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Hierarchically biomimetic scaffold of a collagen–mesoporous bioactive glass nanofiber composite for bone tissue engineering
This content has been downloaded from IOPscience. Please scroll down to see the full text. 2015 Biomed. Mater. 10 025007 (http://iopscience.iop.org/1748-605X/10/2/025007) View the table of contents for this issue, or go to the journal homepage for more
Download details: IP Address: 169.230.243.252 This content was downloaded on 31/03/2015 at 12:02
Please note that terms and conditions apply.
Biomed. Mater. 10 (2015) 025007
doi:10.1088/1748-6041/10/2/025007
Paper
received
17 October 2014
Hierarchically biomimetic scaffold of a collagen–mesoporous bioactive glass nanofiber composite for bone tissue engineering
re vised
4 March 2015 accep ted for publication
Fu-Yin Hsu, Meng-Ru Lu, Ru-Chun Weng and Hsiu-Mei Lin
6 March 2015
Department of Bioscience and Biotechnology, National Taiwan Ocean University, Keelung, Taiwan
published
E-mail: [email protected]
25 March 2015
Keywords: mesoporous bioactive glass nanofiber, collagen, macroporous structure, bone regeneration
Abstract Mesoporous bioactive glass nanofibers (MBGNFs) were prepared by a sol–gel/electrospinning technique. Subsequently, a collagen–MBGNF (CM) composite scaffold that simultaneously possessed a macroporous structure and collagen nanofibers was fabricated by a gelation and freeze-drying process. Additionally, immersing the CM scaffold in a simulated body fluid resulted in the formation of bone-like apatite minerals on the surface. The CM scaffold provided a suitable environment for attachment to the cytoskeleton. Based on the measured alkaline phosphatase activity and protein expression levels of osteocalcin and bone sialoprotein, the CM scaffold promoted the differentiation and mineralization of MG63 osteoblast-like cells. In addition, the bone regeneration ability of the CM scaffold was examined using a rat calvarial defect model in vivo. The results revealed that CM is biodegradable and could promote bone regeneration. Therefore, a CM composite scaffold is a potential bone graft for bone tissue engineering applications. S Online supplementary data available from stacks.iop.org/BMM/10/025007
1. Introduction An ideal synthetic bone graft must be suitably biocompatible, osteoconductive (sustaining cell phenotype), osteoinductive (allowing cell differentiation), and bioresorbable [1]. The bone matrix is approximately 35 wt% organic (mainly type I collagen) and 65 wt% inorganic material (mainly hydroxyapatite). Collagen sponges have been widely used as scaffolds due to their similarity in composition to the extracellular matrix, their high biocompatibility, and their low immunogenicity [2]. Collagen sponges are fabricated from acid-soluble collagen solutions by a direct freeze-drying process. The structure of a collagen sponge is highly porous with interconnected pores. Nevertheless, previous studies have reported that the structure of collagen molecules was globular because of the aggregation of a few collagen molecules at low pH values, whereas the collagen molecules assemble into fibrillar structures near neutral pH values [3–5]. This result correlated with the lower tendency of collagen fibrils to reconstitute at low pH values. Moreover, Tsai et al [6] demonstrated that the unique pattern of collagen nanofibers enhanced the differentiation and mineralization of osteoblastic cells. Thus, one of the aims of this study is to fabricate a collagen scaffold © 2015 IOP Publishing Ltd
that is highly porous with interconnected pores and to maintain the existence of collagen nanofibers. We had previously fabricated a composite containing collagen and bioactive inorganic material for bone tissue regeneration and demonstrated that a composite with a bioactive inorganic material exhibited better in vitro osteogenic properties than collagen alone [7]. Bioactive glasses (BGs) have been widely used as bone grafts due to their excellent bioactivity. Furthermore, some researchers demonstrated that the release of Ca, P and Si ions from BG promoted the differentiation of osteoblasts and enhanced new bone formation [8, 9]. Kim demonstrated that a BG nanofiber–collagen nanocomposite could exhibit excellent bioactivity in vitro and promote the alkaline phosphatase (ALP) activity of osteoblastic cells [10]. Melting-derived BGs are dense and have no texture properties. Yan et al [11] synthesized mesoporous bioactive glasses (MBGs) that possessed an ordered mesoporous channel structure, large pore volume and surface area. Li et al [12] found that MBG had higher drug loading efficiency and drug release capacity than non-mesoporous biological glass. Vallet-Regı et al [9] found that increasing the specific surface area and pore volume of BG greatly accelerated the formation of carbonated hydroxyapatite and therefore enhanced the bioactive behavior. Therefore, MBG
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
possesses excellent bone-forming activity. However, it is very difficult to use MBG as a scaffold for the bone regeneration because the pore size of MBG is too small to promote cell attachment, migration, and tissue ingrowth. Hence, another aim of this study was to fabricate a collagen–mesoporous bioactive glass nanofiber (MBGNF) composite scaffold and evaluate the bioactive properties of the composite in vitro and in vivo.
