Gel (B). A thick multilayer of cells was observed instead on mPCL/Col-Gel. Marginal infiltration was seen in mPCL/Col-PEO, with a thick cell layer still forming on the surface (C). Pronounced cellular infiltration was observed in µmPCL/Col (D, arrow heads) and even more in µmPCL/Col-Hep (E, arrow heads). After 10 days of culture, cells managed to penetrate half the thickness of the mesh in µmPCL/Col (D), whereas in µmPCL/Col-Hep, cells penetrated the entire thickness (E). Bar is 50 µm in all images.

the cells from the surface into the interior. These cells exhibited a more 3D morphology adhering to multiple surrounding fibers. Qualitative cell viability assays were performed to assess the in vitro toxicity of each of the methods used to fabricate the different types of scaffolds, that is, inclusion of PEO, gelatin, and Heprasil. The confocal images of the live-dead cell assay are shown in Figure 5. Cells remain largely viable after 3 and 10 days of culturing. No apparent toxicity effects were observed from the inclusion of gelatin (Figure 5B), PEO (Figure 5C), or Heprasil hydrogel (Figure 5E). Also, no cytotoxicity effects were observed from the electrospinning solvent used. As implantable

biomaterials, minimum toxicity is desired to avoid immune rejection and fibrosis at the implantation site. Cellular penetration into the scaffolds’ structure was studied by transverse sectioning the constructs, followed by H&E histological staining (Figure 6). Conventionally electrospun mPCL/Col (Figure 6A) and mPCL/Col-Gel (Figure 6B) exhibited a monolayer of cells exclusively on the surface lacking any penetration. mPCL/Col-PEO (Figure 6C) showed a small improvement in the infiltration, with cells still predominantly on the surface. The larger resulting PEO fibers could have been

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Figure 7. Infiltration depth analysis was performed by overlaying an array of columns as depicted in A on the H&E stained section images of the different meshes. The deepest cell localization in each column was recorded and averaged for every sample group. The result is shown in B. When compared to conventional mPCL/Col, significant differences were observed in the infiltration depth of mPCL/Col-PEO, µmPCL/Col, and µmPCL/Col-Hep after 3 (#) and 10 (*) days. Inclusion of Heprasil regions in the mesh significantly improved the infiltration depth of µmPCL/ Col, with some cells penetrating the full mesh thickness.

more effective compared to the smaller gelatin fibers in mPCL/ Col-Gel in creating a less dense mPCL/Col mesh post-leaching. However, the fast leaching of the water-soluble fibers resulted in a collapse of the meshes’ structure. We conclude that this was the cause of the poor penetration of cells into this group of meshes. Although quantitative comparison was not performed on the mesh thicknesses before and after leaching of the sacrificial polymers PEO and gelatin, qualitative observation did reveal a more significant reduction in thickness in mPCL/ Col-PEO compared to mPCL/Col-Gel. An explanation for this is the difference in size between PEO and gelatin fibers. The collapse of the larger PEO fibers resulted in a higher thickness reduction of the mesh compared to the smaller gelatin fibers. Increasing mPCL/Col fiber size (Figure 6D) resulted in a less dense fibrous structure. More cells were able to migrate into the interior of the mesh, corroborating the SEM observation. Codeposition of electrosprayed Heprasil hydrogel with the µmPCL/Col (µmPCL/Col-Hep, Figure 6E) improved the cellular penetration considerably. Cells were able to infiltrate the full thickness of the µmPCL/Col-Hep mesh. The inclusion of Heprasil regions in the µmPCL/Col mesh created a reduction in volume density of the fibers and created compartments of biodegradable glycosaminoglycans hydrogel that allowed passage of infiltrating cells. Quantitative comparisons between the samples were performed via an array of vertical columns on the histological sections (Figure 7A). From these results, significant improvements on the cellular penetration depth were observed compared to conventionally electrospun nanofiber mesh, with Heprasil incorporation yielding the deepest infiltration (Figure 7B). The issue with cellular infiltration into the fiber architecture is rapidly gaining attention due to its potential in stagnating further applications of electrospun meshes or scaffolds. The conventional technique of electrospinning onto a static, flat collection screen provides a simple and inexpensive way of obtaining nanofibers. Meshes collected in this manner, however, will have tremendous fiber density built over time. Also, it has been reported that, with decreasing the electrospun fiber diameter, an increase in the number of fiber-to-fiber contacts per unit length and a decrease in the mean pore radius in the mesh are expected.23 These factors create a large size mismatch between the small pores in the structure and the larger physical size of the cells, limiting the ability of the cells to migrate and populate the scaffold’s interior. Several methods have been reported in an attempt to address this issue. Kidoaki et al. proposed the concept of selective leaching to create microvoids

