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Functionalization of poly(caprolactone) scaffolds by the surface mineralization for use in bone regeneration
Materials Letters 65 (2011) 3559–3562
Contents lists available at ScienceDirect
Materials Letters j o u r n a l h o m e p a g e : w w w. e l s ev i e r. c o m / l o c a t e / m a t l e t
Functionalization of poly(caprolactone) scaffolds by the surface mineralization for use in bone regeneration Biligzaya Dorj a, b, 1, Mi-Kyung Kim a, b, 1, Jong-Eun Won a, b, Hae-Won Kim a, b, c,⁎ a b c
Department of Nanobiomedical Science & WCU Research Center, Dankook University Graduate School, South Korea Institute of Tissue Regeneration Engineering (ITREN), Dankook University, South Korea Department of Biomaterials Science, School of Dentistry, Dankook University, South Korea
a r t i c l e
i n f o
Article history: Received 6 June 2011 Accepted 29 July 2011 Available online 4 August 2011 Keywords: Biopolymer Surface treatment Biofunctional scaffold Protein loading Mineralization
a b s t r a c t The surface of a synthetic biopolymer scaffold was tailored by a mineralization with calcium phosphate for use as a functional bone tissue engineering matrix. Poly(ε-caprolactone) scaffold with a defined pore configuration constructed by a robocasting method was treated in a series of solutions involving steps of surface activation and calcium phosphate induction. The scaffold surface was completely covered with calcium phosphate nanocrystallites that had typical characteristics of bone mineral-like carbonate apatite. The scaffold with mineralized-surface demonstrated to support more favorable bone cell responses, including initial cell adhesion and proliferation and to allow higher loading of protein than the untreatedscaffold. The results suggest the developed scaffold has the potential for use as a bone regenerative matrix. © 2011 Elsevier B.V. All rights reserved.
1. Introduction Scaffolds with three-dimensional (3D) pore structure have played vital roles in the biomedical field, such as tissue reconstruction and cell engineering. The interconnected pore space of the scaffolds allows the supply of nutrients and cells and aids their population and tissuespecific development [1]. Therefore, the 3D pore configuration is considered to be of special importance in the success of the scaffold applications [2]. Among other methods used to develop porous scaffolds, robocasting, a class of rapid prototyping techniques, is a facile tool used to produce materials with pore structures tuned to the designer specifications [2–4]. Robocasting scaffolds made from degradable polymers have been found to have pores with complete interconnectivity and 3D pore configurations tuned to the target tissues [4]. In addition to the pore structure, the surface of scaffolds should be considered because it is the first ligand of materials that allows tissue cells to recognize the anchorage site and further them to spread and multiply [5,6]. In practice, the surface properties of synthetic biopolymers including poly(ε-caprolactone) (PCL), which have been widely used as tissue scaffolds, must be improved because they generally are hydrophobic and have poor cell affinity [7,8] Surface treatments of these synthetic polymers include coating with hydrophilic polymers and cell-adhesive proteins [9,10]. Otherwise, bioac-
⁎ Corresponding author. Tel.: + 82 41 550 3081; fax: + 82 41 550 3085. E-mail address: [email protected] (H.-W. Kim). 1 Both authors contributed equally to this work. 0167-577X/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.matlet.2011.07.110
tive inorganics have been incorporated within the compositions to improve the biological affinity of the polymeric scaffolds [11]. In this study, we attempted to improve the surface properties of PCL scaffold targeting bone tissue, which was ultimately useful as a degradable and bioactive 3D matrix for bone regeneration. For this purpose, the surface was tailored with a bone mineral-like apatite. The scaffold was first produced to retain a well-controlled pore dimension using a robocasting method, after which it was treated with a series of solutions to induce surface mineralization. The efficacy of the surface-mineralized 3D scaffold was investigated in terms of the bone cell adhesion and proliferation as well as the loading of biological protein.
