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Effects of crystalline phase on the biological properties of collagen–hydroxyapatite composites
Acta Biomaterialia 6 (2010) 2189–2199
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Effects of crystalline phase on the biological properties of collagen–hydroxyapatite composites L. Zhang, P. Tang, M. Xu, W. Zhang, W. Chai, Y. Wang * Chinese PLA General Hospital, Department of Orthopaedics, Beijing 100853, China
a r t i c l e
i n f o
Article history: Received 18 October 2009 Received in revised form 17 December 2009 Accepted 21 December 2009 Available online 28 December 2009 Keywords: Hydroxyapatite Collagen Crystalline phase Biomineralization Pre-crystallization
a b s t r a c t The objective of this study was to investigate the effects of spatial structure and crystalline phase on the biological performance of collagen–hydroxyapatite (Col–HA) composite prepared by biomineralization crystallization. Two types of Col–HA composites were prepared using mineralization crystallization (MC composites) and pre-crystallization (PC composites), respectively. Structural characteristics were analyzed by scanning electron microscopy and transmission electron microscopy. Surface elemental compositions were measured by electron spectroscopy for chemical analysis (ESCA). These composites were used in in vivo repair of bone defects. The effects of the crystalline phase on the biological performance of Col–HA composites were investigated using radionuclide bone scan, histopathology and morphological observation. It was observed that in MC composites, HA was located on the surface of the collagen fibers and aggregated into crystal balls, whereas HA in PC composites was scattered among the collagen fibers. ESCA showed that phosphorus and calcium were 8.99% and 17.56% on MC composite surface, compared with 4.39% and 5.86% on the PC composite surface. In vivo bone defect repair experiments revealed that radionuclide uptake was significantly higher in the area implanted with the PC composite than in the contralateral area implanted with the MC composite. Throughout the whole repair process, the PC composite proved to be superior to the MC composite with regard to capillary-forming capacity and the amount of newly formed bone tissue. So it could be concluded that HA placement on collagen fibers affected the biological performance of Col–HA composites. Pre-crystallization made HA scattered among collagen fibers, creating a better structure for bone defect repair in comparison with MC Col–HA composites. Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Great efforts have been made to develop biomaterials capable of repairing bone defects resulting from various causes. Bone tissue is a particularly complex composite because it contains multiple levels of organic and inorganic nanophases. At the lowest level of this hierarchy is a three-dimensional composite made up of collagen triple helices, and the hydroxyapatite crystals grow in some different form: the majority of the mineral lies in such a way that its c-axes are oriented along the long axes of the fibrils (called intra-fibrillar mineral) [1–3], some lies between the fibrils, the c-axes of which are perpendicular to the collagen molecular axis (called inter-fibrillar mineral or extra-fibrillar mineral) [4,5]. Many researchers have attempted to construct composites of hydroxyapatite and collagen (Col–HA) for bone defect repair. Self-assembly and biomineralization have been used recently in
* Corresponding author. Address: Chinese PLA General Hospital, Department of Orthopaedics, No. 28 Fuxing Road, Beijing 100853, China. Tel.: +86 13810745151; fax: +86 1068161218. E-mail address: [email protected] (Y. Wang).
biology for the fabrication of bone-repairing materials [6–8]. The core issue in the preparation of Col–HA composite lies in its structure, in particular the relative phase relationship between the two materials. But there are also a few hints of a fresh theme emerging from this work: the design of artificial nanostructures which can interact with and replace natural and biological materials [9]. It is still difficult to design supramolecular structures, particularly starting with designed molecules and forming bone-like objects between nanoscopic and macroscopic dimensions, which affects the composite’s properties, especially its mechanical properties [10]. Bone repair materials prepared by biomineralization differ from natural bones in the crystalline phase [11]. During mineralization of natural bones, organic matter like collagen is first secreted in areas where osteogenesis occurs. No local formation of hydroxyapatite crystals happens in the early stage of osteogenesis. Consequently, collagens form a cross-linking network in an interference-free environment, while association between collagen fibers is enhanced by proteoglycans. This allows collagen fibers in a relatively stable state, i.e., bone matrix gel, before deposition of hydroxyapatite. Meanwhile, a large number of negatively charged proteoglycans
1742-7061/$ - see front matter Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.12.042
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exist among and on the collagen fibers, providing a basis for chemical and electrical bonding. These proteoglycans serve as crystal nuclei during hydroxyapatite deposition [12,13], inducing HA crystallization along the length of collagen fibers [14]. However, in the preparation of Col–HA composites by in vitro biomineralization and self-assembly, no pre-cross-linking of collagen fibers is introduced to form bone matrix gel. To maintain the stability of collagen sol during reaction, cross-linking or bonding should be avoided between collagen fibers. When hydroxyapatite crystals precipitate in the collagen swelling solution, hydroxyapatite crystals grow using polar groups near the ends of collagen fibers as crystal nuclei. Since crystal nuclei localize at the ends of collagen fibers, association between collagen fibers is impaired; meanwhile, aggregation and elongation of the crystals themselves are also affected [15]. Precursor HA crystals deposit on the polar groups of Col fibers in the composites, which may block cell adhesion sites (such as RGD binding sites), thus changing the biological properties of Col–HA composites [16]. Based on the above analysis, this work presents a method of producing Col–HA composites making use of pre-crystallization. Prior to precipitation of collagen fibers, nano-HA crystals with independent crystal nuclei were prepared. These crystals were absorbed among collagen fibers by electric bonding rather than aggregated at the ends of collagen fibers. This process exposes collagen fibers, which have higher biocompatibility, on the composite exterior. Col–HA composites formed by pre-crystallization and mineralization crystallization were also characterized and then their biological performance in bone defect repair was compared both in vivo and in vitro to observe the crystalline phase on the biological properties of Col–HA composites. 2. Materials and methods 2.1. Materials Type I collagen (Col) was obtained from Department of Biomedical Engineering, Peking Union Medical College and Chinese Academy of Medical Science. H3PO4 and Ca(OH)2 used in this study were of analytical pure grade, obtained from Tianjin Chemical Reagent Company (Tianjin, China). 2.2. Preparation of Col–HA composites 2.2.1. Preparation of Col–HA composites by mineralization crystallization Five grams of type I Col was added to a solution of H3PO4 (59.7 mM). The mixture was agitated to ensure uniform distribution. Meanwhile, 199.1 mmol Ca(OH)2 was dispersed in 2 dm3 of distilled water and thoroughly mixed. Next, the Ca(OH)2 solution was added to the solution of swollen collagen at a ratio of 1:4 (Col:HA). The resultant mixture was allowed to react for 12 h. The precipitate was then freeze-dried using a lyophilizer under vacuum. Later, a solution of 0.25% glutaraldehyde was added to the precipitate. The cross-linking reaction proceeded for 2 h at room temperature. Subsequently, the cross-linked composites were washed with distilled water ten times and then relyophilized to yield a product referred to hereafter in this study as MC composite. 2.2.2. Preparation of Col–HA composites by pre-crystallization First, 199.1 mmol Ca(OH)2 was added to 2 dm3 of distilled water and mixed to ensure even distribution. Next, a solution of 59.7 mM H3PO4 was added to the Ca(OH)2 solution. The mixture was ultrasonicated and allowed to react for 2 h at room temperature, yielding a white solid precipitate. The precipitate was harvested by centrifugation at 800 rpm for 10 min and then transferred to a
lyophilizer to freeze dry under vacuum for 24 h, producing white solid HA. The HA was added to a 5% Col swollen solution (5.0 g) at a ratio of 1:4 (Col:HA). The mixture was agitated and allowed to react for 12 h. After that, the precipitate was placed in a lyophilizer to freeze dry under vacuum. Subsequently, a solution of 0.25% glutaraldehyde was added to facilitate the cross-linking reaction for 2 h at room temperature. The cross-linked material was washed with double distilled water 10 times and then relyophilized to yield a composite referred to hereafter in this study as PC composite. 2.3. Structural characterization 2.3.1. Scanning electron microscopy (SEM) MC and PC composites were sectioned, covered with gold and then examined with a Hitachi S-3500N scanning electron microscope to determine their surface features. 2.3.2. Transmission electron microscopy (TEM) To investigate the mineralization properties of the materials, samples were stained, embedded in epoxy resin and then sectioned. A JEM 2010 TEM was used with the positive stain uranyl acetate, which preferentially stains acidic groups. TEM showed higher electron density at the periphery of the fiber, revealing donut-shaped patterns indicating that only the outer portion of the fiber was stained. 2.4. Physicochemical characterization 2.4.1. Electron spectroscopy for chemical analysis (ESCA) The surface elements of MC and PC composites were characterized by electron spectroscopy for chemical analysis using a magnesium anode (Mg K = 1253.6 eV) with a survey scan range of 0– 1000 eV. All electron binding energies were referenced to the C1s hydrocarbon peak at 284.6 eV. 2.4.2. Water absorption of the Col–HA composites After weighing (Wd), three samples from each of the MC and PC composites were immersed in PBS solution (pH 7.2–7.4) at room temperature. The samples were extracted after 2 h and placed on filters to remove excess water. Subsequently, the scaffolds were weighed (Wh), and the amount of water absorbed (Wa) by each sample was calculated according to the equation:
Wa ¼
Wh Wd 100% Wd
ð1Þ
2.5. Evaluation of in vitro biological performance 2.5.1. Separation and seeding of bone marrow stromal stem cells According to the methods previously described [17,18], bone marrow stromal stem cells (BMSC) were isolated from the long bone of 6-week-old Wistar rats. Isolated cells were cultured in an incubator at 37 °C in an atmosphere of 5% CO2. The culture medium was replaced after 24 h of culture and then changed every 2 or 3 days. Four to five days later, when cells reached 80–90% confluence, cells were digested with 0.25% trypsin and passaged. Samples from the MC and PC composites (each with a diameter of 16 mm and a thickness of 3 mm) were each placed into a 24-well culture plate. Third or fourth generation BMSC in the exponential growth phase were digested with 0.25% trypsin and centrifuged. The pelleted cells were resuspended in DMEM culture medium for cell counting and trypan blue cell viability assays. The cell suspension was then adjusted to a concentration of 1.0 106 cells ml1. Subsequently, 1 ml of BMSC suspension was added to each well containing a composite or pure collagen. Cells
