Article pubs.acs.org/Biomac
Polycaprolactone Scaffolds Fabricated with an Advanced Electrohydrodynamic Direct-Printing Method for Bone Tissue Regeneration Seung Hyun Ahn,† Hyeong Jin Lee,† and Geun Hyung Kim*,†,‡ †
Department of Mechanical Engineering, Bio/Nanofluidics Lab and ‡Department of Dental life science, College of Dentistry, Chosun University, 375 Seosok-dong, Dong-gu, Gwangju 501-759, South Korea ABSTRACT: Electrohydrodynamic (EHD) direct writing has been used in diverse microelectromechanical systems and various supplemental methods for biotechnology and electronics. In this work, we expanded the use of EHD-induced direct writing to fabricate 3D biomedical scaffolds designed as porous structures for bone tissue engineering. To prepare the scaffolds, we modified a grounded target used in conventional EHD direct printing using a poly(ethylene oxide) solution bath, elastically cushioning the plotted struts to prevent crumbling. The fabricated scaffolds were assessed for not only physical properties including surface roughness and water uptake ability but also biological capabilities by culturing osteoblast-like cells (MG63) for the EHD-plotted polycaprolactone (PCL) scaffold. The EHD-scaffolds showed significantly roughened surface and enhanced water-absorption ability (400% increase) compared with the pure rapidprototyped PCL. The results of cell viability, alkaline phosphatase activity, and mineralization analyses showed significantly enhanced biological properties of the scaffold (20 times the cell viability and 6 times the mineralization) compared with the scaffolds fabricated using RP technology. Because of the results, the modified EHD direct-writing process can be a promising method for fabricating 3D biomedical scaffolds in tissue engineering.
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printing applications in electronics,18,19 information display,20,21 optical films,22 drug discovery,23 micromechanical devices,24 and other areas.25 EHD printing is a patterning method that uses a fine jet generated at the apex of a liquid cone under an applied electric field.26 When the jet is deposited on a substrate, a fine pattern of micro/nanoparticles remains on the substrate after solvent evaporation. Therefore, EHD printing has been used to create fine patterns consisting of solid micro/ nanoparticles or polymer droplets. Recently, Edirsinghe et al. used EHD direct writing for biomedical scaffold fabrication to produce fine threads (2 mm thick) the process must be improved because the scaffolds are thin and smooth, and layer-by-layer structure control is poor.27,28 To overcome the limitations of the conventional EHD printing for biomedical scaffolds, we propose a modified EHD process for fabricating highly porous and surface-roughened/
INTRODUCTION In tissue regeneration, the physical structure of biomaterials can influence signal expression and subsequent differentiation of seeded cells.1,2 In particular, porosity and pore size can affect osteoblastic differentiation of mesenchymal stem cells (MSCs) and the production of extracellular matrix (ECM) proteins. 3,4 Therefore, the physical structure of various biomaterials (collagen, chitosan, synthetic polymers, and bioceramics) is an important design parameter in fabricating biomedical scaffolds. Several fabrication methods are used to prepare biomedical scaffolds, including phase separation,5 gas formation,6 porogen-leaching,7 and electrospinning.8,9 These methods produce highly porous scaffolds, but control of their intermicroporous structure has been an issue. Recently, solidfreeform fabrication (SFF) methods, including melt-plotting, selective laser sintering, sterolithography, 3D printing, and fuse deposition modeling have attracted attention because they produce scaffolds with controlled porosity, pore size, interconnectivity, and mechanical properties, influencing various cell signals for tissue regeneration and differentiation.10−16 Although these methods possess enormous potential as fabrication techniques, they suffer from limitations, including high cost, low-resolution struts, material restraints, long processing times, and smooth-surfaced scaffolds, which deteriorate initial cell attachment and proliferation.14−16 Electrohydrodynamic (EHD) jet printing, a direct-writing technique, has been used to create micro/nanosized structures on substrates.17 Many reports have described various EHD© 2011 American Chemical Society
Received: August 15, 2011 Revised: October 26, 2011 Published: November 9, 2011 4256
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layer-by-layer 3D poly(ε-caprolactone) (PCL) structures with thickness of 2 to 3 mm. Because PCL has been effectively used to enhance bone in-growth and regeneration in the treatment of bone defects, we used PCL in the EHD direct writing process. According to Boyan et al., adhesion and proliferation of osteoblast are highly sensitive to surface topography and mechanical stimulation29,30 so that we applied the highly roughened PCL scaffold to the bone tissue regeneration. To fabricate pore structure-controlled 3D scaffolds, a modified EHD process was supplemented with a grounded target system in which a viscous poly(ethylene oxide) (PEO) solution was used. To determine the processing window, we studied the effects of PCL weight fraction and applied electric field on the final struts. Using this method, we acquired structurally stable 3D PCL structures (>2 mm thick) with controlled strut sizes and high porosity (>70%). The cell viability, ALP activity, and mineralization of the fabricated scaffolds were evaluated by culturing osteoblast-like cells (MG63).