2. Experimental PCL (MW = 80,000, Sigma-Aldrich) solution was prepared by dissolving in acetone (PCL/acetone= 25 wt.%) at 50 °C. A robocasting machine (Ez-ROBO3, Iwashita, Japan) was used to produce 3D PCL scaffolds. Specifically, the solution was added to a syringe that was maintained at 50 °C using a heating jacket, after which it was injected through a 520 μm diameter needle using a force-controlled plunger to regulate a mass flow rate. To guide the fiber dispensing path, a positional control unit was operated and the speed of the dispensing fiber was set as 10 mm/s. The fiber was deposited into a cooled ethanol bath to effectively solidify the PCL solution. Through layer-by-layer stacking of the 2D fiber configuration, a 3D porous scaffold was produced with a controlled pore size and geometry. The scaffold was then post-treated at 50 °C to aid contact fusion of the fibers.
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For the mineralization of the scaffold, the surface was activated by dipping in 1 N NaOH for 1 h while gently stirring, followed by washing in distilled water. Next, the scaffold was dipped alternatively into 150 mM CaCl2 and 150 mM Na2HPO4 solution five times to induce CaP nucleation. Between the alternative dipping processes the scaffold was fully washed. The CaP-nucleated scaffold was further treated in a mineralization solution (1.5 times of simulated body fluid (1.5SBF), with 213.0 mM Na+, 7.5 mM K +, 2.25 mM Mg2+, 3.75 mM Ca2+, 221.7 mM Cl−, 6.3 mM HCO3−, 1.5 mM HPO42− and 0.75 mM SO42−) at 37 °C for up to 14 days. The 1.5SBF was used after the sterilization by filtering through a sterilization filter paper (Milipore, 0.2 μm pore size). The pore morphology of the scaffold and its surface microstructure were then observed by scanning electron microscopy (SEM; Hitachi S-3000H) and the mineralized composition was characterized by energy dispersive spectroscopy (EDS, Bruker SNE-3000 M). The weight change of the scaffolds according to the mineralization was measured. The chemical bond status of the mineralized surface was examined by Fourier transform infrared (FTIR; Jasco 470 PLUS) spectroscopy. Tissue cell responses to the mineralized PCL scaffold were investigated in terms of the adhesion and proliferation of bone-associated cells (MC3T3-E1, ATCC). Prior to cell tests, samples were sterilized with 70% ethanol for 10 min and then dried overnight under a laminar flow. Cells were seeded onto each scaffold (PCL mineralized or not, 10 mm ×10 mm×3 mm) in a regular culture medium (α-modified minimal essential medium supplemented with fetal bovine serum containing 1% antibiotic/antimycotic solution, without the use of osteogenic factors). To aid cellular adhesion to the scaffold, a 50 μl aliquot of cells (5×104 cells) was soaked into each scaffold which was left in an incubator (humidified atmosphere of 5% CO2 at 37 °C) for 6 h, and then 150 μl culture medium was added. The cell-seeded scaffolds were cultured for further up to 7 days. The cell growth morphology was then examined by SEM after dehydration and fixation of the cells. The cell proliferation was quantified based on the mitochondrial activity using an MTS (3-(4,5dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)2H-tetrazolium) assay kit (CellTiter 96 ®Queous One Solution Cell Proliferation Assay, Promega) [12]. The viable cells were also counted using a hemocytometer after trypan blue screening. Statistical analysis of the cell data was made by Student's t-test and significance was considered at pb 0.05.