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were cultured at 37 °C in the presence of 5% CO2. The culture medium was replaced every 2 days [17,19]. 2.5.2. SEM analysis After 14 days of culture, the MC and PC composites seeded with BMSC were fixed in glutaraldehyde for 2 h and then post-fixed in osmic acid for 1 h. After critical point drying and gold sputtering, the MC and PC composites were analyzed by SEM (Hitachi S-3500). 2.5.3. Measurement of cell adhesion Before cell seeding, the scaffolds were presoaked in DMEM for 1 h. BMSC in the exponential growth phase were digested with 0.25% trypsin and harvested by centrifugation. After cell counting, the cells were resuspended in DMEM at a final concentration of 1.0 106 cells ml1. A 1-ml sample of the cell suspension was seeded into each scaffold and cultured at 37 °C with 5% CO2. These scaffolds were washed with PBS three times after 2, 4, 6 and 8 h. Scaffolds were then treated with 0.25% trypsin for 10 min and thoroughly washed with PBS. After centrifugation, the cells were collected and counted. Scaffolds were also stained with neutral red to confirm that all cells had been removed. The amount of cell adhesion (percent adhesion) was calculated according to the equation below (where Na is the number of the cells adhered, Ns is the number of cells seeded on the matrices and Nr is the number of cells removed during rinsing):
Na ðadheredÞ ¼
Ns ðseededÞ Nr ðrinsedÞ Ns ðseededÞ
ð2Þ
2.5.4. Cell growth curve Scaffolds seeded with BMSC were washed with PBS three times at days 1, 4 and 7. Scaffolds were then treated with 0.25% trypsin for 10 min and washed thoroughly with PBS. After centrifugation, the cells were collected and counted. The scaffolds were stained with neutral red to confirm that all cells had been removed. The cell growth curve was calculated according to Eq. (2). 2.5.5. Alkaline phosphatase activity assay Composites seeded with BMSC were treated with 0.25% trypsin. The cells were centrifuged at 1000 rpm for 5 min and washed twice with Tyrode’s solution. Next, the cells were resuspended in alkaline phosphatase incubation medium to a final concentration of 4 106 cells ml1. Osteoblasts and BMSC were lysed by ultrasonic homogenization. The lysate was incubated at 37 °C for 10 min. Optical densities of the osteoblasts and BMSC were measured at 405 nm using a spectrophotometer. 2.6. Repair of mandibular defects A bilateral mandibular fracture model was established in New Zealand rabbits. The MC and PC composites were implanted into either side of the defect areas. This experiment was approved by the Ethical Committee at the Chinese PLA General Hospital. Rabbits were anesthetized with sodium pentobarbital (30 mg kg1) and placed on their backs (supine). After skin preparation and disinfection, a longitudinal incision 2.5 cm in length was made in the right submandibular area to expose the middle and posterior segments of the mandibular bone. A periosteal elevator was used to perform subperiosteal dissection and fully expose the mandibular bone. Next, a fissure bur with cooling saline solution was used to create a wound surface measuring 10 4 3 mm in the mandibular bone where the two composites were implanted. The wound was washed with gentamicin (40,000 U) and stitched together. Rabbits were allowed free access to water and food after the surgery. Gentamicin (40,000 U) was administered intramuscularly daily for three consecutive days.
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2.6.1. Masson staining At 1, 2 and 4 months following implantation, three rabbits were sacrificed. Bone tissues from the implantation sites were collected, fixed in 10% glutaraldehyde, embedded in resin and sectioned. The sections were stained with Masson composition solution for 5 min and rinsed with 0.2% acetic acid solution. Next, the sections were treated with 5% phosphotungstic acid for 5–10 min and rinsed with 0.2% acetic acid solution. Then, the sections were stained in a light green solution and rinsed with 0.2% acetic acid solution twice. The stained sections were dehydrated in absolute alcohol, cleared in xylene and mounted with neutral gum. 2.6.2. Radionuclide bone scan Newly formed bone in the wound areas was observed under anesthesia by a radioactive nuclide bone scan at 1, 2 and 4 months after implantation. The radionuclide bone scan was performed with a gamma camera (Hamamatsu) after intravenous injection of 3 mG of 99mTc-MDP. Differences in bone defects between the vascular phase (0–30 s) and the delayed phase (2 h) were recorded and compared between the two composites. Regions of interest were selected in the bilateral submandibular areas to calculate radioactive nuclide uptake. 3. Results 3.1. Col–HA composite structure 3.1.1. SEM The results of SEM analysis of the two composites are presented in Fig. 1. There were large numbers of HA crystals on the collagen surface of the MC composite. Needle-like HA crystals showed irregular crystal orientations. However, the surface of the PC composite was relatively smooth. No perpendicularly needle-like HA crystals were apparent on the collagen surface. Macropores with diameters ranging from more than 10 to 300 lm were distributed on the surface of the two composites. Interconnected porous network structures were found within the macropores of the PC composite, while very few micropores were seen on the wall of the macropores in the MC composite. Amplification of the composites (Fig. 1C and D) revealed that the PC composite had an uneven surface with several HA crystals. However, these HA crystals did not aggregate into clusters. In comparison, the MC composite displayed a large number of needle-like HA crystals on the collagen surface of the MC composite which aggregated into clusters in contact with the collagen scaffold. 3.1.2. TEM Fig. 2A and C shows the structure of the MC Col–HA matrix. The composite was an intertwined assembly of fibers (i.e., bundles). Each bundle consisted of many Col fibrils surrounded by coraline-like HA nanocrystals. Positive-staining experiments indicated electron-dense HA nanocrystals packed around the acidic moieties of the peptide displayed on the surface of the fiber. HA was concentrated around the fibers, increasing their contrast. Some parts of the fibers were covered with crystalline mineral, but few plateshaped mineral polycrystals were visible along the fiber surface. HA appeared as discrete, electron-dense clusters on the surface. These HA nanocrystals, 50–100 nm in size, were situated among the Col fibrils and oriented in a slightly disorderly manner. The finer periodic bands of the collagen fibrils were not apparent, and the collagen fiber bundles in matrix were not compact. Some smaller ultracrystalline HA crystals were also dispersed among the collagen fibers (Fig. 2A). The crystals were plate shaped, with dimensions of 2–5 nm. The crystal sizes were relatively uniform, and the range was relatively confined.
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Fig. 1. SEM images of (A and C) MC and (B and D) PC composites. (On the MC matrix surface, HA particles aggregate into clusters, which are not observed on the PC matrix surface.) (A) 900; (B) 200; (C) 15,000; (D) 15,000.
Fig. 2B and D shows the structure of the PC Col–HA matrix. The matrix was an intertwined assembly of fibers (i.e., bundles) more than 20 lm in length, as shown in the figure. Each bundle consisted of many collagen fibrils surrounded by three polypeptide chains in a kinked triple helix (one-quarter stagger of the length produced a striation). At this level, the collagen fiber bundles appeared to be parallel and relatively compact. The striations were distinctly visible, and the arrangement of the rods was compact. Some plate-shaped polycrystalline mineral was seen along the fiber surfaces. The crystallites traversed both the gap and overlap zones, and the finer striations of the collagen fibrils were still discernible. The intra-fibrillar crystallites grew larger than the MC matrix, which were generally spaced with 64 nm periodicity. Very few plate-shaped ultracrystalline HA were observed among the collagen fibers (Fig. 2B). However, aggregation was observed in some regularly shaped HA polycrystals. These crystals were 10–20 nm in length and larger than those in the MC composite. 3.2. Physicochemical characterization 3.2.1. Electron spectroscopy for chemical analysis (ESCA) Fig. 3 shows the atomic ratios of elements Ca, O, P, C and N on the surface of the two composites. The C and N contents on the PC composite surface were markedly higher than those on the MC composite surface. In particular, the N content (from Col) increased from 1.92% in the MC composite to 13.13% in the PC composite, whereas the P and Ca contents decreased from 8.99% and 17.56% in the MC composite to 4.39% and 5.86% in the PC composite. These data suggest that HA precipitation within the collagen fibers and distribution of collagen fibers on the surface of the composites led to higher amounts of N (from Col) and lower levels of Ca and P (from HA) in the PC composite. The polypeptide chains of the collagen fibers consist of more than 300 repeats of the Gly-X-Y sequence, where X is often proline
and Y is often hydroxyproline (Hyp). Hyp contributes the only free hydroxyl group in collagen. The MC composite surface had a relatively high ratio of hydroxyl groups (11.87%), derived mainly from HA located on the surface of the collagen fibers. Conversely, few hydroxyl groups were found on the surface of the PC composite (1.57%). Thus, the hydroxyl group distribution was consistent with the element distribution, confirming the differences in spatial structures between the two composites. CH3C*–N groups were present mainly on the polar domains of collagen. The percentage of CH3C*–N groups was significantly higher in the PC composite than in the MC composite (16.34% vs 3.70% in total). As the primary polar groups in collagen materials, CH3C*–N groups can raise the hydrophilicity and cell adhesion capabilities of collagen materials. 3.2.2. Water absorption As shown by ESCA, the total number of the polar groups on the surface of the PC composite surface was higher than that on the MC composite surface. Polar groups are known to raise the surface energy on the composite surface, and a composite with a higher-energy surface can absorb more water compared with a composite with a low-energy surface. The water-absorbing capacity (Wa) of the two composites was compared. The MC composite had a mean Wa of 2.16, and the PC matrix had a mean Wa of 3.28. There was a significant difference in water absorption between the two composites at room temperature (p = 0.019 < 0.05) (see Table 1). 3.3. Evaluation of in vitro biological performance 3.3.1. Separation and culture of BMSC Trypan blue staining showed that up to 95% of cells were viable. When seeded onto culture flasks, most cells attached within 4 h. When the cells were passaged to the fourth generation, extracellular matrix deposits were found around the cells, suggesting that the cells were actively engaged in proliferation.
