Porous 3D modeled scaffolds of bioactive glass and photocrosslinkable poly(ε-caprolactone) by stereolithography

Composites Science and Technology 74 (2013) 99–106

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Porous 3D modeled scaffolds of bioactive glass and photocrosslinkable poly(e-caprolactone) by stereolithography Laura Elomaa a, Anne Kokkari b, Timo Närhi b, Jukka V. Seppälä a,⇑ a b

Aalto University School of Chemical Technology, Department of Biotechnology and Chemical Technology, Polymer Technology, P.O. Box 16100, FI-00076 Aalto, Finland University of Turku, Faculty of Medicine, Institute of Dentistry, Lemminkäisenkatu 2, FI-20520 Turku, Finland

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Article history: Received 5 September 2012 Received in revised form 12 October 2012 Accepted 22 October 2012 Available online 2 November 2012 Keywords: A. Glasses A. Polymer–matrix composites (PMCs) B. Porosity/Voids D. Scanning electron microscopy (SEM) Photocrosslinking

a b s t r a c t Bioactive glass is known to benefit cell interactions of polymeric tissue engineering scaffolds. Most likely, the best response is obtained when the glass is on the scaffold surface without a cover. We combined a photocrosslinkable poly(e-caprolactone) resin with bioactive glass in a rapid prototyping process. Bioactive glass was homogeneously distributed through the highly porous scaffolds and their surface. Ion release measurements in simulated body fluid revealed a rapid decrease in calcium and phosphorus concentrations. The presence of calcium phosphate deposits on the surface of the composite scaffolds indicated in vitro bioactivity. The bioactive glass increased the metabolic activity of fibroblasts. Our work showed that stereolithography enables the fabrication of well-defined composite scaffolds in which the bioactive glass is homogeneously distributed on the surface and readily available for rapid ion release and cell interactions. By stereolithography, an unwanted polymer layer covering the BG particles on the scaffold surface can be successfully avoided. Ó 2012 Elsevier Ltd. All rights reserved.

1. Introduction There is a clinical need for bioactive, biodegradable materials that enhance the formation of new tissue at the site of a wound resulting from disease or trauma in the body. Hench et al. first published results concerning the bone bonding ability of bioactive glass (BG) in 1971 [1]. Bioactive glasses are inorganic surface-active biomaterials that can interact with bone tissue because of the rapid chemical reactions that can take place on the glass surface. When exposed to a biological fluid, BG shows osteopromotive behavior by rapidly releasing soluble ionic species and forming a carbonated hydroxyapatite layer on the surface that is responsible for strong bonding with bone tissue [2–4]. The released ions regulate the cell cycle of osteoblasts [5] and play an important role in angiogenesis [6]. An approach to overcoming a problem of brittleness and poor processability of BG is to use it with biodegradable polymers, as has been done with polylactide [7], polyglycolide [8], poly(e-caprolactone) (PCL) [9], and their copolymers [10]. These composites mimic bone tissue, which is mainly composed of an organic phase of collagen and an inorganic phase of hydroxyapatite [11]. The field of tissue engineering applies three-dimensional (3D) scaffolds as temporary substitutes for the extracellular matrix in which cell proliferation and differentiation occurs [12]. For producing porous ⇑ Corresponding author. Tel.: +358 9 47022614; fax: +358 9 47022622. E-mail address: jukka.seppala@aalto.fi (J.V. Seppälä). 0266-3538/$ - see front matter Ó 2012 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.compscitech.2012.10.014

BG/polymer scaffolds, conventional fabrication methods, such as salt-leaching [13,14], phase separation [15–17], and modified sugar template-particulate leaching [18], are commonly used. However, the drawbacks of the conventional methods are that the pores in scaffolds are poorly interconnected and BG particles on the scaffold surface are covered by a polymer film during the fabrication. Additive manufacturing technologies (AMTs) and computeraided design (CAD) enable the precise control over pore geometry and the design of 3D structures for the preparation of tissue engineering scaffolds [19,20]. One of the most researched AMT methods is stereolithography, which is based on photocrosslinking of a liquid polymer resin into a precisely defined 3D structure [21]. A variety of acrylated or methacrylated biophotopolymers have been synthesized for use in lithography [22–24] and the features of degradation products of photocrosslinked networks has been discussed [25]. As applied to composite materials, stereolithography has been used to prepare fiber reinforced structures of an acrylic-based urethane photopolymer and short glass fibers [26] and an epoxy photoresin and carbon fibers [27]. Also polypropylene fumarate-based [28] and divinyl ester-based [29] composite scaffolds containing hydroxyapatite particles have been fabricated. Stereolithography, however, has not been commonly used to make biodegradable BG composite structures. Recently, a lithographybased technique was used to produce scaffolds of photosensitive slurry containing BG powder mixed with acrylate-based monomer and organic solvent [30].