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Table 1. Fabrication Condition, Pore Structure (Strut Diameter, Pore Size, and Porosity), and Modulus for EHDPlotted and RP-Plotted PCL Scaffolds RP-PCL pneumatic pressure (kPa) processing temp. (°C) nozzle speed (mm s−1) electric field (kV cm−1) nozzle diameter (μm) flow rate (mL h−1) strut diameter (μm) pore size (μm) porosity (%) Young’s modulus (MPa)
500 120 3.2 inner: 200, outer: 240 220 ± 25 205 ± 27 56 12.5 ± 2.3
EHD-PCL 25 10.5 2.5 inner: 510, outer: 800 1 205 ± 61 347 ± 86 78 6.3 ± 1.8
speed of 0.2 mm s−1. The apparent porosity of the scaffolds was obtained using the following equation: porosity (%) = (1 − apparent density of scaffold/bulk density of scaffold) × 100, where the apparent density is defined as (weight of scaffold)/(volume of rectangular scaffold). In Vitro Osteoblast-Like Cell (MG63) Culture. The scaffolds (5 × 5 × 2 mm3) used with the cell cultures were sterilized with 70% EtOH and UV light and placed in culture medium overnight. As a control, we used the RP-scaffold with the same dimension. Osteoblastlike cells (MG63; ATCC, Manassas, VA) were used to observe cellular behavior in the scaffolds. MG63 cells were cultured in Dulbecco’s modified Eagle’s medium (DMEM; Hyclone, Logan, UT) supplemented with 10% fetal bovine serum (Hyclone) and 1% penicillin/ streptomycin (Hyclone). The tissue culture plate was 24-well and the volume of cell culture medium added to each well was 400 μL. The cells were maintained up to passage seven and collected by trypsin− ethylenediaminetetraacetic acid (EDTA) treatment. Approximately 50 μL of culture medium containing 5 × 104 cells was seeded on the scaffolds in 24-well culture plates. The cells were allowed to attach to the scaffold for 4 h; then, 400 μL of fresh culture medium was added to each well, and the cells were incubated in an atmosphere of 5% CO 2 at 37 °C. The medium was changed every other day. To assess the morphology of cells on the scaffolds, the cells were examined by SEM after 3 days. The cell/scaffold constructs were fixed in 2.5% glutaraldehyde and dehydrated through a graded ethanol series. Dried scaffolds were coated with gold and examined under SEM. Viable cells were determined by MTT assay (Cell Proliferation Kit I; Boehringer Mannheim, Mannheim, Germany) based on the cleavage of yellow tetrazolium salt MTT by mitochondrial dehydrogenases in viable cells to produce purple formazan crystals. Cells on the scaffold were incubated with 0.5 mg mL−1 MTT for 4 h at 37 °C, and absorbance was measured at 570 nm using a microplate reader (EL800; Bio-Tek Instruments, Winnooski, VT). Five samples were tested for each incubation period, and each test was performed in triplicate. After 3 days of cell culture, the scaffolds were fluorescently stained with DAPI (diamidino-2-phenylindole) to characterize the nuclei of the cells on the surface of the scaffolds. The fluorescence images were acquired with a ZEISS axiovision 1 (Germany). Alkaline Phosphatase Activity. For MG63 cells seeded in the scaffolds for 5 and 10 days, alkaline phosphatase (ALP), a marker of osteoblast activity, was assayed by measuring the release of pnitrophenol from p-nitrophenyl phosphate (p-NPP). The scaffolds seeded with MG63 cells were rinsed gently with phosphate-buffered saline (PBS) and incubated in Tris buffer (10 mM, pH 7.5) containing 0.1% Triton X-100 for 10 min. Next, 100 μL of the lysate was added to a 96-well tissue culture plate containing 100 μL of p-NPP solution, prepared using an ALP kit (procedure no. ALP-10; Sigma). In the presence of ALP, p-NPP was transformed to p-nitrophenol and inorganic phosphate. ALP activity was determined by measuring the absorbance at 405 nm using a microplate reader (Spectra III; SLT-Lab Instruments, Salzburg, Austria).