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Fig. 2. TEM images of the fiber matrix: (A and C) MC matrix; (B and D) PC matrix. Each bundle consists of multiple 300-nm Col fibrils surrounded by HA crystals. The transverse stripe structure is typical of collagen fibers and is caused by interlaced parallel collagen fibrils. HA is the electron dense dot: (A) 27,000; (B) 27,000; (C) 1850,000; (D) 1850,000.
3.3.2. Cell adhesion As demonstrated in Fig. 4, the number of BMSC adhering to the scaffold decreased successively with time in the PC and MC composites. Cells adhered to the scaffold predominantly between the 4th and 6th hour following seeding. HA decreased cell adhesion on the MC composites in comparison with the PC composites. Cell adhesion was higher on the PC composite than on the MC composite, mainly due to the non-blocked polar groups and cell adhesionpromoting sites such as RGD on the collagen surface of the PC composite. Serum-free DMEM culture medium was used to exclude the possible effects of adhesion proteins and cytokines in the serum. 3.3.3. Cell growth rate The cell growth curve (Fig. 5) shows that the addition of HA also decreased cell growth on the MC composite surface as compared with the PC composite surface. HA microcrystals were located on the surface of the MC composite, shielding the collagen that enhances cell growth. 3.3.4. Alkaline phosphatase activity assay BMSC alkaline phosphatase activity on the PC composite peaked at day 20 and declined thereafter. In contrast, BMSC alkaline phosphatase activity on the MC composite held at a constant low level. Col could have induced BMSC differentiation into osteoblasts that have higher alkaline phosphatase activity than BMSC (Fig. 6). The present study showed that BMSC differentiation into osteoblasts was improved on the PC composite surface when compared with the MC composite surface.
3.4. In vivo bone defect repair 3.4.1. Overview The effects of the composites on bone defect repair are presented in Fig. 7. No local infections occurred after composite implantation into the bone defect areas. At the end of the first month, massive new bone tissue formation with a capillary network was visible in the area treated with the MC composite. About a quarter of the bone defect area contained residual MC composite (arrow G). Meanwhile, large amounts of newly formed bone tissue filled the bone defect area that was treated with the PC composite. In the inferior border of the bone defect area, there was only a small amount of residual PC composite (arrow H) that established a weak connection with the surrounding tissues. A capillary network was seen on the surface of new bone tissues. At the end of the second month, in the area treated with the MC composite, most of the bone defect area was occupied by new bone tissue. In the border regions, only a little residual MC composite remained, connecting with the surrounding muscle tissues. Most of the bone defect area treated with the PC composite was filled with new bone tissue. A capillary network connecting with surrounding tissues was found in the cortical bone (arrow J). At the end of the fourth month, bone defect repair in the MC-implanted area was essentially completed, and a few capillaries were observed in the cortical bone with extensive connection to the surrounding tissues (arrow K). In the area treated with the PC composite, bone defect repair was also finished. There were a few capillaries with almost no connection to the surrounding tissue.
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Fig. 3. ESCA profiles of the PC and MC composites: (A) elemental peaks of the MC composite; (B) elemental peaks of the PC composite; (C) analysis of groups in C1s peak for the MC composite; (D) analysis of groups in C1s peak for the PC composite.
Table 1 Water absorption by PC and MC composites (g). Variable
Obs.
Mean
SE
SD
PC MC Combined Diff.
3 3 6
3.28 2.16 2.72 1.12
0.29 0.07 0.28 0.30
0.50 0.12 0.70
3.4.2. Masson staining The Masson staining results are shown in Fig. 8. In vivo implantation of the two bone repair composites induced differentiation of autologous BMSC and new bone formation in the bone defect areas. A large amount of type I collagen was secreted to aid in the repair of the bone defects. At the end of the first month, significant amounts of residual scaffold and disordered scaffold structures were observed. The structure of the MC composite was looser than that of the PC composite. Moreover, inflammatory infiltration and a large number of red blood cells were seen in the composite residual (arrow G). Inflammatory infiltration in the area treated with the PC composite was more pronounced than in that of the MC composite. Additionally, a large number of red blood cells were found in the PC-implanted area. By the end of the second month, the composites had been degraded into separate pieces. Multiple layers of the collagen deposits were found in the bone tissue. Layer-like bone lamellae and early bone trabeculae began to develop in certain regions. Blood vessels were noted in the areas surrounding bone trabeculae and
Fig. 4. Adhesion of BMSC on the PC and MC composites.
within the bone matrix (arrow H). More calcium deposits and bone trabeculae had formed in the area treated with the PC composite than in the MC composite area. In addition, new blood vessels in bone trabeculae and a large number of osteocytes in the bone matrix were found in the PC-treated area. At the end of the fourth month, both composites were fully degraded, and new bone trabeculae were formed (arrow J). An exten-
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PC-treated side in the central defect was still slightly higher than that of the MC-treated side at the end of the 4th and 12th weeks. The average ROI was 74.57 pixels in the PC-treated side and 61.64 in the MC-treated side at the end of the 4th week; 95.55 pixels in the PC-treated side and 93.68 pixels in the MC-treated side at the end of the 12th week.
4. Discussion
Fig. 5. BMSC growth curves on the PC and MC composites.
Fig. 6. BMSC alkaline phosphatase activity on the MC and PC composites at 405 nm.
sive capillary network existed between the bone trabeculae. Bone destruction and reconstruction occurred within the trabeculae. There were no marked differences in bone defect repair between the two composites. However, the extent of peripheral vascularization was higher in the PC-treated area. 3.4.3. Radionuclide bone scan Nuclide bone scans were used to measure 99mTc-MDP uptake at the bone defect sites (Fig. 9). Angiographic Phase I (0–30 s) images showed neovascularization around the bone defect sites. Bone-uptake Phase III (2 h) images showed reconstruction and new bone formation at 4, 8 and 12 weeks, which had a different appearance from the implant graft. In angiographic phase analysis (Tables 2–4 and Fig. 9), 99mTcMDP uptake was higher on the Col–CS–HA side than the Col–HA side after 4, 8 and 12 weeks (p = 0.00001 < 0.001). Radiographic evaluation in Phase III revealed that the radiopaque shadow of the PC-treated side in the central defect was higher than that of the MC-treated side. The average ROI in the PC-treated side was 113.64 pixels compared with 76.56 pixels in the MC-treated side at the end of the 8th week. Radiographic evaluation in Phase III also showed that the radiopaque shadow of the
An ideal biomaterial for bone defect repair can function as a scaffold to allow entry and attachment of host blood vessels and cells, facilitating formation of new bone and subsequent degradation and absorption of the implant. Therefore, the usefulness of a biomaterial for bone defect repair is largely determined by its capacity to repair bone defects in vivo. Bony tissues consist mainly of mineralized type I collagen fibrils. These have plate-like crystals of carbonated hydroxyapatite arranged in parallel layers along grooves or gap regions through the fibrils. HA exists chiefly in this form in natural bones [20]. During the biomineralization process, hydroxyapatite microcrystals formed between the ends of collagen fibers, using the polar groups on the collagen surface as crystal nuclei (Fig. 2). During in vitro biomineralization, SEM and TEM findings showed that HA did not form the parallel layers seen in natural bones [4,21]; instead, HA aggregated into crystal balls using polar groups as the core (Figs. 1 and 2). These observations are the difference between biomineralization and natural mineralization. Polar groups such as amino (–NH2), carboxyl (–COOH) and hydroxyl (–OH) groups are thereby blocked on the surface of the collagen fibers. Therefore, in mineralization crystallization, the existence of HA affects the biological properties of collagen [22]. In the present study, the spatial relationship between collagen and HA in Col–HA composites was altered by adjusting the crystalline phase. In the PC composite, HA microcrystals do not precipitate at the ends of the collagen fibers and, thus, HA does not affect fiber polymerization, and active groups existed on the collagen surface (Fig. 2). This strategy draws on the design of layer by layer [23], yet more streamlined in process. As seen by SEM and TEM, the Col–HA composites generated by mineralization crystallization or pre-crystallization exhibited different pore diameters, spatial arrangement and fiber thicknesses, owing to their different structures. In the PC composite, HA microcrystals are located within the collagen fiber scaffold; in the MC composite, however, HA microcrystals are exposed on the surface of the collagen scaffold. This finding is confirmed by ESCA results showing that HA microcrystals deposited on the collagen fiber surface in the MC composite contain N 1.92%, P 8.99% and Ca 17.56%. This suggests that HA, the source of Ca, is deposited on the surface of the collagen fibers, while the collagen fibers, the source of N, are covered. In comparison, when HA microcrystals are deposited within the collagen scaffold in the PC composite, the N content rises to 13.13%, and the P and Ca contents decrease to 4.39% and 5.86%, respectively. Furthermore, the PC composite exhibits a more regular scaffold structure, a smoother surface and more homogeneous pore size [24,25]. Compared with the MC composite, the PC composite surface has more polar groups (–NH2 and –COOH) derived mainly from the ends of the collagen fibers. A previous study showed that, during biological mineralization, these polar groups serve as nuclei in HA mineral crystallization [26]. Consequently, during fabrication of the Col–HA composites by mineralization crystallization, these polar groups are shielded by HA crystals (as shown by SEM). The number of polar groups on these composite surfaces is reduced, attenuating the hydrophilicity and the cell adhesion capacity of these bone repair biomaterials.
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Fig. 7. In vivo repair of rabbit mandibular defects: (A) 1 month after implantation of the MC composite; (B) 1 month after implantation of the PC composite; (C) 2 months after implantation of the MC composite; (D) 2 months after implantation of the PC composite; (E) 4 months after implantation of the MC composite; (F) 4 months after implantation of the PC composite.