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The aim of this study was to obtain highly porous composite scaffolds of BG and PCL with a mathematically defined architecture and enhanced in vitro bioactivity. The objective was to avoid BG particles on the scaffold surface being covered by a polymer layer. We chose the stereolithography as a fabrication method because of its high resolution and mild processing conditions: only room temperature is needed. Previously, we used stereolithography to prepare PCL-based scaffolds without any active components [31]. In this study, we successfully prepared the scaffolds with different BG contents. The scaffolds were characterized by mechanical tests and in vitro bioactivity tests in simulated body fluid (SBF). Cell proliferation on the scaffolds was evaluated by fibroblast culturing. 2. Materials and methods

corner. The structures were multiplied and shaped to meet the desired dimensions. The Magics Envisiontec and Envisiontec Perfactory softwares were used for finishing and slicing the stl-files. The scaffolds were built using the photocrosslinkable resin containing the methacrylated PCL macromer mixed with LucirinÒ TPO-L photoinitiator (3 wt.%), Orasol Orange G dye (0.10 wt.%), and 0, 5, 10, or 20 wt.% of bioactive glass (BG0, BG5, BG10, BG20). A layer thickness was 50 lm with an exposure time of 12 s and a light intensity of 1600 mW dm2. A lens with a focal distance of 75 mm was used. After fabrication, any uncured macromer was removed by extraction from the scaffolds in a mixture (3:1) of acetone and isopropanol. The scaffolds were dried under vacuum conditions until a constant weight was achieved. 2.4. Characterization of resins and scaffolds

2.1. Materials The e-caprolactone monomer (Solvay Chemicals) was dried over calcium hydride (Acros Organics) and distilled under reduced pressure before polymerization. Tin(II) 2-ethylhexanoate (Sn(Oct)2), trimethylolpropane, methacrylic anhydride, and triethylamine (Sigma–Aldrich) were used as received. Analytical grade dichloromethane (Merck Chemicals) was dried over CaH2 (Merck Chemicals) and filtered before use. LucirinÒ TPO-L photoinitiator and OrasolÒ Orange G dye were obtained from BASF. Bioactive glass (S53P4) was obtained from Vivoxid Ltd. A small granule size (<45 lm) was chosen to produce a high glass area/volume ratio. The composition of BG in weight percentages was SiO2 (53); Na2O (23); CaO (20); P2O5 (4). Human gingival fibroblasts were a donation from Dr. Jaana Willberg, University of Turku, Finland. Dulbecco’s Modified Eagle Medium (DMEM), fetal bovine serum, penicillin-streptomycin, and the AlamarBlueÒ (AB) assay were from Life Technologies. The hematoxylin solution was from Fluka Analytical. 2.2. PCL macromer synthesis A three-armed hydroxyl-terminated PCL oligomer was synthesized by the ring-opening polymerization of e-caprolactone monomers using trimethylolpropane as an initiator and 0.02 mol% Sn(Oct)2 as a catalyst. The oligomer was polymerized at 140 °C for 24 h under a nitrogen atmosphere. The photocrosslinkable PCL macromer was produced using methacrylic anhydride to cap the hydroxyl-terminated oligomer. The oligomer was dissolved in dichloromethane, and a 30 mol% excess of methacrylic anhydride and triethylamine was added to the reaction. Methacrylation was continued for 24 h at room temperature under a nitrogen atmosphere. After methacrylation, the reaction solution was washed with a saturated aqueous sodium bicarbonate solution and dried with magnesium sulfate. The residual solvent was evaporated under vacuum. 2.3. Scaffold fabrication Porous 3D scaffolds were prepared using an EnvisionTEC PerfactoryÒ Mini Multilens SLA machine. The gyroid pores in the scaffolds were modeled using the free K3DSurf surface generator (k3dsurf.sourceforge.net). The initial CAD-file for the 3D objects was obtained by visualizing the surface described by Eq. (1) with boundary conditions x, y, z = [10p, 10p].