The protein adsorption test onto the scaffolds mineralized or not was performed using cytochrome C as the model protein. Protein solution was prepared in 10 ml PBS at 100 μg/ml. Each scaffold (6 mm × 6 mm × 3 mm) was dipped into the solution and left at 20 °C while gentle shaking for various periods (up to 17 h) to allow the protein adsorption. The protein quantity that remained in the solution was assessed using UV–VIS spectroscopy. The absorbance measured at 409 nm was converted to the protein quantity based on the standard curve. 3. Results and discussion The 3D polymeric scaffold used for the surface mineralization was produced using the robocasting method. This method is a rapid prototyping tool that has the potential to construct 3D pore geometry tuned to the defect size and shape with the assistance of computerized design and operation, as fully described in the Experimental section. The representative morphology of the scaffold is shown in Fig. 1(a,b). Well-constructed scaffolds with 3D geometry (average fiber diameter of 245 μm, average pore size of 346 μm, and porosity of ~78%) were obtained using this method. The surface of the PCL scaffold was shown to be clean and the fiber stems were welded together to provide a stable 3D structure. The surface of the robocasting PCL scaffold was shown to be highly hydrophobic, making it difficult for water spreading and tissue cell adhesion [8]; therefore, we tailored the surface to be more hydrophilic for use as the 3D matrix for bone tissue regeneration. Calcium phosphate (Ca) mineralization of the surface was conducted though a series of solution-based treatments. To mineralize the scaffold, the surface was first activated by immersion in an alkaline solution (1 N NaOH) to reveal a number of carboxyl and hydroxyl groups [5]. The activated surface was further dipped in sequence within the highly concentrated calcium- and phosphate-containing medium several times (here 5 times used), which was followed by incubation in calcification medium composed of 1.5 times simulated body fluid (1.5SBF). No obvious change was observed on the surface of the scaffold after the NaOH treatment, while the alternative soaking process resulted in the formation of some aggregates of CaP. These aggregates should play a role in enhancing the mineralization process in the following step of
Fig. 1. (a,b) Morphology of the PCL scaffolds produced by a robotic dispensing method, featuring a well-constructed pore geometry. (c,d,e) SEM images of the PCL scaffolds after the mineralization process in 1.5SBF for (c) 1, (d) 3 and (e) 7 days. After 3 days, the surface was covered completely with the nanocrystalline apatite mineral, while prolonged immersion for 7 days resulted in thick layer and severe cracking after drying. (f) Weight increase of the scaffold due to mineral formation, and (g) FT-IR chemical analysis of the scaffolds before and after the mineralization.
B. Dorj et al. / Materials Letters 65 (2011) 3559–3562
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Fig. 2. Bone cell responses to the mineralized PCL scaffold: (a,b) cell growth morphology on the mineralized PCL for 3 days of culturing at low and high magnification, and (b) MTS cell proliferation rate measured for up to 7 days, showing significantly enhanced cell growth on the mineralized scaffold (significance level *p b 0.05).
1.5SBF incubation. A special care needs to eliminate excess calciummedium prior to dipping the scaffold into the phosphate-medium and vice versa because this helps produce homogeneous CaP nucleation and prevent the formation of uneven CaP formation such as clusters. As shown in Fig. 1(c–e), further soaking in the pseudo-saturated Ca–Psolution (1.5SBF used herein) finally accelerates the mineralization process, resulting in the considerable formation of CaP crystals at day 1 (Fig. 1(c)) and complete coverage of the surface with minerals at 3 days (Fig. 1(d)). The magnified image in the inset showed the formation of bone mineral-like nanocrystallites. However, a prolonged incubation for more than 7 days resulted in severe crack formation after drying, indicating this is an excessive stage of the mineralization (Fig. 1(e)). EDS revealed that the Ca/P ratio of the mineral product present on the scaffold surface was ~1.57, which is similar but slightly lower than that of stoichiometric hydroxyapatite, suggesting calcium-deficient bonemineral like apatite. The weight increase associated with the mineral formation was monitored during the SBF immersion (Fig. 1(f)). The increase was continuous with immersion time, resulting in ~3, 7 and 11% of the initial weight, at 1, 2 and 3 weeks of immersion, respectively. The mineralized surface was characterized with FTIR (Fig. 1(g)). Compared with the bare PCL scaffolds, where only the chemical bands related to PCL appeared, the scaffold mineralized for 3 days showed big difference, presenting a sharp phosphate band at ~1020 cm− 1 as well as carbonate bands at 1455 and 870 cm− 1, confirming the apatite is carbonated. The trend was almost similar in the scaffold mineralized for 7 days.