In general, a hydrophobic surface absorbs fewer proteins than a hydrophilic surface, which may lead to denaturation of albumin. As seen in Table 1, the PC composite absorbs more water than the MC composite. This may be partly explained by the fact that Col, which is rich in the polar groups, is situated on the composite surface, thus increasing the number of proteins. Therefore, enhancing the hydrophilicity of bone repair materials increases their capacity for water absorption, promoting biocompatibility [27–29]. Cell culture serves as an important tool for studying material biocompatibility. As presented in Figs. 4–6, both composites provide a favorable microenvironment for the growth and functioning of BMSC after 3 weeks of culture. However, significant differences are seen with respect to cell shape and cell concentration. In particular, osteoblasts are distributed more uniformly and grow at a higher density on the PC composite. In addition, BMSC adhere better and excrete larger amounts of extracellular matrix on the PC composite. BMSC adhesion and growth are superior on the PC composite surface in comparison with the MC composite surface. The number of BMSC on the PC composite surface is similar to that of BMSC grown on the pure collagen scaffold. In addition, alkaline phosphatase activity is higher in the PC com-
posite than in the MC composite, suggesting that BMSC in the PC composite are in direct contact with the collagen fibers on the composite surface, accelerating their differentiation into osteoblasts induced by the active sites on the surface of the collagen fibers. Therefore, a large amount of extracellular matrix is formed. In general, both composites induce the differentiation of BMSC into osteoblasts and the formation of new bones that eventually replace artificial implants and repair bone defects. Nevertheless, differences between the two composites are observed by macroscopic observation and radionuclide bone scans. Macroscopic observation reveals that, during the first and second months after implantation, bone defect repair proceeds significantly faster in the PC composite. Furthermore, Masson staining at the end of the first month shows that both composites essentially maintain their morphological integrity, despite the initiation of decomposition. Disintegration and absorption together with creeping substitution occur in both composites. Substitution occurs mainly in the early stages during the formation of new vascular channels within the composites. By the end of the second month, most of the compos-
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Fig. 8. Masson staining of rabbit mandibular defects (A) 1 month after implantation of the MC composite; (B) 1 month after implantation of the PC composite; (C) 2 months after implantation of the MC composite; (D) 2 months after implantation of the PC composite; (E) 4 months after implantation of the MC composite; (F) 4 months after implantation of the PC composite (100).
ites decompose, leaving only scattered fragments. Substantial amounts of layered collagen precipitate in bone tissues where bone lamellae form, indicating that most of the bone repair is completed during the second month in both composites. This finding is consistent with the radionuclide bone scan results. At the end of the fourth month, the composites are completely decomposed; bone trabeculae formation and reconstruction are seen; osteocytes and bone lacuna are found in the bone trabeculae. The fact that the hydroxyapatite crystals grow with their c-axes oriented along the long axes of the nanofibers could be of interest in the design of materials. The present paper has described the nanostructure of Col–HA composite by self-assembly. HA assembled around the polar groups of collagen fiber in clusters. In pure collagen scaffold, collagen fibers can be cross-linked by the formation of intermolecular disulfide bonds upon oxidation, which results in a chemically robust fiber. However, in the scaffold derived from biomineralization, HA aggregation affected Col cross-linking, thereby affecting the overall structure of the scaffold. As shown in the present study, deficiencies in the biological performance of the scaffold prepared by biomineralization are mani-
fested in several aspects. First, loose association between fibers and poor mechanical performance led to irregular spatial structures. Second, polar groups on the surface of collagen fibers were blocked, resulting in lower hydrophilicity. Third, HA aggregation on Col polar groups blocked cell attachment sites such as RGD. As a result, the overall biological performance of HA–Col composites is affected. Comparison of such materials with pre-crystallized material also proves this finding. In contrast, in the pre-crystallization, HA crystallizes before Col precipitation and thus does not hide the polar groups on the surface of Col. Because of the different order of crystallization, the two composites demonstrate different biological properties.
5. Conclusion It can be concluded that HA placement on the collagen fibers affected the biological performance of Col–HA composites. Pre-crystallization made HA scatter among collagen fibers, creating a better structure for bone defect repair in comparison with MC Col–HA composites.
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Fig. 9. Radionuclide bone scans of rabbit mandibular defects following repair with PC and MC composites: (A–C) angiographic Phase I; (D–F) bone-uptake Phase III. (A and D) 1 month after implantation; (B and E) 2 months after implantation; (C and F) 4 months after implantation (left implant: the PC composite; right implant: the MC composite).
Table 2 The 99mTc-MDP taken up in angiographic phase analysis at 1 M.
Table 3 99m Tc-MDP taken up in angiographic phase analysis at 2 M.
Variable
Obs.
Mean
SE
SD
Variable
Obs.
Mean
SE
SD
PC MC Diff.
30 30
1610 1413 197
441 385
81 70
PC MC Diff.
30 30
2390 2063 327
469 455
86 83
L. Zhang et al. / Acta Biomaterialia 6 (2010) 2189–2199 Table 4 99m Tc-MDP taken up in angiographic phase analysis at 4 M. Variable
Obs.
Mean
SE
SD
PC MC Diff.
30 30
2942 2326 616
801 634
146 116
Acknowledgements This work was supported by grants from the National Natural Science Foundation of China (30700848 and 30772443) and the Natural Science Foundation of Beijing, China (7042058 and 7082090). Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figures 3, 7–9, are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi: 10.1016/j.actbio.2009.12.042. References [1] Keating JF, McQueen MM. Substitute for autologous bone graft in orthopaedic trauma. J Bone Joint Surgery 2001;83B:3–8. [2] Crane GM, Ishaug SL, Mikos AG. Bone tissue engineering. Nat Med 1995;1:1322–4. [3] Damien CJ, Parsons JR. Bone graft and bone graft substitutes: a review of current technology and applications. J Appl Biomater 1991;2:187–208. [4] Lees S. A model for the distribution of HAP crystallites in bone—an hypothesis. Calcif Tissue Int 1979;27:53–6. [5] Prostak KS, Lees S. Visualization of crystal-matrix structure. In situ demineralization of mineralized Turkey leg tendon and bone. Calcif Tissue Int 1996;59:474–9. [6] Hartgerink Jeffrey D, Beniash Elia, Stupp Samuel I. Self-assembly and mineralization of peptide-amphiphile nanofibers. Science 2001;23:1684–8. [7] Rammelt S, Schulze E, Witt M, Petsch E, Biewener A, Pompe W, et al. Collagen type I increases bone remodelling around hydroxyapatite implants in the rat tibia. Cells Tissues Organs 2004;178:146–57. [8] Du C, Falini G, Fermani S, Abbott C, Moradian-Oldak J. Supramolecular assembly of amelogenin nanospheres into birefringent microribbons. Science 2005;307:1419–20. [9] Taton TA. Nanotechnology. Boning up on biology. Nature 2001;412(6846): 491–2.
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[10] Kanungo BP, Gibson LJ. Density–property relationships in mineralized collagen–glycosaminoglycan scaffolds. Acta Biomater 2009;5(4):1006–18 [Epub 2008 Dec 11]. [11] He G, Dahl T, Veis A, George A. Nucleation of apatite crystals in vitro by selfassembled dentin matrix protein 1. Nat Mater 2003;2(8):552–8. [12] Hunter GK, Hauschka PV, Poole AR, Rosenberg LC, Goldberg HA. Nucleation and inhibition of hydroxyapatite formation by mineralized tissue proteins. Biochem J 1996;317:59–64. [13] Boskey AL. Osteopontin and related phosphorylated sialoproteins: effects on mineralization. Ann NY Acad Sci 1995;760:249–56. [14] Weiner S, Traub W. Bone structure: from angstroms to microns. FASEB J 1992;6(3):879–85 [Review]. [15] Butler WT, Ritchie H. The nature and functional significance of dentin extracellular matrix proteins. Int J Dev Biol 1995;39:169–79. [16] Zhang L, Hu M, Wang M, Li Y, Chen H, Chu C, Jiang H. Evaluation of modifying collagen matrix with RGD peptide through periodates oxidation. J Biomed Mater Res 2005;73A:468–75. [17] Martin I, Shastri VP, Padera RF, Yang J, Mackay AJ, Langer R, et al. Selective differentiation of mammalian bone marrow stromal cells cultured on threedimensional polymer foams. J Biomed Mater Res 2001;55:229–35. [18] Mills BG, Singer FR, Weiner LP. Long-term culture of cells from bone affected by Paget’s disease. Calcif Tissue Int 1979;29:79–87. [19] Ishaug SL, Crane GM, Miller MJ, Yasko AW, Yaszemski MJ, Mikos AG. Bone formation by three-dimensional stromal osteoblast culture in biodegradable polymer scaffolds. J Biomed Mater Res 1997;36:17–28. [20] Robinson RA, Watson ML. Collagen–crystal relationships in bone as seen in the electron microscope. Anat Rec 1952;114:383–410. [21] Hellmich C, Ulm FJ. Average hydroxyapatite concentration is uniform in the extracollagenous ultrastructure of mineralized tissues: evidence at the 1–10microm scale. Biomech Model Mechanobiol 2003;2(1):21–36. [22] Blair HC, Zaidi M, Schlesinger PH. Mechanisms balancing skeletal matrix synthesis and degradation. Biochem J 2002;364(Pt 2):329–41. [23] Ficai A, Andronescu E, Voicu G, Manzu D, Ficai M. Layer by layer deposition of hydroxyapatite onto the collagen matrix. Mater Sci Eng C 2009;29: 2217–20. [24] Pieper JS, Oosterhof A. Preparation and characterization of porous crosslinked collagenous matrices containing bioavailable chondroitin sulphate. Biomaterials 1999;20:847–58. [25] Madihally Sundararajan V, Howard WT. Porous chitosan scaffolds for tissue engineering. Biomaterials 1999;20:1133–43. [26] Gajjeraman S, Narayanan K, Hao J, Qin C, George A. Matrix macromolecules in hard tissues control the nucleation and hierarchical assembly of hydroxyapatite. J Biol Chem 2007;282:1193–204. [27] Stephen W, Wolfie T. Bone structure: from angstroms to microns. FASEB J 1992;6:879–85. [28] Baht GS, Hunter GK, Goldberg HA. Bone sialoprotein–collagen interaction promotes hydroxyapatite nucleation. Matrix Biol 2008;27:600–8. [29] Brandao-Burch A, Utting JC, Orriss IR, Arnett TR. Acidosis inhibits bone formation by osteoblasts in vitro by preventing mineralization. Calcif Tissue Int 2005;77:167–74.