MATERIALS AND METHODS
Materials. PCL (density = 1.135 g/cm3; Mw = 60 000; melting point = 60 °C) and PEO (Mw = 900 000) were obtained from SigmaAldrich (St. Louis, MO). To fabricate the struts using EHD plotting, 12 wt % PCL in a 20:80 solvent mixture of methylene chloride (MC; Junsei Chem., Tokyo, Japan) and dimethylformamide (DMF; Junsei Chem.) were used. To fabricate 5 wt % of PEO solution, we dissolved PEO with distilled water. Scaffold Fabrication. An EHD-plotting system connected to a three-axis robot was used to fabricate the PCL scaffold. The PCL solution was injected using a syringe pump at 25 °C, and the EHD apparatus moved automatically according to the structure designed by the CAD system. The procedure was repeated several times to acquire layered PCL struts. The speed of the nozzle was set to 10.5 mm s −1. The electric field was 2.5 kV cm−1, and the flow rate of the PCL solution was 1 mL h−1 (KDS 230; KD Scientific, Holliston, MA). A power supply (SHV300RD-50K; Convertech, Seoul, South Korea) was used to fabricate the struts. To construct the EHD-plotted 3D PCL scaffold, we used a 5 wt % PEO solution as a target solution bath, and a grounded copper plate was immersed in the bath. After fabricating the layered structures, PEO was dissolved with purified water, and the structures were dried for 24 h to retain their 3D architecture. To remove the edge defects of the fabricated scaffolds, the edge of the scaffolds was trimmed about 1 to 2 mm. The physical dimensions of the finally fabricated scaffold were 60 × 60 × 2 mm3. To fabricate a RP scaffold control, a three-axis robot system with a heating dispenser was used. The PCL powder was melted in a heating cylinder at 120 °C, and it was extruded through a 200 μm nozzle (inner diameter) with the plotting speed, 3.2 mm s−1. The applied pneumatic pressure during the extrusion of the scaffold was set as a 500 ± 20 kPa. Details of the processing conditions are listed in Table 1. Scaffold Characterization. The surface topography of the scaffolds was observed under an optical microscope (BX FM-32; Olympus, Tokyo, Japan) connected to a digital camera and a scanning electron microscope (SEM; Sirion, Hillsboro, OR). Water uptake was calculated by weighing the scaffolds before and after soaking in distilled water for 2 h. The percent increase in water absorption was calculated as (W2h − Wo)/Wo × 100, where W2h is the weight of the scaffold after 2 h and Wo is the original weight of the scaffold at time zero. The mechanical properties of the scaffolds were evaluated in the dry state using a tensile test. The scaffolds were cut into small strips (5 × 15 × 2 mm), and five samples were taken from different sites for each scaffold. The rectangular specimens were measured using a digital calliper micrometer (Ultracal III; Sylvac, Bern, Switzerland). To obtain the mean size, we measured and averaged three different sections of each specimen. The tensile test was conducted using a universal tensile machine (Top-tech 2000; Chemilab, Suwon, South Korea). The stress−strain curves of the scaffolds were recorded at a stretching 4257
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Figure 1. Fabrication process for a highly roughened 3D PCL scaffold. (a) Schematic diagrams of an EHD process with a grounded PEO solution bath. (b) Stable cone-jet mode showing an ejected solution line due to the applied voltage (10 kV) between the nozzle tip and the grounded PEO solution (40 mm). SEM images of fabricated struts (c) with and (d) without a PEO solution bath. When using the grounded PEO bath, the surface of the fabricated strut was highly roughened and resembled a mass of electrosprayed particles. (e) Optical micrographs of the fabricated scaffold (80 × 80 × 2 mm3) (up) and side view of the scaffold (down). Alizarine Red-S (ARS) Staining and Qualification of Calcium Deposit. The level of mineralization was determined by ARS using six-well plates for 7 and 14 days. MG63 cells were cultured in DMEM containing 50 μg mL−1 vitamin C and 10 mM β-glycerophosphate for 1 week in the presence of phlorotannins. Next, the cells were washed three times with PBS, fixed in 70% (v/v) cold ethanol (4 °C) for 1 h, and then air-dried. The ethanol-fixed specimens were stained with 40 mM ARS (pH 4.2) for 1 h and washed three times with purified water. The specimens were destained with 10% cetylpyridium chloride in 10 mM sodium phosphate buffer (pH 7.0) for 15 min. The optical density (OD) was measured at 562 nm using a Spectra III UV microplate reader. Statistical Analysis. All quantitative results were obtained from five samples. All data presented are expressed as the mean ± SD. Statistical analyses consisted of single-factor analyses of variance (ANOVAs). The significance level was set at p < 0.05.