Cytochrome C loading (µg)
20 18 16 Mineralized scaffold
14 12 10
The bone-mineral like nanocrystalline phase produced on the scaffold surface is considered to provide favorable substrate conditions for the cell adhesion and growth, a required initial stage for bone tissue engineering. The in vitro pre-osteoblast responses to the mineralized PCL scaffold were investigated, as shown in Fig. 2. The SEM cell morphology on the mineralized PCL scaffold at the growth stage (3 days) was taken as the representative sample. Cells were shown to have good adhesive contacts with the substrate and a number of protruding filopodia (Fig. 2(a,b)). Quantification of the cell proliferation based on the mitochondrial activity (Fig. 2(c)) revealed significantly enhanced cell growth potential on the mineralized PCL scaffold when compared to that on the bare scaffold, particularly at days 3 and 7 (*pb 0.05). The cells were highly viable during the growth stage; when measured by using a hemocytometer after the trypan blue screening, the cell viability was about 90–110% with respect to culture dish control. Based on these cellular responses, the surface mineralization of the PCL scaffold is considered to be highly favorable for supporting cell adhesion and growth and can thus be used as a bone tissue engineering scaffold. Another merit of the mineralized scaffold was found in the significant improvement of the loading capacity of the biological proteins. The utilization of biological proteins with specific functions such as growth factors in concert with the bone scaffolds is considered potential strategy to improve the regeneration ability of the defective bone [13,14]. Here we chose one model protein, cytochrome C, and its loading amount onto the scaffold was monitored with time. As shown in Fig. 3, the mineralized scaffold showed significantly higher loading quantity than the base scaffold (approximately 2.5 times higher); the maximal loading was ~17.7 μg for the mineralized and ~7.4 μg for the bare scaffold. After the maximal loading the protein appeared to be released from the scaffold as deduced from the decrease in the protein quantity with increasing time. In fact, cytochrome C is known to have similar characteristics such as molecular weight and surface charge with respect to growth factors, mainly basic fibroblast growth factor (bFGF) [15], therefore this study may be useful in designing bone scaffolds with delivering of bFGF to elicit therapeutic effects such as the increase in cell proliferative potential and blood vessel formation for bone tissue engineering, which remains as further study.
8 6 Bare scaffold
4 2 0 0
2
4
6
8
10
12
14
16
18
Loading time (h) Fig. 3. Protein (cytochrome C) loading capacity of the mineralized PCL scaffold in comparison with that of bare scaffold. The maximal loading amount was about 2.5 times higher in the mineralized scaffold.
4. Conclusions The surface of the PCL biopolymer scaffold constructed using a robocasting method was mineralized with carbonate apatite nanocrystallites by a series of solution-based treatments. Tissue cells actively adhered and proliferated on the surface-mineralized scaffold, which was in direct contrast with the poor cell affinity of the untreated PCL scaffold. Moreover, the surface mineralization facilitated the scaffold to hold significantly enhanced protein loading capacity. The study
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demonstrates the surface-mineralized polymer scaffolds are potentially useful in bone tissue engineering. Acknowledgment This work was supported by Priority Research Centers Program (grant#: 2009–0093829) and WCU (World Class University) program (grant#: R31-10069) through the National Research Foundation (NRF) funded by the Ministry of Education, Science and Technology. Authors thank the support of IBST in Dankook University.
[4] [5] [6] [7] [8] [9] [10] [11] [12] [13]
References [1] Lu L, Mikos A. MRS Bull 1996;21:28. [2] Hollister SJ, Maddox RD, Taboas JM. Biomaterials 2002;23:4095. [3] Chua CK, Leong KF, Lim CS. Rapid prototyping2nd ed. . Singapore: World scientific; 2003.
[14] [15]
Hutmacher DW, Sittinger M, Risbud MV. Trends Biotechnol 2004;22:354. Yu HS, Jang JH, Kim TI, Lee HH, Kim HW. J Biomed Mater Res A 2009;88:747. Stevens MM, George JH. Science 2005;310:1135. Kimura Y. Biodegradable polymers. In: Tsuruta T, editor. Biomedical applications of polymeric materials. Boca Raton: CRC press; 1993. p. 163. Hong SJ, Jeong I, Noh KT, Yu HS, Lee KS, Kim HW. J Mater Sci Mater Med 2009;20: 1955. Desmet T, Morent R, Geyter ND, Leys C, Schacht E, Dubruel P. Biomacromolecules 2009;10:2351. Kim JE, Noh KT, Yu HS, Lee HY, Jang JH, Kim HW. Adv Eng Mater 2010;12:94. Koh YH, Jun IK, Kim HW. Mater Lett 2006;60:1184. Oh SA, Kim SH, Won JE, Kim JJ, Shin US, Kim HW. J Tissue Eng 2011;2010:10 (Article ID 475260). Yun YR, Won JE, Jeon E, Lee S, Kang W, Jo H, et al. J Tissue Eng 2010;2010:18 (Article ID 218142). Young CS, Hart BG, Karunanidhi A, Street RM, Schutte L, Hollinger JO. J Tissue Eng 2010;2010:8 (Article ID 246215). Tsurushima H, Marushima A, Suzuki K, Oyane A, Sogo Y, Nakamura K, et al. Acta Biomater 2010;6:2751.