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Effects of crystalline phase on the biological properties of collagen–hydroxyapatite composites L. Zhang, P. Tang, M. Xu, W. Zhang, W. Chai, Y. Wang * Chinese PLA General Hospital, Department of Orthopaedics, Beijing 100853, China
a r t i c l e
i n f o
Article history: Received 18 October 2009 Received in revised form 17 December 2009 Accepted 21 December 2009 Available online 28 December 2009 Keywords: Hydroxyapatite Collagen Crystalline phase Biomineralization Pre-crystallization
a b s t r a c t The objective of this study was to investigate the effects of spatial structure and crystalline phase on the biological performance of collagen–hydroxyapatite (Col–HA) composite prepared by biomineralization crystallization. Two types of Col–HA composites were prepared using mineralization crystallization (MC composites) and pre-crystallization (PC composites), respectively. Structural characteristics were analyzed by scanning electron microscopy and transmission electron microscopy. Surface elemental compositions were measured by electron spectroscopy for chemical analysis (ESCA). These composites were used in in vivo repair of bone defects. The effects of the crystalline phase on the biological performance of Col–HA composites were investigated using radionuclide bone scan, histopathology and morphological observation. It was observed that in MC composites, HA was located on the surface of the collagen fibers and aggregated into crystal balls, whereas HA in PC composites was scattered among the collagen fibers. ESCA showed that phosphorus and calcium were 8.99% and 17.56% on MC composite surface, compared with 4.39% and 5.86% on the PC composite surface. In vivo bone defect repair experiments revealed that radionuclide uptake was significantly higher in the area implanted with the PC composite than in the contralateral area implanted with the MC composite. Throughout the whole repair process, the PC composite proved to be superior to the MC composite with regard to capillary-forming capacity and the amount of newly formed bone tissue. So it could be concluded that HA placement on collagen fibers affected the biological performance of Col–HA composites. Pre-crystallization made HA scattered among collagen fibers, creating a better structure for bone defect repair in comparison with MC Col–HA composites. Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Great efforts have been made to develop biomaterials capable of repairing bone defects resulting from various causes. Bone tissue is a particularly complex composite because it contains multiple levels of organic and inorganic nanophases. At the lowest level of this hierarchy is a three-dimensional composite made up of collagen triple helices, and the hydroxyapatite crystals grow in some different form: the majority of the mineral lies in such a way that its c-axes are oriented along the long axes of the fibrils (called intra-fibrillar mineral) [1–3], some lies between the fibrils, the c-axes of which are perpendicular to the collagen molecular axis (called inter-fibrillar mineral or extra-fibrillar mineral) [4,5]. Many researchers have attempted to construct composites of hydroxyapatite and collagen (Col–HA) for bone defect repair. Self-assembly and biomineralization have been used recently in
* Corresponding author. Address: Chinese PLA General Hospital, Department of Orthopaedics, No. 28 Fuxing Road, Beijing 100853, China. Tel.: +86 13810745151; fax: +86 1068161218. E-mail address: [email protected] (Y. Wang).
biology for the fabrication of bone-repairing materials [6–8]. The core issue in the preparation of Col–HA composite lies in its structure, in particular the relative phase relationship between the two materials. But there are also a few hints of a fresh theme emerging from this work: the design of artificial nanostructures which can interact with and replace natural and biological materials [9]. It is still difficult to design supramolecular structures, particularly starting with designed molecules and forming bone-like objects between nanoscopic and macroscopic dimensions, which affects the composite’s properties, especially its mechanical properties [10]. Bone repair materials prepared by biomineralization differ from natural bones in the crystalline phase [11]. During mineralization of natural bones, organic matter like collagen is first secreted in areas where osteogenesis occurs. No local formation of hydroxyapatite crystals happens in the early stage of osteogenesis. Consequently, collagens form a cross-linking network in an interference-free environment, while association between collagen fibers is enhanced by proteoglycans. This allows collagen fibers in a relatively stable state, i.e., bone matrix gel, before deposition of hydroxyapatite. Meanwhile, a large number of negatively charged proteoglycans
1742-7061/$ - see front matter Ó 2010 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.12.042
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exist among and on the collagen fibers, providing a basis for chemical and electrical bonding. These proteoglycans serve as crystal nuclei during hydroxyapatite deposition [12,13], inducing HA crystallization along the length of collagen fibers [14]. However, in the preparation of Col–HA composites by in vitro biomineralization and self-assembly, no pre-cross-linking of collagen fibers is introduced to form bone matrix gel. To maintain the stability of collagen sol during reaction, cross-linking or bonding should be avoided between collagen fibers. When hydroxyapatite crystals precipitate in the collagen swelling solution, hydroxyapatite crystals grow using polar groups near the ends of collagen fibers as crystal nuclei. Since crystal nuclei localize at the ends of collagen fibers, association between collagen fibers is impaired; meanwhile, aggregation and elongation of the crystals themselves are also affected [15]. Precursor HA crystals deposit on the polar groups of Col fibers in the composites, which may block cell adhesion sites (such as RGD binding sites), thus changing the biological properties of Col–HA composites [16]. Based on the above analysis, this work presents a method of producing Col–HA composites making use of pre-crystallization. Prior to precipitation of collagen fibers, nano-HA crystals with independent crystal nuclei were prepared. These crystals were absorbed among collagen fibers by electric bonding rather than aggregated at the ends of collagen fibers. This process exposes collagen fibers, which have higher biocompatibility, on the composite exterior. Col–HA composites formed by pre-crystallization and mineralization crystallization were also characterized and then their biological performance in bone defect repair was compared both in vivo and in vitro to observe the crystalline phase on the biological properties of Col–HA composites. 2. Materials and methods 2.1. Materials Type I collagen (Col) was obtained from Department of Biomedical Engineering, Peking Union Medical College and Chinese Academy of Medical Science. H3PO4 and Ca(OH)2 used in this study were of analytical pure grade, obtained from Tianjin Chemical Reagent Company (Tianjin, China). 2.2. Preparation of Col–HA composites 2.2.1. Preparation of Col–HA composites by mineralization crystallization Five grams of type I Col was added to a solution of H3PO4 (59.7 mM). The mixture was agitated to ensure uniform distribution. Meanwhile, 199.1 mmol Ca(OH)2 was dispersed in 2 dm3 of distilled water and thoroughly mixed. Next, the Ca(OH)2 solution was added to the solution of swollen collagen at a ratio of 1:4 (Col:HA). The resultant mixture was allowed to react for 12 h. The precipitate was then freeze-dried using a lyophilizer under vacuum. Later, a solution of 0.25% glutaraldehyde was added to the precipitate. The cross-linking reaction proceeded for 2 h at room temperature. Subsequently, the cross-linked composites were washed with distilled water ten times and then relyophilized to yield a product referred to hereafter in this study as MC composite. 2.2.2. Preparation of Col–HA composites by pre-crystallization First, 199.1 mmol Ca(OH)2 was added to 2 dm3 of distilled water and mixed to ensure even distribution. Next, a solution of 59.7 mM H3PO4 was added to the Ca(OH)2 solution. The mixture was ultrasonicated and allowed to react for 2 h at room temperature, yielding a white solid precipitate. The precipitate was harvested by centrifugation at 800 rpm for 10 min and then transferred to a
lyophilizer to freeze dry under vacuum for 24 h, producing white solid HA. The HA was added to a 5% Col swollen solution (5.0 g) at a ratio of 1:4 (Col:HA). The mixture was agitated and allowed to react for 12 h. After that, the precipitate was placed in a lyophilizer to freeze dry under vacuum. Subsequently, a solution of 0.25% glutaraldehyde was added to facilitate the cross-linking reaction for 2 h at room temperature. The cross-linked material was washed with double distilled water 10 times and then relyophilized to yield a composite referred to hereafter in this study as PC composite. 2.3. Structural characterization 2.3.1. Scanning electron microscopy (SEM) MC and PC composites were sectioned, covered with gold and then examined with a Hitachi S-3500N scanning electron microscope to determine their surface features. 2.3.2. Transmission electron microscopy (TEM) To investigate the mineralization properties of the materials, samples were stained, embedded in epoxy resin and then sectioned. A JEM 2010 TEM was used with the positive stain uranyl acetate, which preferentially stains acidic groups. TEM showed higher electron density at the periphery of the fiber, revealing donut-shaped patterns indicating that only the outer portion of the fiber was stained. 2.4. Physicochemical characterization 2.4.1. Electron spectroscopy for chemical analysis (ESCA) The surface elements of MC and PC composites were characterized by electron spectroscopy for chemical analysis using a magnesium anode (Mg K = 1253.6 eV) with a survey scan range of 0– 1000 eV. All electron binding energies were referenced to the C1s hydrocarbon peak at 284.6 eV. 2.4.2. Water absorption of the Col–HA composites After weighing (Wd), three samples from each of the MC and PC composites were immersed in PBS solution (pH 7.2–7.4) at room temperature. The samples were extracted after 2 h and placed on filters to remove excess water. Subsequently, the scaffolds were weighed (Wh), and the amount of water absorbed (Wa) by each sample was calculated according to the equation:
Wa ¼
Wh Wd 100% Wd
ð1Þ
2.5. Evaluation of in vitro biological performance 2.5.1. Separation and seeding of bone marrow stromal stem cells According to the methods previously described [17,18], bone marrow stromal stem cells (BMSC) were isolated from the long bone of 6-week-old Wistar rats. Isolated cells were cultured in an incubator at 37 °C in an atmosphere of 5% CO2. The culture medium was replaced after 24 h of culture and then changed every 2 or 3 days. Four to five days later, when cells reached 80–90% confluence, cells were digested with 0.25% trypsin and passaged. Samples from the MC and PC composites (each with a diameter of 16 mm and a thickness of 3 mm) were each placed into a 24-well culture plate. Third or fourth generation BMSC in the exponential growth phase were digested with 0.25% trypsin and centrifuged. The pelleted cells were resuspended in DMEM culture medium for cell counting and trypan blue cell viability assays. The cell suspension was then adjusted to a concentration of 1.0 106 cells ml1. Subsequently, 1 ml of BMSC suspension was added to each well containing a composite or pure collagen. Cells