2. Materials and methods 2.1. Reagents Pluronic P123 (MW = 5800), tetraethyl orthosilicate (TEOS), triethyl phosphate (TEP), calcium nitrate tetrahydrate, and poly(vinyl pyrrolidone) (PVP) were purchased from Sigma-Aldrich Chemical Company (St. Louis, MO). The other chemicals used were of reagent grade, unless otherwise stated. Type I collagen was prepared from calf skin, as previously described [13], and acid-soluble collagen was used for all experiments. 2.2. Synthesis of MBGNFs The MBGNFs were fabricated by utilizing an electrospinning technique based on the use of a sol–gel precursor. The MBG precursor solution was prepared by mixing Pluronic P123 (1.0 g), TEOS (1.59 ml), calcium nitrate tetrahydrate (0.315 g), TEP (99.8%, 0.114 ml), hydrochloric acid (1 M, 0.804 ml) and ethanol (10 ml). Subsequently, Pluronic P123 (0.2 g), PVP (0.5 g) and ethanol (2.5 ml) were prepared and incorporated into MBG precursor solution (5 ml) to obtain a transparent MBGNF precursor solution. The MBGNF precursor solution was drawn into a plastic syringe with a needle (18 gauge). The plastic syringe was then placed into a syringe pump (RAZEL, Model R99-E), which provided a steady flow rate of the solution at 12.7 μl min−1. A high electric field (1.0 kV cm−1) was applied between the aluminum grounded collector and the needle tip. The MBGNF precursor solution formed non-woven nanofibers on the grounded collector. The collected nanofibers were calcined at 600 °C for 5 h to completely eliminate the organic sources and obtain the MBGNF matrices. The diameter of MBGNF was measured from scanning electron microscopy (SEM) images, and the average pore diameter of MBGNF was measured from transmission electron microscopy (TEM) images using image analysis software (Image J, NIH, USA). 2.3. Fabrication of collagen–MBGNF composite scaffold The reconstitution of the collagen to a native fibrillar structure was initiated by adjusting the pH of the acid-soluble collagen solution (6.2 mg ml−1, 1 ml) to 7.0 using a disodium hydrogen phosphate solution (1.5 M, 80 μl). The final concentration of the neutral collagen solution was diluted by phosphate buffer solution (PBS (pH 7.0)). The MBGNF matrices were reduced to fragments by sonication. The MBGNFs were 2
mixed and suspended in a neutral collagen solution with a weight ratio of 65 : 35 (collagen to MBGNFs) at 4 °C. The scaffolds were prepared following a classical gelation and freeze-drying process. Briefly, 0.5 ml of the collagen or collagen/MBGNF solution (pH 7.0) was added to the wells of a 24-well plate at 37 °C for 12 h to induce collagen fibrillogenesis and then form the gel. Subsequently, the gel was frozen at −80 °C for at least 4 h and then placed into a freeze-dryer and dried overnight. Freeze-dried scaffolds were cross-linked with saturated glutaraldehyde vapor at 37 °C for 6 h. After exposure to glutaraldehyde vapor, the scaffolds were removed and washed with deionized water for 5 min three times to remove any unbound glutaraldehyde. After washing, the scaffolds were lyophilized and sterilized by UV irradiation before cell culture and implantation. 2.4. Morphology observation of scaffold The morphology of the collagen (COL) and collagen– MBGNF (CM) scaffolds was examined using SEM (HITACHI S-3000N). Briefly, the scaffold was sputtercoated with gold and visualized by SEM. The average pore diameter of the scaffold was measured from the SEM images using image analysis software (Image J, NIH, USA). 2.4.1. Cross-linking degree of scaffold The cross-linking treatment of COL is one of the most important issues for the porous COL scaffold because the cross-linking degree of the porous COL scaffold influences the mechanical properties and biodegradation rate of the scaffold [14]. The crosslinking degree of the COL and CM scaffolds was determined by ninhydrin assay [15]. Briefly, the sample was heated with ninhydrin solution (1 ml) for 35 min at 95 °C, and 50% isopropanol (200 μl) was added to all samples. The quantity of free amino groups was determined based on the absorbance value of the solution at 570 nm measured with a spectrophotometer. The quantity of free amino groups in a scaffold is proportional to the optical absorbance of the solution. The degree of cross-linking of the scaffold was calculated according to the following equation: Degree of cross-linking (%) = [(OD b − ODa )/OD b] × 100%,
where ODb is the optical absorbance of the solution before cross-linking treatment as a blank, and ODa is the optical absorbance of the solution after crosslinking treatment. 2.4.2. Shrinkage ratios of scaffold after cross-linking treatment To quantify the degree of shrinkage of the COL and CM scaffolds after cross-linking treatment, the diameter of scaffolds both before and after cross-linking treatment was measured. The shrinkage of the scaffold was quantified using a dimensionless shrinkage ratio, calculated by the following equation.