cosðxÞ sinðyÞ þ cosðyÞ sinðzÞ þ cosðzÞ sinðxÞ  1 ¼ 0

ð1Þ Ò

The stl-files were prepared using the Rhinoceros software (McNeel). The 3D structures were rotated so that the pore channels were not open from top to bottom vertically but from corner to

1 H NMR spectra of the oligomer and the macromer were recorded on a Bruker Avance-III 400 MHz spectrometer. The sample was dissolved in deuterated chloroform and measured in a 5 mm NMR tube at room temperature. The molecular weight of the oligomer was calculated by a 1H NMR end-group analysis. Thermal properties of the macromer were determined from measurements using a Mettler Toledo Stare DSC 821e differential scanning calorimeter (DSC). The sample was first heated from 25 to 100 °C at a rate of 20 °C/min. The temperature was later decreased to 100 °C at 10 °C/min. After 2 min at 100 °C, the sample was heated from 100 to 100 °C at a recorded rate of 10 °C/min. Glass transition and melting temperatures were determined from the second heating scan. The compression moduli of the scaffolds were measured with a stress controlled rotational rheometer AR G2 (TA instruments) using a 20 mm, plate-and-plate geometry. The samples were discoid scaffolds with a 20 mm diameter and a 3 mm height, and the measurements were taken in triplicate. For a comparison purpose, similar nonporous PCL samples were prepared by SLA and their compression moduli were also tested. The compression was performed at a temperature of 23 °C and a speed of 10 lm/s until the stress normal was 45 N. To test wet samples, the scaffolds were immersed in SBF at 37 °C for 24 h before measurements. The water uptake of the scaffolds was defined as the relative increase in the weight of the polymer fraction. A micro-CT system (l-CT 40, operated by Scanco Medical AG, Switzerland) was used to analyze the 3D morphology of the scaffolds at a 6.6 lm voxel resolution.

2.5. In vitro bioactivity and ion release studies The SBF was prepared by dissolving reagent grade NaCl, NaHCO3, KCl, K2HPO43H2O, MgCl26H2O, CaCl22H2O, and Na2SO4 in deionized water. The solution was buffered with tris(hydroxymethyl)aminomethane and hydrochloric acid to physiological pH of 7.40 at 37 °C. The ion composition of the SBF corresponded to the inorganic portion of human blood plasma [32]. Each scaffold was immersed in 50 mL of SBF and enclosed in a test tube that was incubated and shaken in a 37 °C water bath and were monitored over 21 days. Four replicate scaffolds (10  10  3 mm3) were analyzed at predetermined time points for ion concentrations of calcium, phosphorus, and silica. Two scaffolds were removed from the SBF at predetermined times for SEM analysis. In addition, small samples of SBF were collected from each test tube and the calcium, silica, and phosphorus concentrations were measured in triplicate. The SBF solution was not changed between the assayed time points. Colorimetric measurements of silica and orthophosphate concentrations were based on the molybdenum blue method. Any silicomolybdate complex was reduced with a mixture of

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100,000 cells/cm2 and cultured for up to 14 days. For the control, a standard number of fibroblasts were seeded in the 24-well plates. The culture medium was changed three times a week. Cell proliferation was determined using the AlamarBlueÒ assay. The scaffolds (10  10  3 mm3, n = 4) and fresh assay solution (phenol red-free DMEM supplemented with 10% serum and 10% AB reagent) were added to the wells. After 3 h of incubation, the absorbance of the solution was measured at 570 nm and 595 nm using the ELISA plate reader. The measured absorption was used to calculate the reduction of AB reagent. Cell activities were normalized in relation to the cell activity of the scaffolds without BG after day 1. The seeding efficiency of the scaffolds was determined at day 1 in relation to the activity of the fibroblasts seeded on 24 plates. 2.7. Scanning electron microscopy and elemental analysis Fig. 1. 1H NMR spectrum of the methacrylated PCL macromer.

1-amino-2-naphthol-4-sulphonic acid and sulphite, and tartaric acid was used to eliminate interference from phosphates. The antimonyphosphomolybdate complex was reduced with ascorbic acid. Calcium concentrations were determined using the ortho-cresolphtalein complexone (OCPC) method. This assay consists of OCPC and 8-hydroxyquinol in an ethanolamine/boric acid buffer. Absorption at the following wavelengths of 820 nm for silica, 690 nm for phosphorus, and 570 nm for calcium were measured using UV1601 spectrophotometer (Shimadzu) or Multiskan MS ELISA plate reader (Labsystems).