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were cultured at 37 °C in the presence of 5% CO2. The culture medium was replaced every 2 days [17,19]. 2.5.2. SEM analysis After 14 days of culture, the MC and PC composites seeded with BMSC were fixed in glutaraldehyde for 2 h and then post-fixed in osmic acid for 1 h. After critical point drying and gold sputtering, the MC and PC composites were analyzed by SEM (Hitachi S-3500). 2.5.3. Measurement of cell adhesion Before cell seeding, the scaffolds were presoaked in DMEM for 1 h. BMSC in the exponential growth phase were digested with 0.25% trypsin and harvested by centrifugation. After cell counting, the cells were resuspended in DMEM at a final concentration of 1.0 106 cells ml1. A 1-ml sample of the cell suspension was seeded into each scaffold and cultured at 37 °C with 5% CO2. These scaffolds were washed with PBS three times after 2, 4, 6 and 8 h. Scaffolds were then treated with 0.25% trypsin for 10 min and thoroughly washed with PBS. After centrifugation, the cells were collected and counted. Scaffolds were also stained with neutral red to confirm that all cells had been removed. The amount of cell adhesion (percent adhesion) was calculated according to the equation below (where Na is the number of the cells adhered, Ns is the number of cells seeded on the matrices and Nr is the number of cells removed during rinsing):
Na ðadheredÞ ¼
Ns ðseededÞ Nr ðrinsedÞ Ns ðseededÞ
ð2Þ
2.5.4. Cell growth curve Scaffolds seeded with BMSC were washed with PBS three times at days 1, 4 and 7. Scaffolds were then treated with 0.25% trypsin for 10 min and washed thoroughly with PBS. After centrifugation, the cells were collected and counted. The scaffolds were stained with neutral red to confirm that all cells had been removed. The cell growth curve was calculated according to Eq. (2). 2.5.5. Alkaline phosphatase activity assay Composites seeded with BMSC were treated with 0.25% trypsin. The cells were centrifuged at 1000 rpm for 5 min and washed twice with Tyrode’s solution. Next, the cells were resuspended in alkaline phosphatase incubation medium to a final concentration of 4 106 cells ml1. Osteoblasts and BMSC were lysed by ultrasonic homogenization. The lysate was incubated at 37 °C for 10 min. Optical densities of the osteoblasts and BMSC were measured at 405 nm using a spectrophotometer. 2.6. Repair of mandibular defects A bilateral mandibular fracture model was established in New Zealand rabbits. The MC and PC composites were implanted into either side of the defect areas. This experiment was approved by the Ethical Committee at the Chinese PLA General Hospital. Rabbits were anesthetized with sodium pentobarbital (30 mg kg1) and placed on their backs (supine). After skin preparation and disinfection, a longitudinal incision 2.5 cm in length was made in the right submandibular area to expose the middle and posterior segments of the mandibular bone. A periosteal elevator was used to perform subperiosteal dissection and fully expose the mandibular bone. Next, a fissure bur with cooling saline solution was used to create a wound surface measuring 10 4 3 mm in the mandibular bone where the two composites were implanted. The wound was washed with gentamicin (40,000 U) and stitched together. Rabbits were allowed free access to water and food after the surgery. Gentamicin (40,000 U) was administered intramuscularly daily for three consecutive days.
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2.6.1. Masson staining At 1, 2 and 4 months following implantation, three rabbits were sacrificed. Bone tissues from the implantation sites were collected, fixed in 10% glutaraldehyde, embedded in resin and sectioned. The sections were stained with Masson composition solution for 5 min and rinsed with 0.2% acetic acid solution. Next, the sections were treated with 5% phosphotungstic acid for 5–10 min and rinsed with 0.2% acetic acid solution. Then, the sections were stained in a light green solution and rinsed with 0.2% acetic acid solution twice. The stained sections were dehydrated in absolute alcohol, cleared in xylene and mounted with neutral gum. 2.6.2. Radionuclide bone scan Newly formed bone in the wound areas was observed under anesthesia by a radioactive nuclide bone scan at 1, 2 and 4 months after implantation. The radionuclide bone scan was performed with a gamma camera (Hamamatsu) after intravenous injection of 3 mG of 99mTc-MDP. Differences in bone defects between the vascular phase (0–30 s) and the delayed phase (2 h) were recorded and compared between the two composites. Regions of interest were selected in the bilateral submandibular areas to calculate radioactive nuclide uptake. 3. Results 3.1. Col–HA composite structure 3.1.1. SEM The results of SEM analysis of the two composites are presented in Fig. 1. There were large numbers of HA crystals on the collagen surface of the MC composite. Needle-like HA crystals showed irregular crystal orientations. However, the surface of the PC composite was relatively smooth. No perpendicularly needle-like HA crystals were apparent on the collagen surface. Macropores with diameters ranging from more than 10 to 300 lm were distributed on the surface of the two composites. Interconnected porous network structures were found within the macropores of the PC composite, while very few micropores were seen on the wall of the macropores in the MC composite. Amplification of the composites (Fig. 1C and D) revealed that the PC composite had an uneven surface with several HA crystals. However, these HA crystals did not aggregate into clusters. In comparison, the MC composite displayed a large number of needle-like HA crystals on the collagen surface of the MC composite which aggregated into clusters in contact with the collagen scaffold. 3.1.2. TEM Fig. 2A and C shows the structure of the MC Col–HA matrix. The composite was an intertwined assembly of fibers (i.e., bundles). Each bundle consisted of many Col fibrils surrounded by coraline-like HA nanocrystals. Positive-staining experiments indicated electron-dense HA nanocrystals packed around the acidic moieties of the peptide displayed on the surface of the fiber. HA was concentrated around the fibers, increasing their contrast. Some parts of the fibers were covered with crystalline mineral, but few plateshaped mineral polycrystals were visible along the fiber surface. HA appeared as discrete, electron-dense clusters on the surface. These HA nanocrystals, 50–100 nm in size, were situated among the Col fibrils and oriented in a slightly disorderly manner. The finer periodic bands of the collagen fibrils were not apparent, and the collagen fiber bundles in matrix were not compact. Some smaller ultracrystalline HA crystals were also dispersed among the collagen fibers (Fig. 2A). The crystals were plate shaped, with dimensions of 2–5 nm. The crystal sizes were relatively uniform, and the range was relatively confined.
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Fig. 1. SEM images of (A and C) MC and (B and D) PC composites. (On the MC matrix surface, HA particles aggregate into clusters, which are not observed on the PC matrix surface.) (A) 900; (B) 200; (C) 15,000; (D) 15,000.
Fig. 2B and D shows the structure of the PC Col–HA matrix. The matrix was an intertwined assembly of fibers (i.e., bundles) more than 20 lm in length, as shown in the figure. Each bundle consisted of many collagen fibrils surrounded by three polypeptide chains in a kinked triple helix (one-quarter stagger of the length produced a striation). At this level, the collagen fiber bundles appeared to be parallel and relatively compact. The striations were distinctly visible, and the arrangement of the rods was compact. Some plate-shaped polycrystalline mineral was seen along the fiber surfaces. The crystallites traversed both the gap and overlap zones, and the finer striations of the collagen fibrils were still discernible. The intra-fibrillar crystallites grew larger than the MC matrix, which were generally spaced with 64 nm periodicity. Very few plate-shaped ultracrystalline HA were observed among the collagen fibers (Fig. 2B). However, aggregation was observed in some regularly shaped HA polycrystals. These crystals were 10–20 nm in length and larger than those in the MC composite. 3.2. Physicochemical characterization 3.2.1. Electron spectroscopy for chemical analysis (ESCA) Fig. 3 shows the atomic ratios of elements Ca, O, P, C and N on the surface of the two composites. The C and N contents on the PC composite surface were markedly higher than those on the MC composite surface. In particular, the N content (from Col) increased from 1.92% in the MC composite to 13.13% in the PC composite, whereas the P and Ca contents decreased from 8.99% and 17.56% in the MC composite to 4.39% and 5.86% in the PC composite. These data suggest that HA precipitation within the collagen fibers and distribution of collagen fibers on the surface of the composites led to higher amounts of N (from Col) and lower levels of Ca and P (from HA) in the PC composite. The polypeptide chains of the collagen fibers consist of more than 300 repeats of the Gly-X-Y sequence, where X is often proline
and Y is often hydroxyproline (Hyp). Hyp contributes the only free hydroxyl group in collagen. The MC composite surface had a relatively high ratio of hydroxyl groups (11.87%), derived mainly from HA located on the surface of the collagen fibers. Conversely, few hydroxyl groups were found on the surface of the PC composite (1.57%). Thus, the hydroxyl group distribution was consistent with the element distribution, confirming the differences in spatial structures between the two composites. CH3C*–N groups were present mainly on the polar domains of collagen. The percentage of CH3C*–N groups was significantly higher in the PC composite than in the MC composite (16.34% vs 3.70% in total). As the primary polar groups in collagen materials, CH3C*–N groups can raise the hydrophilicity and cell adhesion capabilities of collagen materials. 3.2.2. Water absorption As shown by ESCA, the total number of the polar groups on the surface of the PC composite surface was higher than that on the MC composite surface. Polar groups are known to raise the surface energy on the composite surface, and a composite with a higher-energy surface can absorb more water compared with a composite with a low-energy surface. The water-absorbing capacity (Wa) of the two composites was compared. The MC composite had a mean Wa of 2.16, and the PC matrix had a mean Wa of 3.28. There was a significant difference in water absorption between the two composites at room temperature (p = 0.019 < 0.05) (see Table 1). 3.3. Evaluation of in vitro biological performance 3.3.1. Separation and culture of BMSC Trypan blue staining showed that up to 95% of cells were viable. When seeded onto culture flasks, most cells attached within 4 h. When the cells were passaged to the fourth generation, extracellular matrix deposits were found around the cells, suggesting that the cells were actively engaged in proliferation.