the apex of the cone was reduced until the Maxwell stress equaled the maximum capillary stress, resulting in charged-fluid jet ejection.32 To obtain a stable cone-jet mode, the applied voltage was appropriately controlled between dripping mode and multijet mode.33 Because voltages greater than the Rayleigh limit can induce multijet mode, which can generate micro/ nanosized droplets/fibers, an appropriate voltage should be selected. Figure 1b shows a stable single jet obtained during EHD plotting under an applied voltage (10 kV) to a flowing medium (15 wt % PCL solution) at a fixed rate (1 mL h−1). To achieve stable 3D structures, we used a PEO solution (5 wt %) in the grounded bath so that the deposited struts did not crumble because of elastic cushioning of the PEO solution target (Figure 1c). The deposited struts can crumble on the solid target because of the high electrostatic force between the nozzle and the target (Figure 1d). When EHD-plotted struts are soaked in a grounded bath, remnant DMF and MC within the struts can induce interadhesion between the struts. Adhesion between struts, facilitated by remnant solvents, improved because of the high viscosity of the target PEO solution. In a pure water bath, adhesion decreased because of rapid sedimentation of the struts. This phenomenon can be explained by Fick’s second law, which states that the mass loss of remnant solvents in the struts as a function of time is an exponential decay curve, from which the interadhesion between the deposited struts can be calculated. Therefore, an appropriate amount of remnant solvent is required. However, if sedimentation occurs rapidly in a water bath, the struts do not
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RESULTS AND DISCUSSION Figure 1a shows a schematic diagram of our EHD-plotting system. A syringe pump (flow rate: 1 mL h−1) connected to a stainless-steel nozzle (i.d., 510 μm; o.d., 800 μm) delivered the PCL solution to the nozzle tip. The stainless-steel nozzle was used as an anode by connecting it to a power supply to form an electric field with a grounded electrode. A three-axis robot moving stage was controlled by a computer-aided design (CAD) system. By applying voltage to the nozzle, a spherical meniscus of polymer solution at the nozzle tip transformed slowly into a cone known as a Taylor cone31 due to accumulation of surface charges.32 The radius of curvature at 4258
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Figure 2. (a−c) Solution meniscus at the nozzle tip for various applied voltages (8, 10, 15 kV) under the same flow rate (1 mL h −1) and weight fraction (15 wt %) of PCL solution. (d) Profiles of the strain rate for applied voltages (V a). The inset shows measured mean diameters (dl) of liquid jets from the apex of the Taylor cone. (e) Effect of applied voltage (V a) and weight fraction (10, 15, and 20 wt %) of PCL on the measured strut diameter (ds) of the scaffold.