using a water-soluble polymer, for example, PEO intermixed within a structural fiber architecture using their mixing electrospinning system.21 Testing this hypothesis, we successfully created fibrous mesh of mPCL/Col cospun with PEO using our coelectrospinning setup. However, this mesh only offered very limited improvements compared to conventionally electrospun fibers in terms of cell infiltration. Similarly, blending of gelatin into PCL electrospinning solution was reported to improve the ability of cell migration into the PCL mesh due to the rapid dissolution of gelatin.24 This claim, however, needs to be further substantiated with data confirming the distribution of cells across the thickness of the mesh. An interesting technique developed by Nam et al. involved simultaneous mechanical dispersion of NaCl particles and electrospinning of fibers followed by salt leaching.25 Although large pores were created with this approach, very dense fiber sheaths were still present in between the created macropores. Cellular infiltration was possible through surface migration and continuous proliferation into the delaminations or voids and not through the fiber architecture. Pham et al. reported excellent penetration by the cells into an electrospun mesh when large fibers of around 5-10 µm in diameters were used in conjunction with dynamic flow perfusion culture techniques.20 The use of larger fibers, however, would divert away from creating a more ECM-mimetic environment using electrospinning. Furthermore, they also observed a significant reduction in cell penetration when more submicron fibers were intermixed in the mesh, even with dynamic culturing. The hybrid mesh fabricated using our electrospinning-electrospraying technique allowed superior cellular infiltration over conventionally electrospun fibers. Increasing the fiber size to 1-2 µm diameters in combination with the presence of biodegradable hyaluronic acid hydrogel throughout the structure provided the cells with a sufficiently large pore system and interconnectivity for vertical migration. Another prospective advantage in the Heprasil codeposition is the capability of loading bioactive molecules into the hybrid mesh. As reported previously, the HA hydrogel can be utilized as a controlled release vehicle for growth factors such as vascular endothelial growth factor (VEGF) or fibroblast growth factor (FGF) due to the presence of cross-linked heparin in the structure.14,26,27 The potential of creating 3D fibrous scaffolds and loading specific bioactive molecules for tissue-tailored applications is currently underway in our laboratory.

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Conclusions With the intention of addressing the issue of lack of cellular penetration into an electrospun fibrous mesh, we attempted three different electrospinning methodologies: selective fiber leaching, micron-sized structural fibers, and a combination of microfibers with the codeposition of a biodegradable glycosaminoglycan hydrogel, Heprasil. To achieve this, a simultaneous coelectrospinning and electrospinning-electrospraying setup were devised. Each method conferred improvements over conventionally electrospun fibers, albeit to differing extents. By far, the most successful method was using a combination of micron-sized fibers with Heprasil, which created a fibrous mesh that was found to be biocompatible and allowed significant infiltration of cells into its structure. The inclusion of Heprasil regions in the fiber architecture would not only allow better infiltration of cells into the mesh but also potentially permit the controlled release of bioactive factors for specific tissue applications. Acknowledgment. The authors thank Glycosan Biosystems for their generous contribution in terms of the Heprasil hydrogel kits.

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