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Fig. 2. TEM images of the fiber matrix: (A and C) MC matrix; (B and D) PC matrix. Each bundle consists of multiple 300-nm Col fibrils surrounded by HA crystals. The transverse stripe structure is typical of collagen fibers and is caused by interlaced parallel collagen fibrils. HA is the electron dense dot: (A) 27,000; (B) 27,000; (C) 1850,000; (D) 1850,000.
3.3.2. Cell adhesion As demonstrated in Fig. 4, the number of BMSC adhering to the scaffold decreased successively with time in the PC and MC composites. Cells adhered to the scaffold predominantly between the 4th and 6th hour following seeding. HA decreased cell adhesion on the MC composites in comparison with the PC composites. Cell adhesion was higher on the PC composite than on the MC composite, mainly due to the non-blocked polar groups and cell adhesionpromoting sites such as RGD on the collagen surface of the PC composite. Serum-free DMEM culture medium was used to exclude the possible effects of adhesion proteins and cytokines in the serum. 3.3.3. Cell growth rate The cell growth curve (Fig. 5) shows that the addition of HA also decreased cell growth on the MC composite surface as compared with the PC composite surface. HA microcrystals were located on the surface of the MC composite, shielding the collagen that enhances cell growth. 3.3.4. Alkaline phosphatase activity assay BMSC alkaline phosphatase activity on the PC composite peaked at day 20 and declined thereafter. In contrast, BMSC alkaline phosphatase activity on the MC composite held at a constant low level. Col could have induced BMSC differentiation into osteoblasts that have higher alkaline phosphatase activity than BMSC (Fig. 6). The present study showed that BMSC differentiation into osteoblasts was improved on the PC composite surface when compared with the MC composite surface.
3.4. In vivo bone defect repair 3.4.1. Overview The effects of the composites on bone defect repair are presented in Fig. 7. No local infections occurred after composite implantation into the bone defect areas. At the end of the first month, massive new bone tissue formation with a capillary network was visible in the area treated with the MC composite. About a quarter of the bone defect area contained residual MC composite (arrow G). Meanwhile, large amounts of newly formed bone tissue filled the bone defect area that was treated with the PC composite. In the inferior border of the bone defect area, there was only a small amount of residual PC composite (arrow H) that established a weak connection with the surrounding tissues. A capillary network was seen on the surface of new bone tissues. At the end of the second month, in the area treated with the MC composite, most of the bone defect area was occupied by new bone tissue. In the border regions, only a little residual MC composite remained, connecting with the surrounding muscle tissues. Most of the bone defect area treated with the PC composite was filled with new bone tissue. A capillary network connecting with surrounding tissues was found in the cortical bone (arrow J). At the end of the fourth month, bone defect repair in the MC-implanted area was essentially completed, and a few capillaries were observed in the cortical bone with extensive connection to the surrounding tissues (arrow K). In the area treated with the PC composite, bone defect repair was also finished. There were a few capillaries with almost no connection to the surrounding tissue.
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Fig. 3. ESCA profiles of the PC and MC composites: (A) elemental peaks of the MC composite; (B) elemental peaks of the PC composite; (C) analysis of groups in C1s peak for the MC composite; (D) analysis of groups in C1s peak for the PC composite.
Table 1 Water absorption by PC and MC composites (g). Variable
Obs.
Mean
SE
SD
PC MC Combined Diff.
3 3 6
3.28 2.16 2.72 1.12
0.29 0.07 0.28 0.30
0.50 0.12 0.70
3.4.2. Masson staining The Masson staining results are shown in Fig. 8. In vivo implantation of the two bone repair composites induced differentiation of autologous BMSC and new bone formation in the bone defect areas. A large amount of type I collagen was secreted to aid in the repair of the bone defects. At the end of the first month, significant amounts of residual scaffold and disordered scaffold structures were observed. The structure of the MC composite was looser than that of the PC composite. Moreover, inflammatory infiltration and a large number of red blood cells were seen in the composite residual (arrow G). Inflammatory infiltration in the area treated with the PC composite was more pronounced than in that of the MC composite. Additionally, a large number of red blood cells were found in the PC-implanted area. By the end of the second month, the composites had been degraded into separate pieces. Multiple layers of the collagen deposits were found in the bone tissue. Layer-like bone lamellae and early bone trabeculae began to develop in certain regions. Blood vessels were noted in the areas surrounding bone trabeculae and
Fig. 4. Adhesion of BMSC on the PC and MC composites.
within the bone matrix (arrow H). More calcium deposits and bone trabeculae had formed in the area treated with the PC composite than in the MC composite area. In addition, new blood vessels in bone trabeculae and a large number of osteocytes in the bone matrix were found in the PC-treated area. At the end of the fourth month, both composites were fully degraded, and new bone trabeculae were formed (arrow J). An exten-
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PC-treated side in the central defect was still slightly higher than that of the MC-treated side at the end of the 4th and 12th weeks. The average ROI was 74.57 pixels in the PC-treated side and 61.64 in the MC-treated side at the end of the 4th week; 95.55 pixels in the PC-treated side and 93.68 pixels in the MC-treated side at the end of the 12th week.
4. Discussion
Fig. 5. BMSC growth curves on the PC and MC composites.
Fig. 6. BMSC alkaline phosphatase activity on the MC and PC composites at 405 nm.
sive capillary network existed between the bone trabeculae. Bone destruction and reconstruction occurred within the trabeculae. There were no marked differences in bone defect repair between the two composites. However, the extent of peripheral vascularization was higher in the PC-treated area. 3.4.3. Radionuclide bone scan Nuclide bone scans were used to measure 99mTc-MDP uptake at the bone defect sites (Fig. 9). Angiographic Phase I (0–30 s) images showed neovascularization around the bone defect sites. Bone-uptake Phase III (2 h) images showed reconstruction and new bone formation at 4, 8 and 12 weeks, which had a different appearance from the implant graft. In angiographic phase analysis (Tables 2–4 and Fig. 9), 99mTcMDP uptake was higher on the Col–CS–HA side than the Col–HA side after 4, 8 and 12 weeks (p = 0.00001 < 0.001). Radiographic evaluation in Phase III revealed that the radiopaque shadow of the PC-treated side in the central defect was higher than that of the MC-treated side. The average ROI in the PC-treated side was 113.64 pixels compared with 76.56 pixels in the MC-treated side at the end of the 8th week. Radiographic evaluation in Phase III also showed that the radiopaque shadow of the
An ideal biomaterial for bone defect repair can function as a scaffold to allow entry and attachment of host blood vessels and cells, facilitating formation of new bone and subsequent degradation and absorption of the implant. Therefore, the usefulness of a biomaterial for bone defect repair is largely determined by its capacity to repair bone defects in vivo. Bony tissues consist mainly of mineralized type I collagen fibrils. These have plate-like crystals of carbonated hydroxyapatite arranged in parallel layers along grooves or gap regions through the fibrils. HA exists chiefly in this form in natural bones [20]. During the biomineralization process, hydroxyapatite microcrystals formed between the ends of collagen fibers, using the polar groups on the collagen surface as crystal nuclei (Fig. 2). During in vitro biomineralization, SEM and TEM findings showed that HA did not form the parallel layers seen in natural bones [4,21]; instead, HA aggregated into crystal balls using polar groups as the core (Figs. 1 and 2). These observations are the difference between biomineralization and natural mineralization. Polar groups such as amino (–NH2), carboxyl (–COOH) and hydroxyl (–OH) groups are thereby blocked on the surface of the collagen fibers. Therefore, in mineralization crystallization, the existence of HA affects the biological properties of collagen [22]. In the present study, the spatial relationship between collagen and HA in Col–HA composites was altered by adjusting the crystalline phase. In the PC composite, HA microcrystals do not precipitate at the ends of the collagen fibers and, thus, HA does not affect fiber polymerization, and active groups existed on the collagen surface (Fig. 2). This strategy draws on the design of layer by layer [23], yet more streamlined in process. As seen by SEM and TEM, the Col–HA composites generated by mineralization crystallization or pre-crystallization exhibited different pore diameters, spatial arrangement and fiber thicknesses, owing to their different structures. In the PC composite, HA microcrystals are located within the collagen fiber scaffold; in the MC composite, however, HA microcrystals are exposed on the surface of the collagen scaffold. This finding is confirmed by ESCA results showing that HA microcrystals deposited on the collagen fiber surface in the MC composite contain N 1.92%, P 8.99% and Ca 17.56%. This suggests that HA, the source of Ca, is deposited on the surface of the collagen fibers, while the collagen fibers, the source of N, are covered. In comparison, when HA microcrystals are deposited within the collagen scaffold in the PC composite, the N content rises to 13.13%, and the P and Ca contents decrease to 4.39% and 5.86%, respectively. Furthermore, the PC composite exhibits a more regular scaffold structure, a smoother surface and more homogeneous pore size [24,25]. Compared with the MC composite, the PC composite surface has more polar groups (–NH2 and –COOH) derived mainly from the ends of the collagen fibers. A previous study showed that, during biological mineralization, these polar groups serve as nuclei in HA mineral crystallization [26]. Consequently, during fabrication of the Col–HA composites by mineralization crystallization, these polar groups are shielded by HA crystals (as shown by SEM). The number of polar groups on these composite surfaces is reduced, attenuating the hydrophilicity and the cell adhesion capacity of these bone repair biomaterials.
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Fig. 7. In vivo repair of rabbit mandibular defects: (A) 1 month after implantation of the MC composite; (B) 1 month after implantation of the PC composite; (C) 2 months after implantation of the MC composite; (D) 2 months after implantation of the PC composite; (E) 4 months after implantation of the MC composite; (F) 4 months after implantation of the PC composite.