have enough time to adhere to each other, which is one of the reasons we used PEO in the grounded bath. In our previous work, a similar process using electrostatic force and a solution bath as a target was used to fabricate hydrophobic surfaces and 3D nanofiber structures.34,35 Another reason we used PEO solution as a target was that by plotting the PCL struts in a bath, electric interruption occurred because the PCL struts were soaked in solution. Therefore, in low-viscosity solution, EHD-plotted struts moved easily. If unstable movement of the struts occurs in the bath, then fabricating stable layer-by-layer 3D structures is very difficult. However, because we used a PEO solution with an appropriate viscosity, the movement of the struts was reduced dramatically. Figure 1e shows a final PCL scaffold with dimensions of 80 × 80 × 2 mm. Figure 2a−c shows that the meniscus at the nozzle tip contracted as the electric field increased under a constant flow rate (1 mL h−1) of PCL solution. The elongational deformation of the jet exhibited exponential growth;36 thus, as shown in the insets of Figure 2c, the diameter (d1) of the jet ejected under the same flow rate and weight fraction (15 wt %) of PCL solution decreased exponentially, and the rate of reduction accelerated with increasing applied voltage (Va). As shown in Figure 2c, the elongational strain rate, (1/vo)(dvx/dx), of the x axis was analyzed using the diameter change of the jets. The strain−rate curve showed a characteristic bell shape with a maximum near the nozzle tip at the deformation region along the jet line. The maximum point increased and shifted to the nozzle tip as the applied voltage increased. This phenomenon is similar to the general melt-spinning process for increasing uptake speed. Figure 2d shows an operating diagram delineating the effects of electric field (Va) and polymer weight fraction on the diameter (ds) of the final struts in the EHD-plotting system. The breakup voltages, which resulted in electrospray or
electrospinning, for various weight fractions of PCL were measured under the same flow rate. For 10, 15, and 20 wt % PCL solutions, the breakup voltages were 15.5, 16.7, and 18.5 kV, respectively. Although the flow rate to the nozzle was the same, the voltage for the viscous solution can be expressed as V ≈ η1/2, where V and η are the breakup voltage and viscosity, respectively.37 As shown in Figure 2d, the strut diameter in the high-viscosity (20 wt %) PCL solution was smaller than that in the low-viscosity (10 and 15 wt %) PCL solutions because the breakup voltage, generating multijets, in the high-viscosity solution was higher. At voltages greater than the breakup voltage, the strut width was large because the electrosprayed droplets spread out due to electrostatic repulsion between droplets. In general, EHD plotting has been used to obtain fine electronic circuits and 2D coating surfaces for biomedical applications.38,39 However, in tissue engineering, biomedical scaffolds should have a 3D, rough, porous structure providing physical transport of both nutrients and biological agents from the surface through the width of the 3D scaffold.40 To fabricate an EHD-plotted 3D structure for tissue engineering, we used 12 wt % PCL and applied a voltage of 2.5 kV cm−1. Detailed experimental conditions are described in the Methods section. Figure 3a shows SEM images of the EHD-plotted struts using the conventional method. Struts with a width distribution of 253 ± 27 μm exhibited layer-by-layer construction but crumbled due to remnant solvent within the plotted struts. The shrunken shape of the struts was related to the evaporation rate (Peclet number) of the solvent used in the process.41,42 This phenomenon was accelerated under higher electric fields due to increased electrostatic forces between the positively charged solution and the solid target. Additionally, as the electric field increased beyond the breakup voltage, a mixture of struts and electrospun fibers was observed (Figure 4259
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that the surface of the struts resembled a lump of slightly dissolved droplets (mean size = 11 ± 3 μm). On the basis of surface characterization, this scaffold is likely to improve cell attachment and proliferation significantly compared with general rapid-prototyped scaffolds. Figure 3d,e shows optical images of surface and cross-section of the finally fabricated scaffold (60 × 60 × 3 mm3). The ability of scaffolds to absorb water is an important factor affecting cell distribution, cellular proliferation, and structural morphology of regrown tissues.43 Water-absorption capacity was calculated by weighing the scaffolds before and after soaking in distilled water for 2 h. To compare the efficiency of the EHD-plotted PCL (EHD-PCL) scaffold, we used a control sample fabricated using a rapid-prototyped PCL (RP-PCL) scaffold with a pore size of 205 ± 27 μm and strut diameter of 220 ± 25 μm. The fabrication procedure was the same as that used in our previous work,14,44 and the fabrication parameters are listed in Table 1. Figure 4a compares the relative water-absorption capacity of RP-PCL and EHD-PCL scaffolds. The EHD-PCL scaffold had a much higher water-absorption capacity than the RP-PCL scaffold, even though the same material (PCL) was used to fabricate both scaffolds. This phenomenon was due to differences between the physical structures of the two scaffolds: a high porosity and slightly offset structure of the EHD-PCL struts compared with the smooth, 3D structure of the RP-PCL struts. The results were validated by time-lapse images showing the absorption of one drop of water (5 μL) mixed with purified water and red dye (Figure 4b−g). The mechanical properties of biomedical scaffolds for bone tissue regeneration should be strong enough to sustain new bone tissue regeneration because the damaged area can be stressed by surrounding native tissue.38 However, because of high porosity, which is required in 3D scaffolds to lead direct osteogenesis,3 some mechanical strength is sacrificed. Therefore, the relationship between porosity and mechanical properties should be appropriately balanced according to the required mechanical and biological functions. Therefore, control of porosity and mechanical properties is important when designing scaffolds.40 Typical stress (σ)−strain (ε) curves for scaffolds (RP-PCL and EHD-PCL) with different porosities (56 and 78%) at a constant stretching velocity 0.2 mm s−1 are shown in Figure 4h. In the RP-PCL scaffold, the curve fluctuated due to break-up of the interlocked struts during the
Figure 3. SEM images of fabricated PCL scaffolds. (a) SEM micrographs of the scaffold fabricated by a normal EHD-printing method showing the pore size (347 ± 86 μm) and strut size (205 ± 61 μm). (b) SEM images showing a mixture of struts and fibers due to the high electric field. (c) SEM images showing the highly roughened surface morphology of the EHD-plotted 3D scaffold on the PEO solution bath. Optical images of (d) surface and (e) cross-section of a fabricated EHD-PCL scaffold (60 × 60 × 3 mm3).