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
Shrinkage ratio (%) = [(D b − Da)/D b] × 100%,
where Db and Da are the diameters of the scaffold before and after cross-linking treatment, respectively. 2.4.3. Swelling ratio of scaffold Cross-linked COL and CM scaffolds were weighed and immersed in PBS at room temperature for 3 h. Finally, the wet scaffold was weighed and calculated according to the following equation: Swelling ratio (%) = [(Ww − Wd)/Wd] × 100%,
where Wd is the weight of the dry scaffold, and Ww is the weight of the wet scaffold. 2.4.4. Compressive modulus For mechanical testing, the loads for the compressive testing of the COL (diameter: 6.2 mm, length: 21 mm) and CM (diameter: 6.9 mm, length: 24 mm) scaffolds were obtained for a crosshead speed of 1 mm min−1 using a materials testing machine (ElectroForce® 3100, BOSE). The compressive modulus was determined from the compressive curve. 2.5. Bioactivity of scaffold Apatite mineralizes on scaffolds in simulated body fluid (SBF), which has been considered an important phenomenon and regarded as evidence for the bioactivity of materials [16]. COL and CM scaffolds were immersed in 1.5 × SBF for up to 7 d. The 1.5 × SBF was prepared by dissolving reagent-grade NaCl (213 mM), KCl (4.5 mM), CaCl2 (3.8 mM), MgCl2· 6H2O (2.3 mM), K2HPO4 · 3H2O (1.5 mM), NaHCO3 (6.3 mM), and Na2SO4 (0.75 mM) in deionized water. The solution was buffered at pH 7.4 with tris(hydroxymethyl) aminomethane and HCl (1 M) at 37 °C. After soaking, samples were removed from the SBF, washed with deionized water, and dried at 25 °C. The surface structure and the morphology of the scaffold before and after soaking in the SBF solutions were characterized using SEM. The phase of the crystals deposited on the scaffold was determined using x-ray diffraction (XRD). The angle of diffraction (2θ) was from 15° to 55°. The standard diffraction pattern of hydroxyapatite #9-432 was used as the reference standard. 2.6. Cellular proliferation on scaffold The MG63 osteoblast-like cell line (BCRC no. 60279) was purchased from the Bioresource Collection and Research Center, Taiwan. The CM scaffolds were placed into 24-well tissue culture dishes containing a suspension of MG63 cells (1 × 105 cells per well) in minimum essential media (MEM) supplemented with ascorbic acid (50 μg ml−1), β-glycerophosphate (10 mM), penicillin (100 U ml −1), streptomycin (100 μg ml −1) and fetal bovine serum (10%) and were incubated at 37 °C in a humidified atmosphere with 5% CO2. The cultures of cell-seeded scaffolds were harvested on days 1, 3, 7 d for cell proliferation assessments. 3
Cel l v iabilit ies were de ter mined using 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assays. The cell-containing scaffolds were incubated with MTT solution (0.5 mg ml−1) at 37 °C for 3 h. After removal of the supernatants, dimethyl sulfoxide (DMSO) was added, and the absorbance was measured at a wavelength of 570 nm by a microplate reader (Biotek uQuant). 2.7. Cell morphology and cytoskeleton organization The scaffolds were collected at various time intervals after incubation. The scaffolds were washed with PBS and fixed with glutaraldehyde solution (2.5%) for 1 h. After washing in PBS, the samples were post-fixed in osmium tetroxide (1%) for 1 h. The fixed scaffolds were dehydrated in a gradient of ethanol solution and dried by critical point drying. Finally, the dehydrated scaffolds were sputter-coated with gold and observed by SEM. 2.8. Immunofluorescence At the predetermined time points, the scaffold was fixed with paraformaldehyde (3.7% in PBS) for 10 min. The scaffold was subsequently rinsed in PBS with 0.1% Triton X-100 for 5 min. To reduce nonspecific binding, the scaffold was blocked with bovine serum albumin (BSA) (1% in PBS) for 1 h. After blocking, the scaffold was incubated with osteoblast-specific marker protein primary antibody osteocalcin (OCN, Millipore AB10911) or bone sialoprotein (BSP, Millipore AB1854) overnight at 4 °C. Following incubation with the primary antibody, the scaffold was washed and incubated with a secondary antibody (goat anti-rabbit antibody conjugated rhodamine, Santa Cruz, sc2091, USA) for 30 min. Finally, the scaffold was incubated with fluorescein isothiocyanate (FITC)-conjugated phalloidin for 20 min to stain the cytoskeleton organization and incubated with 4’,6-diamidino-2-phenylindole (DAPI) solution for 3 min to stain the DNA in the cells. After washing with PBS, the scaffold was observed under a laser scanning confocal microscope (LSCM, Zeiss LSM 510 META). Osteoblast-specific marker protein, cytoskeleton and nuclei were stained in red, green and blue, respectively. 