Contents lists available at ScienceDirect
Materials Letters j o u r n a l h o m e p a g e : w w w. e l s ev i e r. c o m / l o c a t e / m a t l e t
Functionalization of poly(caprolactone) scaffolds by the surface mineralization for use in bone regeneration Biligzaya Dorj a, b, 1, Mi-Kyung Kim a, b, 1, Jong-Eun Won a, b, Hae-Won Kim a, b, c,⁎ a b c
Department of Nanobiomedical Science & WCU Research Center, Dankook University Graduate School, South Korea Institute of Tissue Regeneration Engineering (ITREN), Dankook University, South Korea Department of Biomaterials Science, School of Dentistry, Dankook University, South Korea
a r t i c l e
i n f o
Article history: Received 6 June 2011 Accepted 29 July 2011 Available online 4 August 2011 Keywords: Biopolymer Surface treatment Biofunctional scaffold Protein loading Mineralization
a b s t r a c t The surface of a synthetic biopolymer scaffold was tailored by a mineralization with calcium phosphate for use as a functional bone tissue engineering matrix. Poly(ε-caprolactone) scaffold with a defined pore configuration constructed by a robocasting method was treated in a series of solutions involving steps of surface activation and calcium phosphate induction. The scaffold surface was completely covered with calcium phosphate nanocrystallites that had typical characteristics of bone mineral-like carbonate apatite. The scaffold with mineralized-surface demonstrated to support more favorable bone cell responses, including initial cell adhesion and proliferation and to allow higher loading of protein than the untreatedscaffold. The results suggest the developed scaffold has the potential for use as a bone regenerative matrix. © 2011 Elsevier B.V. All rights reserved.
1. Introduction Scaffolds with three-dimensional (3D) pore structure have played vital roles in the biomedical field, such as tissue reconstruction and cell engineering. The interconnected pore space of the scaffolds allows the supply of nutrients and cells and aids their population and tissuespecific development [1]. Therefore, the 3D pore configuration is considered to be of special importance in the success of the scaffold applications [2]. Among other methods used to develop porous scaffolds, robocasting, a class of rapid prototyping techniques, is a facile tool used to produce materials with pore structures tuned to the designer specifications [2–4]. Robocasting scaffolds made from degradable polymers have been found to have pores with complete interconnectivity and 3D pore configurations tuned to the target tissues [4]. In addition to the pore structure, the surface of scaffolds should be considered because it is the first ligand of materials that allows tissue cells to recognize the anchorage site and further them to spread and multiply [5,6]. In practice, the surface properties of synthetic biopolymers including poly(ε-caprolactone) (PCL), which have been widely used as tissue scaffolds, must be improved because they generally are hydrophobic and have poor cell affinity [7,8] Surface treatments of these synthetic polymers include coating with hydrophilic polymers and cell-adhesive proteins [9,10]. Otherwise, bioac-
⁎ Corresponding author. Tel.: + 82 41 550 3081; fax: + 82 41 550 3085. E-mail address: [email protected] (H.-W. Kim). 1 Both authors contributed equally to this work. 0167-577X/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.matlet.2011.07.110
tive inorganics have been incorporated within the compositions to improve the biological affinity of the polymeric scaffolds [11]. In this study, we attempted to improve the surface properties of PCL scaffold targeting bone tissue, which was ultimately useful as a degradable and bioactive 3D matrix for bone regeneration. For this purpose, the surface was tailored with a bone mineral-like apatite. The scaffold was first produced to retain a well-controlled pore dimension using a robocasting method, after which it was treated with a series of solutions to induce surface mineralization. The efficacy of the surface-mineralized 3D scaffold was investigated in terms of the bone cell adhesion and proliferation as well as the loading of biological protein.