3b). However, our goal was to obtain 3D scaffolds using our modified EHD-plotting process. Using a previous process, we applied the same electric field, but the target was replaced with a grounded PEO solution bath. Figure 3c shows SEM images of the structures fabricated using this process. As shown in the Figure, the strut did not crumble, and the surface was rough, enhancing initial cell attachment. The magnified image shows
Figure 4. (a) Relative water-absorption capacities of RP-PCL and EHD-PCL scaffolds (n = 5). Time-lapse images of water absorbed from one drop of water (5 μL) mixed with purified water and red dye for (b−d) RP-PCL and (e−g) EHD-PCL. (h) Stress (σ)−strain (ε) curves of RP-PCL (porosity, ϕ: 56%) and EHD-PCL (porosity, ϕ: 78%) scaffolds with a constant stretching velocity (0.2 mm s −1) (n = 5). *p < 0.05 indicates a significant difference. 4260
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Figure 5. SEM micrographs of MG63 cells cultured on (a,b) an EHD-PCL scaffold and (c) RP-PCL scaffold after 3 days of cell culture. (d) Cell viability is indicated by the MTT assay of MG63 cells seeded on the scaffolds (n = 5). (e) Alkaline phosphatase (ALP) activity of MG63 cells on RPPCL and EHD-PCL scaffolds from 5 to 10 days (n = 5). (f) Calcium mineralization for the scaffolds from 7 to 14 days (n = 5). (g,h) Optical micrographs of ARS-staining indicating mineralization for the RP-PCL and EHD-PCL scaffold after 7 days. *p < 0.05 indicates a significant difference.
tensile test. In the EHD-PCL scaffold, the curve was a smooth line due to homogeneous stress distribution over the crosssectional area. As shown in the curves, the Young’s modulus (RP-PCL: 12.5 ± 2.3 MPa) of the low-porosity (56%) scaffold was higher than that (EHD-PCL: 6.3 ± 1.8 MPa) of the highporosity (78%) scaffold. The relationship between porosity and modulus was evaluated using an equation in which pores are assumed to be cubic:45 E(ϕ) = Eo(1 − ϕ2/3), where Eo and ϕ are the Young’s modulus of the pore-free materials and porosity, respectively. The PCL modulus with no porosity was 41 ± 2.1 MPa. Using the equation, the calculated moduli for RP-PCL and EHD-PCL were 13.1 and 6.3 MPa, respectively. We concluded that the relatively low modulus of EHD-PCL compared with that of RP-PCL was due to the higher porosity of the EHD-PCL scaffold. In this study, our regenerating target is trabecular bone, which has a modulus from 38 to 130 MPa. The value is highly dependent on the direction of the applied force and the inherent structure.46 The measured modulus of EHD-PCL scaffold was relatively low compared with the modulus of trabecular bone. However, the mechanical properties might be improved by incorporating various composite systems containing synthetic polymers or ceramics, such as hydroxyapatite and beta-tricalcium phosphate. Future studies will focus on strengthening the structured scaffolds with various supporting information. To assess the cellular response of the scaffolds fabricated via EHD plotting, we used osteoblast-like cells (MG63). Although the optimal pore size of a scaffold for bone tissue regeneration has been debated, the pore size of EHD-PCL was obtained as 347 ± 86 μm because the optimal pore size of bone scaffolds is recommended to be between 100 and 400 μm for osteoconduction.47 To observe interactions between cells (MG63) and scaffolds, we acquired SEM and fluorescence images 3 days after cell culture (Figure 5a−c). The images indicated that the
proliferated osteoblast-like cells adhered not only on the strut surface but also on the small regions consisting of struts, indicating that the EHD-PCL scaffold has a higher cell affinity than the RP-PCL scaffold due to the rough layer-by-layer structure. Cell viability was evaluated based on the number of viable cells that adhered and proliferated on the scaffolds after 1, 3, and 7 days. Figure 5d shows that the cell-viability differed significantly on the RP-PCL and EHD-PCL scaffolds for all intervals. The 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) assay results indicated that the EHDplotted scaffold had an enormous increase in cell viability compared