2.9. ALP activity of MG63 on scaffold The cell-seeded scaffold was washed with PBS and suspended in PBS (0.5 ml) containing glycine (0.1 M), MgCl 2 (1 mM) and Triton X-100 (1%). Following lysis, 50 μl of the lysate was incubated with 150 μl of ALP reagent (Randox ALP detection kit) at 37 °C for 10 min. The absorbance of p-nitrophenol formed from p-nitrophenyl phosphate was determined by monitoring the light absorbance at 405 nm, and the amount of p-nitrophenol should be equivalent to the ALP activity. The results of the ALP activity assay were normalized by the amount of DNA on the scaffold.
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
Figure 1. (a) SEM and (b) TEM images of MBGNF.
The amount of DNA was measured using the Quant-iT™ PicoGreen® dsDNA Assay Kit (Invitrogen, Carlsbad, CA, USA). The cell-seeded scaffold was washed with PBS and lysed with radioimmunoprecipitation assay (RIPA) buffer. Then, 10 μl of cell lysate solution was added into 100 μl of PicoGreen reagent (diluted 1 : 200 in Quant-iT™ dsDNA HS buffer), and the intensity of fluorescence was measured at an emission wavelength of 502 nm and an excitation wavelength of 540 nm using a spectrofluorometer (SpectraMax M2e, Molecular devices, Sunnyvale, CA, USA). 2.10. Animal study The animal experiment was approved by the Institutional Animal Care and Use Committee of National Taiwan Ocean University (approval ID: 99010). The scaffolds were sterilized by ultraviolet irradiation before implantation. Adult male Sprague–Dawley (SD) rats (approximately 300 g) were used in this experiment. The animals were anesthetized with an intramuscular injection of Zoletil 50 (20 mg kg −1) and Rompun (0.5 mg kg−1). The anesthetized rats were fixed, and the calvarial skin and periosteum were incised after shaving. A circular defect of the parietal bone approximately 6 mm in diameter was created using a 0.5 mm drill. To trace the newly formed bone, the animals were intramuscularly injected with tetracycline HCl (TC, 20 mg kg−1) at 1 and 5 weeks and with xylenol orange (XO, 90 mg kg−1) at 3 and 7 weeks after implantation. The animals were sacrificed at 4 and 8 weeks after implantation. The rat cranium was dissected and examined by x-ray radiography. The remaining bone defect was measured by x-ray radiography using image analysis software (Image J, NIH, USA). The percentage of newly formed bone in the defect (n-Bone%) was calculated by the following equation: n-Bone% = [(Aod − Ard )/Aod ] × 100%,
where Aod and Ard are the area of the original bone defect and remaining bone defect before and after implant, respectively. To observe the new bone formation, the dissected tissue was fixed in 10% neutral buffered formalin. The fixed specimens were dehydrated using gradations of alcohol, acetone, and methylmethacrylate and then embedded in polymethylmethacrylate. After hardening, the speci4
men was cut into approximately 1 mm thick slices using a diamond-saw microtome. The tissue sections were glued to glass slides and polished manually to approximately 50 μm thickness under cooling water. The specimens were observed using a fluorescence microscope [17]. 2.11. Statistical analyses Results are expressed as the mean ± standard deviation. Statistical analyses were performed using the SPSS v.10 software. The non-parametric Mann–Whitney U test was used to compare MBGNF effects versus COL alone. The differences were considered statistically significant at p
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My IOPscience
Hierarchically biomimetic scaffold of a collagen–mesoporous bioactive glass nanofiber composite for bone tissue engineering
This content has been downloaded from IOPscience. Please scroll down to see the full text. 2015 Biomed. Mater. 10 025007 (http://iopscience.iop.org/1748-605X/10/2/025007) View the table of contents for this issue, or go to the journal homepage for more
Download details: IP Address: 169.230.243.252 This content was downloaded on 31/03/2015 at 12:02
Please note that terms and conditions apply.