2. Experimental PCL (MW = 80,000, Sigma-Aldrich) solution was prepared by dissolving in acetone (PCL/acetone= 25 wt.%) at 50 °C. A robocasting machine (Ez-ROBO3, Iwashita, Japan) was used to produce 3D PCL scaffolds. Specifically, the solution was added to a syringe that was maintained at 50 °C using a heating jacket, after which it was injected through a 520 μm diameter needle using a force-controlled plunger to regulate a mass flow rate. To guide the fiber dispensing path, a positional control unit was operated and the speed of the dispensing fiber was set as 10 mm/s. The fiber was deposited into a cooled ethanol bath to effectively solidify the PCL solution. Through layer-by-layer stacking of the 2D fiber configuration, a 3D porous scaffold was produced with a controlled pore size and geometry. The scaffold was then post-treated at 50 °C to aid contact fusion of the fibers.
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B. Dorj et al. / Materials Letters 65 (2011) 3559–3562
For the mineralization of the scaffold, the surface was activated by dipping in 1 N NaOH for 1 h while gently stirring, followed by washing in distilled water. Next, the scaffold was dipped alternatively into 150 mM CaCl2 and 150 mM Na2HPO4 solution five times to induce CaP nucleation. Between the alternative dipping processes the scaffold was fully washed. The CaP-nucleated scaffold was further treated in a mineralization solution (1.5 times of simulated body fluid (1.5SBF), with 213.0 mM Na+, 7.5 mM K +, 2.25 mM Mg2+, 3.75 mM Ca2+, 221.7 mM Cl−, 6.3 mM HCO3−, 1.5 mM HPO42− and 0.75 mM SO42−) at 37 °C for up to 14 days. The 1.5SBF was used after the sterilization by filtering through a sterilization filter paper (Milipore, 0.2 μm pore size). The pore morphology of the scaffold and its surface microstructure were then observed by scanning electron microscopy (SEM; Hitachi S-3000H) and the mineralized composition was characterized by energy dispersive spectroscopy (EDS, Bruker SNE-3000 M). The weight change of the scaffolds according to the mineralization was measured. The chemical bond status of the mineralized surface was examined by Fourier transform infrared (FTIR; Jasco 470 PLUS) spectroscopy. Tissue cell responses to the mineralized PCL scaffold were investigated in terms of the adhesion and proliferation of bone-associated cells (MC3T3-E1, ATCC). Prior to cell tests, samples were sterilized with 70% ethanol for 10 min and then dried overnight under a laminar flow. Cells were seeded onto each scaffold (PCL mineralized or not, 10 mm ×10 mm×3 mm) in a regular culture medium (α-modified minimal essential medium supplemented with fetal bovine serum containing 1% antibiotic/antimycotic solution, without the use of osteogenic factors). To aid cellular adhesion to the scaffold, a 50 μl aliquot of cells (5×104 cells) was soaked into each scaffold which was left in an incubator (humidified atmosphere of 5% CO2 at 37 °C) for 6 h, and then 150 μl culture medium was added. The cell-seeded scaffolds were cultured for further up to 7 days. The cell growth morphology was then examined by SEM after dehydration and fixation of the cells. The cell proliferation was quantified based on the mitochondrial activity using an MTS (3-(4,5dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)2H-tetrazolium) assay kit (CellTiter 96 ®Queous One Solution Cell Proliferation Assay, Promega) [12]. The viable cells were also counted using a hemocytometer after trypan blue screening. Statistical analysis of the cell data was made by Student's t-test and significance was considered at pb 0.05.