The scaffolds were examined by scanning electron microscopy (SEM) after immersion in SBF and cell culturing. The scaffolds from the SBF test were washed twice with deionized water and dried prior to carbon coating. The scaffolds were subsequently rinsed with PBS and dried in an alcohol series of increasing alcohol composition. The samples were carbon coated and secondary electron images recorded using a JSM-5500 scanning electron microscope (JEOL). After immersion in SBF, the microstructure and elemental composition of the surfaces of the scaffolds were analyzed by a SEM equipped with a PGT Prism 2000 Energy Dispersive X-Ray Spectrometer (EDS) (Princeton Gamma-Tech Instruments Inc.). 2.8. Stereomicroscopy and confocal fluorescence microscopy

2.6. Cell cultures Human gingival fibroblasts were maintained in DMEM supplemented with 10% fetal bovine serum and 100 U/lg penicillinstreptomycin and incubated at 37 °C in a 5% CO2 environment. Semi-confluent cultures were trypsinized, and the cells were counted and resuspended in complete culture medium. Passage 15 cells were used in the culture. Prior to seeding, the scaffolds were pre-incubated to increase their wettability and to overcome some of the ion release from the BG, which can cause a harmful pH increase in the culture conditions. The scaffolds were subsequently placed in 24-well plates and pre-incubated at room temperature two times in phosphate buffered saline (PBS) and two times in cell culture media for 1 h each, totaling 4 h of incubation time. Fibroblasts were later plated on the scaffolds at a density of

The scaffolds were subsequently washed with deionized water and dried in an increasing alcohol series prior to imaging using a Wild’s stereomicroscope (Leica Geosystems) equipped with a Leica DC 300 V2.0 digital camera (Leica Microsystems). For the confocal microscopy, the cultured scaffolds were washed twice with PBS, and the cells were fixed with 4% paraformaldehyde. The fixed cells were permeabilized with 0.1% Triton X-100 for 5 min, and non-specific binding was blocked with 5% bovine serum albumin and 2.5% of fetal bovine serum. The actin cytoskeletons were stained with TRITC (tetramethylrhodamine isothiocyanate)-labeled phalloidin, and the cell nuclei were stained with DAPI (diamidino-2-phenylindole). The cells were imaged using a Leica TCS-SP confocal laser scanning microscope equipped with an argon-krypton laser (Leica Microsystems).

Fig. 2. l-CT micrograph of the pore structure within scaffolds (a) BG0 and (b) BG20.

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Fig. 3. (a) SEM image of a BG0 scaffold, (b) optical stereomicroscope image of a BG20 scaffold and (c and d) SEM images of a surface of a BG20 scaffold.

Fig. 4. (a) Pore size distribution of the scaffolds estimated by cubic thickness and (b) accessible pore volumes of the scaffolds determined by l-CT.

Statistical analyses were performed using the IBM SPSS Statistics 20 software package (SPSS Inc.). Data were analyzed with one-way ANOVA followed by Tukey’s post hoc test.

integral of the methylene group of the PCL backbone (peak e). The calculated molecular weight of the oligomer was 702 g/mol. The peaks i, j, and k in Fig. 1 are attributed to the functionalized ends of the macromer. The melting temperature of the macromer was 15.1 °C, and the glass transition temperature was 60.8 °C.

3. Results

3.2. Surface roughness and porosity of scaffolds

3.1. Photocrosslinkable macromer and resins

Three-dimensional scaffolds with a gyroid pore design were successfully prepared by stereolithography. Fig. 2 shows the l-CT micrograph of (a) a BG0 scaffold and (b) a BG20 scaffold. The pores were open throughout the scaffolds, and the roughness of the BG20 scaffold was an indication that the bioactive glass is homogeneously distributed on the surface of the scaffolds. The surface roughness increased with the amount of BG in the scaffolds. The SEM and optical stereomicroscopy images Fig. 3 revealed the big difference between the surfaces of BG0 and BG20 scaffolds. In contrast to the smooth surface of a BG0 scaffold, the surface of a

2.9. Statistics

Polymerization of the hydroxyl-terminated PCL oligomer was allowed to proceed to completion. The hydroxyl-terminated PCL was functionalized with methacrylic anhydride until all the hydroxyl ends had reacted. Fig. 1 shows the 1H NMR spectrum of the resulting methacrylated PCL macromer. The average number of e-caprolactone units per PCL oligomer was calculated from a 1H NMR spectrum of the oligomer by comparing the peak integral of the methyl group of the TMP initiator (peak h) with the peak

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Fig. 5. Compression moduli of the scaffolds with different BG contents measured under dry conditions and after a 24 h soak in SBF. The error bars represent standard deviations (n = 3).