Biomed. Mater. 10 (2015) 025007
doi:10.1088/1748-6041/10/2/025007
Paper
received
17 October 2014
Hierarchically biomimetic scaffold of a collagen–mesoporous bioactive glass nanofiber composite for bone tissue engineering
re vised
4 March 2015 accep ted for publication
Fu-Yin Hsu, Meng-Ru Lu, Ru-Chun Weng and Hsiu-Mei Lin
6 March 2015
Department of Bioscience and Biotechnology, National Taiwan Ocean University, Keelung, Taiwan
published
E-mail: [email protected]
25 March 2015
Keywords: mesoporous bioactive glass nanofiber, collagen, macroporous structure, bone regeneration
Abstract Mesoporous bioactive glass nanofibers (MBGNFs) were prepared by a sol–gel/electrospinning technique. Subsequently, a collagen–MBGNF (CM) composite scaffold that simultaneously possessed a macroporous structure and collagen nanofibers was fabricated by a gelation and freeze-drying process. Additionally, immersing the CM scaffold in a simulated body fluid resulted in the formation of bone-like apatite minerals on the surface. The CM scaffold provided a suitable environment for attachment to the cytoskeleton. Based on the measured alkaline phosphatase activity and protein expression levels of osteocalcin and bone sialoprotein, the CM scaffold promoted the differentiation and mineralization of MG63 osteoblast-like cells. In addition, the bone regeneration ability of the CM scaffold was examined using a rat calvarial defect model in vivo. The results revealed that CM is biodegradable and could promote bone regeneration. Therefore, a CM composite scaffold is a potential bone graft for bone tissue engineering applications. S Online supplementary data available from stacks.iop.org/BMM/10/025007
1. Introduction An ideal synthetic bone graft must be suitably biocompatible, osteoconductive (sustaining cell phenotype), osteoinductive (allowing cell differentiation), and bioresorbable [1]. The bone matrix is approximately 35 wt% organic (mainly type I collagen) and 65 wt% inorganic material (mainly hydroxyapatite). Collagen sponges have been widely used as scaffolds due to their similarity in composition to the extracellular matrix, their high biocompatibility, and their low immunogenicity [2]. Collagen sponges are fabricated from acid-soluble collagen solutions by a direct freeze-drying process. The structure of a collagen sponge is highly porous with interconnected pores. Nevertheless, previous studies have reported that the structure of collagen molecules was globular because of the aggregation of a few collagen molecules at low pH values, whereas the collagen molecules assemble into fibrillar structures near neutral pH values [3–5]. This result correlated with the lower tendency of collagen fibrils to reconstitute at low pH values. Moreover, Tsai et al [6] demonstrated that the unique pattern of collagen nanofibers enhanced the differentiation and mineralization of osteoblastic cells. Thus, one of the aims of this study is to fabricate a collagen scaffold © 2015 IOP Publishing Ltd
that is highly porous with interconnected pores and to maintain the existence of collagen nanofibers. We had previously fabricated a composite containing collagen and bioactive inorganic material for bone tissue regeneration and demonstrated that a composite with a bioactive inorganic material exhibited better in vitro osteogenic properties than collagen alone [7]. Bioactive glasses (BGs) have been widely used as bone grafts due to their excellent bioactivity. Furthermore, some researchers demonstrated that the release of Ca, P and Si ions from BG promoted the differentiation of osteoblasts and enhanced new bone formation [8, 9]. Kim demonstrated that a BG nanofiber–collagen nanocomposite could exhibit excellent bioactivity in vitro and promote the alkaline phosphatase (ALP) activity of osteoblastic cells [10]. Melting-derived BGs are dense and have no texture properties. Yan et al [11] synthesized mesoporous bioactive glasses (MBGs) that possessed an ordered mesoporous channel structure, large pore volume and surface area. Li et al [12] found that MBG had higher drug loading efficiency and drug release capacity than non-mesoporous biological glass. Vallet-Regı et al [9] found that increasing the specific surface area and pore volume of BG greatly accelerated the formation of carbonated hydroxyapatite and therefore enhanced the bioactive behavior. Therefore, MBG
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
possesses excellent bone-forming activity. However, it is very difficult to use MBG as a scaffold for the bone regeneration because the pore size of MBG is too small to promote cell attachment, migration, and tissue ingrowth. Hence, another aim of this study was to fabricate a collagen–mesoporous bioactive glass nanofiber (MBGNF) composite scaffold and evaluate the bioactive properties of the composite in vitro and in vivo.