The protein adsorption test onto the scaffolds mineralized or not was performed using cytochrome C as the model protein. Protein solution was prepared in 10 ml PBS at 100 μg/ml. Each scaffold (6 mm × 6 mm × 3 mm) was dipped into the solution and left at 20 °C while gentle shaking for various periods (up to 17 h) to allow the protein adsorption. The protein quantity that remained in the solution was assessed using UV–VIS spectroscopy. The absorbance measured at 409 nm was converted to the protein quantity based on the standard curve. 3. Results and discussion The 3D polymeric scaffold used for the surface mineralization was produced using the robocasting method. This method is a rapid prototyping tool that has the potential to construct 3D pore geometry tuned to the defect size and shape with the assistance of computerized design and operation, as fully described in the Experimental section. The representative morphology of the scaffold is shown in Fig. 1(a,b). Well-constructed scaffolds with 3D geometry (average fiber diameter of 245 μm, average pore size of 346 μm, and porosity of ~78%) were obtained using this method. The surface of the PCL scaffold was shown to be clean and the fiber stems were welded together to provide a stable 3D structure. The surface of the robocasting PCL scaffold was shown to be highly hydrophobic, making it difficult for water spreading and tissue cell adhesion [8]; therefore, we tailored the surface to be more hydrophilic for use as the 3D matrix for bone tissue regeneration. Calcium phosphate (Ca) mineralization of the surface was conducted though a series of solution-based treatments. To mineralize the scaffold, the surface was first activated by immersion in an alkaline solution (1 N NaOH) to reveal a number of carboxyl and hydroxyl groups [5]. The activated surface was further dipped in sequence within the highly concentrated calcium- and phosphate-containing medium several times (here 5 times used), which was followed by incubation in calcification medium composed of 1.5 times simulated body fluid (1.5SBF). No obvious change was observed on the surface of the scaffold after the NaOH treatment, while the alternative soaking process resulted in the formation of some aggregates of CaP. These aggregates should play a role in enhancing the mineralization process in the following step of
Fig. 1. (a,b) Morphology of the PCL scaffolds produced by a robotic dispensing method, featuring a well-constructed pore geometry. (c,d,e) SEM images of the PCL scaffolds after the mineralization process in 1.5SBF for (c) 1, (d) 3 and (e) 7 days. After 3 days, the surface was covered completely with the nanocrystalline apatite mineral, while prolonged immersion for 7 days resulted in thick layer and severe cracking after drying. (f) Weight increase of the scaffold due to mineral formation, and (g) FT-IR chemical analysis of the scaffolds before and after the mineralization.
B. Dorj et al. / Materials Letters 65 (2011) 3559–3562
3561
Fig. 2. Bone cell responses to the mineralized PCL scaffold: (a,b) cell growth morphology on the mineralized PCL for 3 days of culturing at low and high magnification, and (b) MTS cell proliferation rate measured for up to 7 days, showing significantly enhanced cell growth on the mineralized scaffold (significance level *p b 0.05).
1.5SBF incubation. A special care needs to eliminate excess calciummedium prior to dipping the scaffold into the phosphate-medium and vice versa because this helps produce homogeneous CaP nucleation and prevent the formation of uneven CaP formation such as clusters. As shown in Fig. 1(c–e), further soaking in the pseudo-saturated Ca–Psolution (1.5SBF used herein) finally accelerates the mineralization process, resulting in the considerable formation of CaP crystals at day 1 (Fig. 1(c)) and complete coverage of the surface with minerals at 3 days (Fig. 1(d)). The magnified image in the inset showed the formation of bone mineral-like nanocrystallites. However, a prolonged incubation for more than 7 days resulted in severe crack formation after drying, indicating this is an excessive stage of the mineralization (Fig. 1(e)). EDS revealed that the Ca/P ratio of the mineral product present on the scaffold surface was ~1.57, which is similar but slightly lower than that of stoichiometric hydroxyapatite, suggesting calcium-deficient bonemineral like apatite. The weight increase associated with the mineral formation was monitored during the SBF immersion (Fig. 1(f)). The increase was continuous with immersion time, resulting in ~3, 7 and 11% of the initial weight, at 1, 2 and 3 weeks of immersion, respectively. The mineralized surface was characterized with FTIR (Fig. 1(g)). Compared with the bare PCL scaffolds, where only the chemical bands related to PCL appeared, the scaffold mineralized for 3 days showed big difference, presenting a sharp phosphate band at ~1020 cm− 1 as well as carbonate bands at 1455 and 870 cm− 1, confirming the apatite is carbonated. The trend was almost similar in the scaffold mineralized for 7 days.