with the control scaffold due to the cell-attractive surface structure. These results were consistent with previous research reporting a correlation between surface roughness and cellular behavior. According to several reports, 48,49 the proliferation of rat calvarial osteoblasts increased significantly with increasing roughness of an epoxy sputter-coated with a 60 nm layer of titanium, and the number of cells increased when the roughness parameter, which was calculated with a method (DIN EN ISO 4288-98), was >2 μm. ALP activity was measured to observe the effect of the EHDPCL scaffold structure. According to Lee et al., the responses between the surface morphology of scaffold and cellular behavior are closely related with a complex biological system that includes protein adsorption, receptor−ligand binding, and signal transduction.50 As shown in Figure 5e, ALP activity increased with time. After 5 days of incubation, the ALP activity of the EHD scaffold was much higher than that of the RP-PCL scaffold. After 10 days of incubation, the ALP activity of the EHD-PCL scaffold was three times that of the control scaffold. This result indicates that the electrohydrodynamically designed structures played an important role in supporting ALP activity, even though the same material was used for both scaffolds. Alizarine red-S (ARS) was used to determine the amount of calcium mineral. The OD of ARS extracted from stained 4261
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cultures is shown in Figure 5f. The intensity after 14 days increased considerably compared with that after 7 days. Additionally, the EHD scaffold had higher OD intensity compared with the RP-PCL scaffold. The ALP activity and ARS results can be compared with the data of cell viability (MTT assay). When comparing the results, the ALP activity (Figure 5e) and mineralization (Figure 5f) were highly correlated with the cell viability, and the increased ALP activity of Figure 5e and mineralization of Figure 5f were only because of the highly proliferated cells. Figure 5g,h shows optical images of ARS staining for the RP-PCL and EHD-PCL scaffolds after 7 days of cell culture. In the images, red indicates calcium mineral: red fully covered the struts and even pores of the EHD-PCL scaffold compared with that of RP-PCL. Therefore, we concluded that the EHD-PCL scaffolds are a potential biomaterial for bone tissue regeneration.
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CONCLUSIONS In summary, although EHD printing has been widely used in designing micro/nanosized electronics and coating surfaces of various biomaterials in cone-jet mode, biomedical applications of scaffolds have been limited because EHD-printed struts have substantial charges, which are problematic during layered deposition of struts, and remnant solvents can dissolve previously deposited layers of struts, crushing the 3D structure. In tissue engineering, a 3D scaffold structure is necessary to provide spatial paths for transport of nutrients and injected cells. The primary advantage of our modified EHD-plotting method over conventional EHD printing is the fabrication of a 3D structure with a rough surface. Although stable size control of micro/nanosized struts remains a challenge, we expect that control of process parameters (e.g., applied electric field, viscosity of target solution, reduction in the nozzle dimensions) will enable control of strut size to fabricate more stable 3D structures. We believe that the modified EHD process may be one of promising methods for the fabrication of bone scaffold due to significantly enhanced dramatic cell viability and high mineralization in the EHD-plotted PCL scaffold.
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AUTHOR INFORMATION
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REFERENCES
Corresponding Author *Tel: +82-62-230-7180. Fax: +82-62-236-1534. E-mail: gkim@ chosun.ac.kr.
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dx.doi.org/10.1021/bm201126j | Biomacromolecules 2011, 12, 4256−4263
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dx.doi.org/10.1021/bm201126j | Biomacromolecules 2011, 12, 4256−4263