2. Materials and methods 2.1. Reagents Pluronic P123 (MW = 5800), tetraethyl orthosilicate (TEOS), triethyl phosphate (TEP), calcium nitrate tetrahydrate, and poly(vinyl pyrrolidone) (PVP) were purchased from Sigma-Aldrich Chemical Company (St. Louis, MO). The other chemicals used were of reagent grade, unless otherwise stated. Type I collagen was prepared from calf skin, as previously described [13], and acid-soluble collagen was used for all experiments. 2.2. Synthesis of MBGNFs The MBGNFs were fabricated by utilizing an electrospinning technique based on the use of a sol–gel precursor. The MBG precursor solution was prepared by mixing Pluronic P123 (1.0 g), TEOS (1.59 ml), calcium nitrate tetrahydrate (0.315 g), TEP (99.8%, 0.114 ml), hydrochloric acid (1 M, 0.804 ml) and ethanol (10 ml). Subsequently, Pluronic P123 (0.2 g), PVP (0.5 g) and ethanol (2.5 ml) were prepared and incorporated into MBG precursor solution (5 ml) to obtain a transparent MBGNF precursor solution. The MBGNF precursor solution was drawn into a plastic syringe with a needle (18 gauge). The plastic syringe was then placed into a syringe pump (RAZEL, Model R99-E), which provided a steady flow rate of the solution at 12.7 μl min−1. A high electric field (1.0 kV cm−1) was applied between the aluminum grounded collector and the needle tip. The MBGNF precursor solution formed non-woven nanofibers on the grounded collector. The collected nanofibers were calcined at 600 °C for 5 h to completely eliminate the organic sources and obtain the MBGNF matrices. The diameter of MBGNF was measured from scanning electron microscopy (SEM) images, and the average pore diameter of MBGNF was measured from transmission electron microscopy (TEM) images using image analysis software (Image J, NIH, USA). 2.3. Fabrication of collagen–MBGNF composite scaffold The reconstitution of the collagen to a native fibrillar structure was initiated by adjusting the pH of the acid-soluble collagen solution (6.2 mg ml−1, 1 ml) to 7.0 using a disodium hydrogen phosphate solution (1.5 M, 80 μl). The final concentration of the neutral collagen solution was diluted by phosphate buffer solution (PBS (pH 7.0)). The MBGNF matrices were reduced to fragments by sonication. The MBGNFs were 2
mixed and suspended in a neutral collagen solution with a weight ratio of 65 : 35 (collagen to MBGNFs) at 4 °C. The scaffolds were prepared following a classical gelation and freeze-drying process. Briefly, 0.5 ml of the collagen or collagen/MBGNF solution (pH 7.0) was added to the wells of a 24-well plate at 37 °C for 12 h to induce collagen fibrillogenesis and then form the gel. Subsequently, the gel was frozen at −80 °C for at least 4 h and then placed into a freeze-dryer and dried overnight. Freeze-dried scaffolds were cross-linked with saturated glutaraldehyde vapor at 37 °C for 6 h. After exposure to glutaraldehyde vapor, the scaffolds were removed and washed with deionized water for 5 min three times to remove any unbound glutaraldehyde. After washing, the scaffolds were lyophilized and sterilized by UV irradiation before cell culture and implantation. 2.4. Morphology observation of scaffold The morphology of the collagen (COL) and collagen– MBGNF (CM) scaffolds was examined using SEM (HITACHI S-3000N). Briefly, the scaffold was sputtercoated with gold and visualized by SEM. The average pore diameter of the scaffold was measured from the SEM images using image analysis software (Image J, NIH, USA). 2.4.1. Cross-linking degree of scaffold The cross-linking treatment of COL is one of the most important issues for the porous COL scaffold because the cross-linking degree of the porous COL scaffold influences the mechanical properties and biodegradation rate of the scaffold [14]. The crosslinking degree of the COL and CM scaffolds was determined by ninhydrin assay [15]. Briefly, the sample was heated with ninhydrin solution (1 ml) for 35 min at 95 °C, and 50% isopropanol (200 μl) was added to all samples. The quantity of free amino groups was determined based on the absorbance value of the solution at 570 nm measured with a spectrophotometer. The quantity of free amino groups in a scaffold is proportional to the optical absorbance of the solution. The degree of cross-linking of the scaffold was calculated according to the following equation: Degree of cross-linking (%) = [(OD b − ODa )/OD b] × 100%,
where ODb is the optical absorbance of the solution before cross-linking treatment as a blank, and ODa is the optical absorbance of the solution after crosslinking treatment. 2.4.2. Shrinkage ratios of scaffold after cross-linking treatment To quantify the degree of shrinkage of the COL and CM scaffolds after cross-linking treatment, the diameter of scaffolds both before and after cross-linking treatment was measured. The shrinkage of the scaffold was quantified using a dimensionless shrinkage ratio, calculated by the following equation.