Cytochrome C loading (µg)
20 18 16 Mineralized scaffold
14 12 10
The bone-mineral like nanocrystalline phase produced on the scaffold surface is considered to provide favorable substrate conditions for the cell adhesion and growth, a required initial stage for bone tissue engineering. The in vitro pre-osteoblast responses to the mineralized PCL scaffold were investigated, as shown in Fig. 2. The SEM cell morphology on the mineralized PCL scaffold at the growth stage (3 days) was taken as the representative sample. Cells were shown to have good adhesive contacts with the substrate and a number of protruding filopodia (Fig. 2(a,b)). Quantification of the cell proliferation based on the mitochondrial activity (Fig. 2(c)) revealed significantly enhanced cell growth potential on the mineralized PCL scaffold when compared to that on the bare scaffold, particularly at days 3 and 7 (*pb 0.05). The cells were highly viable during the growth stage; when measured by using a hemocytometer after the trypan blue screening, the cell viability was about 90–110% with respect to culture dish control. Based on these cellular responses, the surface mineralization of the PCL scaffold is considered to be highly favorable for supporting cell adhesion and growth and can thus be used as a bone tissue engineering scaffold. Another merit of the mineralized scaffold was found in the significant improvement of the loading capacity of the biological proteins. The utilization of biological proteins with specific functions such as growth factors in concert with the bone scaffolds is considered potential strategy to improve the regeneration ability of the defective bone [13,14]. Here we chose one model protein, cytochrome C, and its loading amount onto the scaffold was monitored with time. As shown in Fig. 3, the mineralized scaffold showed significantly higher loading quantity than the base scaffold (approximately 2.5 times higher); the maximal loading was ~17.7 μg for the mineralized and ~7.4 μg for the bare scaffold. After the maximal loading the protein appeared to be released from the scaffold as deduced from the decrease in the protein quantity with increasing time. In fact, cytochrome C is known to have similar characteristics such as molecular weight and surface charge with respect to growth factors, mainly basic fibroblast growth factor (bFGF) [15], therefore this study may be useful in designing bone scaffolds with delivering of bFGF to elicit therapeutic effects such as the increase in cell proliferative potential and blood vessel formation for bone tissue engineering, which remains as further study.
8 6 Bare scaffold
4 2 0 0
2
4
6
8
10
12
14
16
18
Loading time (h) Fig. 3. Protein (cytochrome C) loading capacity of the mineralized PCL scaffold in comparison with that of bare scaffold. The maximal loading amount was about 2.5 times higher in the mineralized scaffold.
4. Conclusions The surface of the PCL biopolymer scaffold constructed using a robocasting method was mineralized with carbonate apatite nanocrystallites by a series of solution-based treatments. Tissue cells actively adhered and proliferated on the surface-mineralized scaffold, which was in direct contrast with the poor cell affinity of the untreated PCL scaffold. Moreover, the surface mineralization facilitated the scaffold to hold significantly enhanced protein loading capacity. The study
3562
B. Dorj et al. / Materials Letters 65 (2011) 3559–3562
demonstrates the surface-mineralized polymer scaffolds are potentially useful in bone tissue engineering. Acknowledgment This work was supported by Priority Research Centers Program (grant#: 2009–0093829) and WCU (World Class University) program (grant#: R31-10069) through the National Research Foundation (NRF) funded by the Ministry of Education, Science and Technology. Authors thank the support of IBST in Dankook University.
[4] [5] [6] [7] [8] [9] [10] [11] [12] [13]
References [1] Lu L, Mikos A. MRS Bull 1996;21:28. [2] Hollister SJ, Maddox RD, Taboas JM. Biomaterials 2002;23:4095. [3] Chua CK, Leong KF, Lim CS. Rapid prototyping2nd ed. . Singapore: World scientific; 2003.
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