In general, a hydrophobic surface absorbs fewer proteins than a hydrophilic surface, which may lead to denaturation of albumin. As seen in Table 1, the PC composite absorbs more water than the MC composite. This may be partly explained by the fact that Col, which is rich in the polar groups, is situated on the composite surface, thus increasing the number of proteins. Therefore, enhancing the hydrophilicity of bone repair materials increases their capacity for water absorption, promoting biocompatibility [27–29]. Cell culture serves as an important tool for studying material biocompatibility. As presented in Figs. 4–6, both composites provide a favorable microenvironment for the growth and functioning of BMSC after 3 weeks of culture. However, significant differences are seen with respect to cell shape and cell concentration. In particular, osteoblasts are distributed more uniformly and grow at a higher density on the PC composite. In addition, BMSC adhere better and excrete larger amounts of extracellular matrix on the PC composite. BMSC adhesion and growth are superior on the PC composite surface in comparison with the MC composite surface. The number of BMSC on the PC composite surface is similar to that of BMSC grown on the pure collagen scaffold. In addition, alkaline phosphatase activity is higher in the PC com-
posite than in the MC composite, suggesting that BMSC in the PC composite are in direct contact with the collagen fibers on the composite surface, accelerating their differentiation into osteoblasts induced by the active sites on the surface of the collagen fibers. Therefore, a large amount of extracellular matrix is formed. In general, both composites induce the differentiation of BMSC into osteoblasts and the formation of new bones that eventually replace artificial implants and repair bone defects. Nevertheless, differences between the two composites are observed by macroscopic observation and radionuclide bone scans. Macroscopic observation reveals that, during the first and second months after implantation, bone defect repair proceeds significantly faster in the PC composite. Furthermore, Masson staining at the end of the first month shows that both composites essentially maintain their morphological integrity, despite the initiation of decomposition. Disintegration and absorption together with creeping substitution occur in both composites. Substitution occurs mainly in the early stages during the formation of new vascular channels within the composites. By the end of the second month, most of the compos-
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Fig. 8. Masson staining of rabbit mandibular defects (A) 1 month after implantation of the MC composite; (B) 1 month after implantation of the PC composite; (C) 2 months after implantation of the MC composite; (D) 2 months after implantation of the PC composite; (E) 4 months after implantation of the MC composite; (F) 4 months after implantation of the PC composite (100).
ites decompose, leaving only scattered fragments. Substantial amounts of layered collagen precipitate in bone tissues where bone lamellae form, indicating that most of the bone repair is completed during the second month in both composites. This finding is consistent with the radionuclide bone scan results. At the end of the fourth month, the composites are completely decomposed; bone trabeculae formation and reconstruction are seen; osteocytes and bone lacuna are found in the bone trabeculae. The fact that the hydroxyapatite crystals grow with their c-axes oriented along the long axes of the nanofibers could be of interest in the design of materials. The present paper has described the nanostructure of Col–HA composite by self-assembly. HA assembled around the polar groups of collagen fiber in clusters. In pure collagen scaffold, collagen fibers can be cross-linked by the formation of intermolecular disulfide bonds upon oxidation, which results in a chemically robust fiber. However, in the scaffold derived from biomineralization, HA aggregation affected Col cross-linking, thereby affecting the overall structure of the scaffold. As shown in the present study, deficiencies in the biological performance of the scaffold prepared by biomineralization are mani-
fested in several aspects. First, loose association between fibers and poor mechanical performance led to irregular spatial structures. Second, polar groups on the surface of collagen fibers were blocked, resulting in lower hydrophilicity. Third, HA aggregation on Col polar groups blocked cell attachment sites such as RGD. As a result, the overall biological performance of HA–Col composites is affected. Comparison of such materials with pre-crystallized material also proves this finding. In contrast, in the pre-crystallization, HA crystallizes before Col precipitation and thus does not hide the polar groups on the surface of Col. Because of the different order of crystallization, the two composites demonstrate different biological properties.
5. Conclusion It can be concluded that HA placement on the collagen fibers affected the biological performance of Col–HA composites. Pre-crystallization made HA scatter among collagen fibers, creating a better structure for bone defect repair in comparison with MC Col–HA composites.
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Fig. 9. Radionuclide bone scans of rabbit mandibular defects following repair with PC and MC composites: (A–C) angiographic Phase I; (D–F) bone-uptake Phase III. (A and D) 1 month after implantation; (B and E) 2 months after implantation; (C and F) 4 months after implantation (left implant: the PC composite; right implant: the MC composite).
Table 2 The 99mTc-MDP taken up in angiographic phase analysis at 1 M.
Table 3 99m Tc-MDP taken up in angiographic phase analysis at 2 M.
Variable
Obs.
Mean
SE
SD
Variable
Obs.
Mean
SE
SD
PC MC Diff.
30 30
1610 1413 197
441 385
81 70
PC MC Diff.
30 30
2390 2063 327
469 455
86 83
L. Zhang et al. / Acta Biomaterialia 6 (2010) 2189–2199 Table 4 99m Tc-MDP taken up in angiographic phase analysis at 4 M. Variable
Obs.
Mean
SE
SD
PC MC Diff.
30 30
2942 2326 616
801 634
146 116
Acknowledgements This work was supported by grants from the National Natural Science Foundation of China (30700848 and 30772443) and the Natural Science Foundation of Beijing, China (7042058 and 7082090). Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figures 3, 7–9, are difficult to interpret in black and white. The full colour images can be found in the on-line version, at doi: 10.1016/j.actbio.2009.12.042. References [1] Keating JF, McQueen MM. Substitute for autologous bone graft in orthopaedic trauma. J Bone Joint Surgery 2001;83B:3–8. [2] Crane GM, Ishaug SL, Mikos AG. Bone tissue engineering. Nat Med 1995;1:1322–4. [3] Damien CJ, Parsons JR. Bone graft and bone graft substitutes: a review of current technology and applications. J Appl Biomater 1991;2:187–208. [4] Lees S. A model for the distribution of HAP crystallites in bone—an hypothesis. Calcif Tissue Int 1979;27:53–6. [5] Prostak KS, Lees S. Visualization of crystal-matrix structure. In situ demineralization of mineralized Turkey leg tendon and bone. Calcif Tissue Int 1996;59:474–9. [6] Hartgerink Jeffrey D, Beniash Elia, Stupp Samuel I. Self-assembly and mineralization of peptide-amphiphile nanofibers. Science 2001;23:1684–8. [7] Rammelt S, Schulze E, Witt M, Petsch E, Biewener A, Pompe W, et al. Collagen type I increases bone remodelling around hydroxyapatite implants in the rat tibia. Cells Tissues Organs 2004;178:146–57. [8] Du C, Falini G, Fermani S, Abbott C, Moradian-Oldak J. Supramolecular assembly of amelogenin nanospheres into birefringent microribbons. Science 2005;307:1419–20. [9] Taton TA. Nanotechnology. Boning up on biology. Nature 2001;412(6846): 491–2.
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[10] Kanungo BP, Gibson LJ. Density–property relationships in mineralized collagen–glycosaminoglycan scaffolds. Acta Biomater 2009;5(4):1006–18 [Epub 2008 Dec 11]. [11] He G, Dahl T, Veis A, George A. Nucleation of apatite crystals in vitro by selfassembled dentin matrix protein 1. Nat Mater 2003;2(8):552–8. [12] Hunter GK, Hauschka PV, Poole AR, Rosenberg LC, Goldberg HA. Nucleation and inhibition of hydroxyapatite formation by mineralized tissue proteins. Biochem J 1996;317:59–64. [13] Boskey AL. Osteopontin and related phosphorylated sialoproteins: effects on mineralization. Ann NY Acad Sci 1995;760:249–56. [14] Weiner S, Traub W. Bone structure: from angstroms to microns. FASEB J 1992;6(3):879–85 [Review]. [15] Butler WT, Ritchie H. The nature and functional significance of dentin extracellular matrix proteins. Int J Dev Biol 1995;39:169–79. [16] Zhang L, Hu M, Wang M, Li Y, Chen H, Chu C, Jiang H. Evaluation of modifying collagen matrix with RGD peptide through periodates oxidation. J Biomed Mater Res 2005;73A:468–75. [17] Martin I, Shastri VP, Padera RF, Yang J, Mackay AJ, Langer R, et al. Selective differentiation of mammalian bone marrow stromal cells cultured on threedimensional polymer foams. J Biomed Mater Res 2001;55:229–35. [18] Mills BG, Singer FR, Weiner LP. Long-term culture of cells from bone affected by Paget’s disease. Calcif Tissue Int 1979;29:79–87. [19] Ishaug SL, Crane GM, Miller MJ, Yasko AW, Yaszemski MJ, Mikos AG. Bone formation by three-dimensional stromal osteoblast culture in biodegradable polymer scaffolds. J Biomed Mater Res 1997;36:17–28. [20] Robinson RA, Watson ML. Collagen–crystal relationships in bone as seen in the electron microscope. Anat Rec 1952;114:383–410. [21] Hellmich C, Ulm FJ. Average hydroxyapatite concentration is uniform in the extracollagenous ultrastructure of mineralized tissues: evidence at the 1–10microm scale. Biomech Model Mechanobiol 2003;2(1):21–36. [22] Blair HC, Zaidi M, Schlesinger PH. Mechanisms balancing skeletal matrix synthesis and degradation. Biochem J 2002;364(Pt 2):329–41. [23] Ficai A, Andronescu E, Voicu G, Manzu D, Ficai M. Layer by layer deposition of hydroxyapatite onto the collagen matrix. Mater Sci Eng C 2009;29: 2217–20. [24] Pieper JS, Oosterhof A. Preparation and characterization of porous crosslinked collagenous matrices containing bioavailable chondroitin sulphate. Biomaterials 1999;20:847–58. [25] Madihally Sundararajan V, Howard WT. Porous chitosan scaffolds for tissue engineering. Biomaterials 1999;20:1133–43. [26] Gajjeraman S, Narayanan K, Hao J, Qin C, George A. Matrix macromolecules in hard tissues control the nucleation and hierarchical assembly of hydroxyapatite. J Biol Chem 2007;282:1193–204. [27] Stephen W, Wolfie T. Bone structure: from angstroms to microns. FASEB J 1992;6:879–85. [28] Baht GS, Hunter GK, Goldberg HA. Bone sialoprotein–collagen interaction promotes hydroxyapatite nucleation. Matrix Biol 2008;27:600–8. [29] Brandao-Burch A, Utting JC, Orriss IR, Arnett TR. Acidosis inhibits bone formation by osteoblasts in vitro by preventing mineralization. Calcif Tissue Int 2005;77:167–74.