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
Shrinkage ratio (%) = [(D b − Da)/D b] × 100%,
where Db and Da are the diameters of the scaffold before and after cross-linking treatment, respectively. 2.4.3. Swelling ratio of scaffold Cross-linked COL and CM scaffolds were weighed and immersed in PBS at room temperature for 3 h. Finally, the wet scaffold was weighed and calculated according to the following equation: Swelling ratio (%) = [(Ww − Wd)/Wd] × 100%,
where Wd is the weight of the dry scaffold, and Ww is the weight of the wet scaffold. 2.4.4. Compressive modulus For mechanical testing, the loads for the compressive testing of the COL (diameter: 6.2 mm, length: 21 mm) and CM (diameter: 6.9 mm, length: 24 mm) scaffolds were obtained for a crosshead speed of 1 mm min−1 using a materials testing machine (ElectroForce® 3100, BOSE). The compressive modulus was determined from the compressive curve. 2.5. Bioactivity of scaffold Apatite mineralizes on scaffolds in simulated body fluid (SBF), which has been considered an important phenomenon and regarded as evidence for the bioactivity of materials [16]. COL and CM scaffolds were immersed in 1.5 × SBF for up to 7 d. The 1.5 × SBF was prepared by dissolving reagent-grade NaCl (213 mM), KCl (4.5 mM), CaCl2 (3.8 mM), MgCl2· 6H2O (2.3 mM), K2HPO4 · 3H2O (1.5 mM), NaHCO3 (6.3 mM), and Na2SO4 (0.75 mM) in deionized water. The solution was buffered at pH 7.4 with tris(hydroxymethyl) aminomethane and HCl (1 M) at 37 °C. After soaking, samples were removed from the SBF, washed with deionized water, and dried at 25 °C. The surface structure and the morphology of the scaffold before and after soaking in the SBF solutions were characterized using SEM. The phase of the crystals deposited on the scaffold was determined using x-ray diffraction (XRD). The angle of diffraction (2θ) was from 15° to 55°. The standard diffraction pattern of hydroxyapatite #9-432 was used as the reference standard. 2.6. Cellular proliferation on scaffold The MG63 osteoblast-like cell line (BCRC no. 60279) was purchased from the Bioresource Collection and Research Center, Taiwan. The CM scaffolds were placed into 24-well tissue culture dishes containing a suspension of MG63 cells (1 × 105 cells per well) in minimum essential media (MEM) supplemented with ascorbic acid (50 μg ml−1), β-glycerophosphate (10 mM), penicillin (100 U ml −1), streptomycin (100 μg ml −1) and fetal bovine serum (10%) and were incubated at 37 °C in a humidified atmosphere with 5% CO2. The cultures of cell-seeded scaffolds were harvested on days 1, 3, 7 d for cell proliferation assessments. 3
Cel l v iabilit ies were de ter mined using 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assays. The cell-containing scaffolds were incubated with MTT solution (0.5 mg ml−1) at 37 °C for 3 h. After removal of the supernatants, dimethyl sulfoxide (DMSO) was added, and the absorbance was measured at a wavelength of 570 nm by a microplate reader (Biotek uQuant). 2.7. Cell morphology and cytoskeleton organization The scaffolds were collected at various time intervals after incubation. The scaffolds were washed with PBS and fixed with glutaraldehyde solution (2.5%) for 1 h. After washing in PBS, the samples were post-fixed in osmium tetroxide (1%) for 1 h. The fixed scaffolds were dehydrated in a gradient of ethanol solution and dried by critical point drying. Finally, the dehydrated scaffolds were sputter-coated with gold and observed by SEM. 2.8. Immunofluorescence At the predetermined time points, the scaffold was fixed with paraformaldehyde (3.7% in PBS) for 10 min. The scaffold was subsequently rinsed in PBS with 0.1% Triton X-100 for 5 min. To reduce nonspecific binding, the scaffold was blocked with bovine serum albumin (BSA) (1% in PBS) for 1 h. After blocking, the scaffold was incubated with osteoblast-specific marker protein primary antibody osteocalcin (OCN, Millipore AB10911) or bone sialoprotein (BSP, Millipore AB1854) overnight at 4 °C. Following incubation with the primary antibody, the scaffold