BG20 scaffold was fully covered by the BG particles. Especially the stereomicroscopy image Fig. 3b shows the BG particles on the scaffold surface to be without any polymer layer covering them. The total porosity of the scaffolds measured with l-CT were 77 vol.% (BG0), 75 vol.% (BG5), 70 vol.% (BG10), and 63 vol.% (BG20). Fig. 4a shows the distribution of pore sizes in the different scaffolds. The mean pore sizes were 594 lm (BG0), 555 lm (BG5), 517 lm (BG10), and 476 lm (BG20). The pore size distributions were narrow, which indicates that the pores were relatively uniform in size. Fig. 4b shows accessible pore volumes that were high for all the scaffolds, indicating good interconnectivity of the pore networks. 3.3. Compression moduli The effect of BG on the mechanical properties of the scaffolds was evaluated by measuring the compressive modulus of the scaffolds when they were either dry or wet. Fig. 5 shows that in dry conditions, the BG20 scaffolds increased the compression modulus from 1.4 to 3.4 MPa compared to the BG0 scaffolds. There was no statistically (p < 0.01) significant difference between the compressive moduli of the BG5 and BG10 scaffolds. Under wet conditions, the compression moduli of the scaffolds decreased. The wet BG20 scaffolds increased the compression modulus from 1.1 to 2.5 MPa compared to the wet BG0 scaffolds. Otherwise, there was no statistically significant difference between the wet BG5, BG10, and BG20 scaffolds. The compression moduli of nonporous PCL samples were 10.4 ± 0.8 MPa (dry) and 10.3 ± 0.6 MPa (wet). The water uptake values for each type of scaffold were 150 ± 7% (BG0), 160 ± 10% (BG5), 176 ± 8% (BG10), and 190 ± 10% (BG20). The water uptake of photocrosslinked nonporous PCL samples was 1.66 ± 0.02%. 3.4. Bioactivity and ion release in SBF The concentration changes of calcium, silica, and phosphorus in the SBF were monitored for 21 days Fig. 6. The BG20 scaffolds showed a large decrease in calcium and phosphorus concentrations from the beginning onwards, whereas the BG10 scaffolds showed only a moderate decline in calcium concentration after 9 days. The BG5 scaffolds showed slightly elevated calcium levels compared to that of the PCL scaffolds (BG0); however, the overall ionic concentrations remained steady in both sample groups. The con-

Fig. 6. Concentrations of calcium, phosphorus, and silica over 21 days in SBF. The error bars represent standard deviation (n = 4).

centrations of phosphorus remained steady for the BG0, BG5, and BG10 scaffolds. No silica was detected in SBF for the pure polymer. SEM imaging and elemental analysis of the surfaces were performed after SBF immersion. Fig. 7 presents SEM images of the BG scaffolds after (a) 3 days, (b) 10 days, and (c, d) 21 days in SBF. The BG20 scaffolds immersed in SBF for 3 days and longer exhibited the formation of precipitate Fig. 7a. The surface cut from the middle of the BG20 scaffold was covered completely with calcium phosphate deposits after 10 days in SBF Fig. 7b. After 21 days, a portion of the calcium phosphate was detected on the surface of the BG10 scaffold Fig. 7c. The BG20 scaffolds were fully covered with calcium phosphate Fig. 7d. The BG5 and BG0 scaffolds did

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Fig. 7. SEM images of the scaffolds after immersion in SBF. (a) BG20 after 3 days, (b) BG20 after 10 days, (c) BG10 after 21 days, and (d) BG20 after 21 days.

on the BG10 and BG20 scaffolds after days 3 and 7 were significantly higher (p < 0.05) than on the BG0 and BG5 scaffolds. On day 1, all the composite scaffolds showed significant increases in proliferation rates compared to those determined for the BG0 scaffolds. SEM images show the cells on a surface of a BG10 scaffold Fig. 9a and a BG20 scaffold Fig. 9b after 14 days in cell culture. Typical fibroblast morphology was observed by confocal fluorescence microscopy. The cells were distributed on the surfaces of the scaffolds regularly, as shown in fluorescence images taken from the top and the middle of the samples Fig. 9c and d. 4. Discussion

Fig. 8. Cell activity during 14 days of fibroblast culturing. Statistically significant differences were found between groups A, B, and C (p < 0.05) within the defined time points. The error bars represent standard deviations (n = 4).

not exhibit the precipitation of calcium. The elemental analysis of the BG20 scaffolds after 3 days in SBF revealed the molar ratio of calcium and phosphorus to be 1.64–1.86. 3.5. Cell proliferation The seeding efficiency of the scaffolds changed between 17% (BG0) and 41% (BG20). The proliferation of fibroblasts on different types of scaffolds was measured over a period of 14 days Fig. 8, where the cell activity was observed to increase from day 1 to day 7. From 7 to 14 days, the cells continued to proliferate in a number of the sample groups, while others exhibited a slight decrease in cell activity at day 14. The BG20 scaffolds showed the most active cell proliferation. The BG5 and BG10 scaffolds showed a twofold increase in activity compared to the activity on day 1. The cell proliferation on the BG0 scaffolds increased until day 14 to over threefold compared to the activity on day 1. Cell growth

Bioactive glass was successfully used in stereolithography to prepare bioactive, polymer-based tissue engineering scaffolds. Surfaces of the composite scaffolds showed a homogeneous distribution of BG, predicting rapid ion release from the scaffolds in SBF. Addition of BG increased the compression modulus of the scaffolds, which is a desired effect on bone applications. Although the same pore design was used for all the scaffolds, the inclusion of BG decreased the pore size and the total porosity. The mean pore sizes, however, remained at approximately 500 lm, which has been reported to be optimal for bone formation [33]. Accessible pore volume analysis confirmed that the pores penetrated through the scaffolds, which is essential for nutrients and body fluids to flow freely inside the 3D structure. High interconnectivity of the pores and open BG particles on the scaffold surface are significant advantages of these new composite structures compared to previously reported structures prepared by conventional methods. The amounts of calcium and phosphorus decreased rapidly in SBF tests of BG20 scaffolds, whereas the amounts were more stable in samples of other scaffolds. Calcium phosphate precipitates were visible on the BG20 scaffolds after 3 days in SBF, which indicates strong in vitro bioactivity. Although deposits were also observed on the BG10 scaffolds after 21 days, we concluded that they may not contain an adequate amount of BG to promote bone healing.

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Fig. 9. Images after 14 days in cell culture. SEM images from (a) top of a BG10 scaffold and (b) a cross-section from the middle of a BG20 scaffold. Fluorescence microscopy images, where the actin filaments and cell nuclei are stained: (c) the second layer of a BG10 scaffold imaged through a pore from the top surface and (d) a stack of 87 individual images from the middle of a BG20 scaffold. The line and asterisk (⁄) show where the surface was cut (no cells), and the arrows indicate the edges of the pores.

Cells were spread over the scaffolds at a 3 mm thickness. This distribution is suitable for static culture, and diffusion does not restrict the cell proliferation. Cell proliferation rates increased by a factor of two in all sample groups compared with their maximum rate and the rate at day 1. The number of cells that were attached to the scaffolds originally appeared to have a strong effect on the maximum rate of proliferation. The differences in cell attachment were not further studied; however, we hypothesize that the changes in surface chemistry and topography induced by the addition of the BG play a major role. Only the most highly populated BG20 group showed a decrease in activity after day 14. This decrease in cell proliferation rate on day 14 was expected, since the slowing down of cell division is attributed to an increase in the number of cells within the available spaces on the surface of the scaffolds. Fluorescence microscopy revealed that the cells adhered to the scaffold surface also inside the 3D structures, which is an essential requirement for proper tissue formation. 5. Conclusions In this study, we successfully prepared well-defined porous tissue engineering scaffolds of bioactive glass and methacrylated poly(e-caprolactone). Stereolithography enabled the homogeneous incorporation of BG into the 3D structures, and the highly interconnected pore networks were obtained. The composite scaffolds were bioactive in vitro, demonstrating the rapid release of silica and the precipitation of calcium phosphate in simulated body fluid. This bioactivity resulted in an enhancement in the attachment and proliferation of human fibroblasts in the 3D scaffolds. Our study suggests that photocrosslinked composite scaffolds of BG and poly(ecaprolactone) prepared by stereolithography have great potential

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