was washed and incubated with a secondary antibody (goat anti-rabbit antibody conjugated rhodamine, Santa Cruz, sc2091, USA) for 30 min. Finally, the scaffold was incubated with fluorescein isothiocyanate (FITC)-conjugated phalloidin for 20 min to stain the cytoskeleton organization and incubated with 4’,6-diamidino-2-phenylindole (DAPI) solution for 3 min to stain the DNA in the cells. After washing with PBS, the scaffold was observed under a laser scanning confocal microscope (LSCM, Zeiss LSM 510 META). Osteoblast-specific marker protein, cytoskeleton and nuclei were stained in red, green and blue, respectively. 2.9. ALP activity of MG63 on scaffold The cell-seeded scaffold was washed with PBS and suspended in PBS (0.5 ml) containing glycine (0.1 M), MgCl 2 (1 mM) and Triton X-100 (1%). Following lysis, 50 μl of the lysate was incubated with 150 μl of ALP reagent (Randox ALP detection kit) at 37 °C for 10 min. The absorbance of p-nitrophenol formed from p-nitrophenyl phosphate was determined by monitoring the light absorbance at 405 nm, and the amount of p-nitrophenol should be equivalent to the ALP activity. The results of the ALP activity assay were normalized by the amount of DNA on the scaffold.
Biomed. Mater. 10 (2015) 025007
F-Y Hsu et al
Figure 1. (a) SEM and (b) TEM images of MBGNF.
The amount of DNA was measured using the Quant-iT™ PicoGreen® dsDNA Assay Kit (Invitrogen, Carlsbad, CA, USA). The cell-seeded scaffold was washed with PBS and lysed with radioimmunoprecipitation assay (RIPA) buffer. Then, 10 μl of cell lysate solution was added into 100 μl of PicoGreen reagent (diluted 1 : 200 in Quant-iT™ dsDNA HS buffer), and the intensity of fluorescence was measured at an emission wavelength of 502 nm and an excitation wavelength of 540 nm using a spectrofluorometer (SpectraMax M2e, Molecular devices, Sunnyvale, CA, USA). 2.10. Animal study The animal experiment was approved by the Institutional Animal Care and Use Committee of National Taiwan Ocean University (approval ID: 99010). The scaffolds were sterilized by ultraviolet irradiation before implantation. Adult male Sprague–Dawley (SD) rats (approximately 300 g) were used in this experiment. The animals were anesthetized with an intramuscular injection of Zoletil 50 (20 mg kg −1) and Rompun (0.5 mg kg−1). The anesthetized rats were fixed, and the calvarial skin and periosteum were incised after shaving. A circular defect of the parietal bone approximately 6 mm in diameter was created using a 0.5 mm drill. To trace the newly formed bone, the animals were intramuscularly injected with tetracycline HCl (TC, 20 mg kg−1) at 1 and 5 weeks and with xylenol orange (XO, 90 mg kg−1) at 3 and 7 weeks after implantation. The animals were sacrificed at 4 and 8 weeks after implantation. The rat cranium was dissected and examined by x-ray radiography. The remaining bone defect was measured by x-ray radiography using image analysis software (Image J, NIH, USA). The percentage of newly formed bone in the defect (n-Bone%) was calculated by the following equation: n-Bone% = [(Aod − Ard )/Aod ] × 100%,
where Aod and Ard are the area of the original bone defect and remaining bone defect before and after implant, respectively. To observe the new bone formation, the dissected tissue was fixed in 10% neutral buffered formalin. The fixed specimens were dehydrated using gradations of alcohol, acetone, and methylmethacrylate and then embedded in polymethylmethacrylate. After hardening, the speci4
men was cut into approximately 1 mm thick slices using a diamond-saw microtome. The tissue sections were glued to glass slides and polished manually to approximately 50 μm thickness under cooling water. The specimens were observed using a fluorescence microscope [17]. 2.11. Statistical analyses Results are expressed as the mean ± standard deviation. Statistical analyses were performed using the SPSS v.10 software. The non-parametric Mann–Whitney U test was used to compare MBGNF effects versus COL alone. The differences were considered statistically significant at p
