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Using calcium silicate to regulate the physicochemical and biological properties when using β-tricalcium phosphate as bone cement
Materials Science and Engineering C 43 (2014) 126–134
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Using calcium silicate to regulate the physicochemical and biological properties when using β-tricalcium phosphate as bone cement Chia-Tze Kao a,b, Tsui-Hsien Huang a,b, Yi-Jyun Chen a,b,c, Chi-Jr Hung a,b, Chi-Chang Lin d,⁎,1, Ming-You Shie d,⁎⁎,1 a
School of Dentistry, Chung Shan Medical University, Taichung, Taiwan Department of Dentistry, Chung Shan Medical University Hospital, Taichung, Taiwan Dental Department, Taichung Hospital, Ministry of Health and Welfare, Taichung City, Taiwan d Department of Chemical and Materials Engineering, Tunghai University, Taichung, Taiwan b c
a r t i c l e
i n f o
Article history: Received 1 December 2013 Received in revised form 19 May 2014 Accepted 30 June 2014 Available online 6 July 2014 Keywords: β-Tricalcium phosphate Calcium silicate Biocomposites Osteogenic Angiogenic Antibacterial
a b s t r a c t β-Tricalcium phosphate (β-TCP) is an osteoconductive material. For this research we have combined it with a low degradation calcium silicate (CS) to enhance its bioactive and osteostimulative properties. To check its effectiveness, a series of β-TCP/CS composites with different ratios were prepared to make new bioactive and biodegradable biocomposites for bone repair. Regarding the formation of bone-like apatite, the diametral tensile strength as well as the ion release and weight loss of composites were compared both before and after immersions in simulated body fluid (SBF). In addition, we also examined the behavior of human dental pulp cells (hDPCs) cultured on β-TCP/CS composites. The results show that the apatite deposition ability of the β-TCP/CS composites improves as the CS content is increased. For composites with more than a 60% CS content, the samples become completely covered by a dense bone-like apatite layer. At the end of the immersion period, weight losses of 24%, 32%, 34%, 38%, 41%, and 45% were observed for the composites containing 0%, 20%, 40%, 80%, 80% and 100% β-TCP cements, respectively. In addition, the antibacterial activity of CS/β-TCP composite improves as the CS-content is increased. In vitro cell experiments show that the CS-rich composites promote human dental pulp cell (hDPC) proliferation and differentiation. However, when the CS quantity in the composite is less than 60%, the quantity of cells and osteogenesis protein of hDPCs is stimulated by Si released from the β-TCP/CS composites. The degradation of β-TCP and the osteogenesis of CS give strong reason to believe that these calcium-based composite cements will prove to be effective bone repair materials. © 2014 Elsevier B.V. All rights reserved.
1. Introduction Materials with varying degrees of bioactivity and degradation are required so as to conform to differing clinical requirements for hard tissue repair. Autograft possesses all the characteristics indispensable for new bone formation, namely, osteogenesis, osteoconductivity, and osteoinductivity, and it is currently considered the gold standard in this field. In spite of its strengths, however, there are still various disadvantages when using this material for certain medical clinical applications. Therefore, the design of an even better composite material would be very beneficial for controlling material physicochemical properties. Silicate-based ceramics have received a considerable amount of positive attention in recent years as these materials have better
⁎ Corresponding author. Tel.: +886 4 23590262x210. ⁎⁎ Corresponding author. Tel.: +886 4 24718668x55511; fax: +886 4 24759065. E-mail addresses: [email protected] (C.-C. Lin), [email protected] (M.-Y. Shie). 1 Both authors contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2014.06.030 0928-4931/© 2014 Elsevier B.V. All rights reserved.
bioactivity than calcium phosphate-based materials [1,2]. Recently, several studies have indicated that calcium silicate (CS) can play an important role in bone formation, at least based upon these materials' patterns of Si ion release [3] and its fast apatite formation ability [4]. CS can promote the proliferation of human mesenchymal stem cells (hMSCs), human dental pulp cells (hDPCs), and osteoblast-like cell adhesion, proliferation, and differentiation [5–9]. A previous study demonstrates that CS cement is biocompatible with cultured lipopolysaccharide-treated dental pulp cells and inhibits inflammation in lipopolysaccharide-treated dental pulp cells, leading to higher IL-1β expression and apoptosis [10]. Furthermore, this CS cement is able to activate angiogenesis [11] and anti-osteoclastogenesis [12]. These studies demonstrate that CS has the ability to promote the apatite layer's formation in vitro [13] and also show that CS is better at promoting bone regeneration in vivo [14,15]. However, the slow degradation rate of CS may result in a decrease in osteoconductivity, which may limit its clinical application [13]. In order to ameliorate its relative disadvantage in regard to material degradation, we used β-tricalcium phosphate (β-TCP) as an additive to see how it would affect CS's rate of decay. βTCP is a bioceramic material that is widely used for hard tissue repair.
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It has a chemical composition similar to apatite present in bone tissue, and has been applied extensively as a bone grafting material [16–18]. Previous studies have shown that wollastonite-doped TCP bioceramics have a higher degradation rate and promote new bone formation better than TCP in vivo [13,14]. Wang et al. assert that composite scaffolds containing 50% and 80% CS not only have good osteoconductivity, but also promote rapid bone formation compared with scaffolding comprised of pure β-TCP and CS [19]. Thus, to obtain both osteostimulation and osteoconductivity by taking advantage of the favorable bioactivity of calcium silicate and the high degradability of β-TCP, β-TCP/CS composite materials have been produced in the hopes that the right mixture can both help to control the degradation rate and improve interactions of the material with human tissue. In this study, β-TCP/CS composite cements with varied ratios were prepared so that we could observe the changes in physiochemical properties, bioactivity, in vitro degradation behavior, osteogenesis, angiogenesis and anti-bacterial activity in composites with different β-TCP/CS ratios. It is our hope that this knowledge may prove useful in the design of optimal biomaterials for bone regeneration. 2. Materials and methods 2.1. Preparation of β-TCP/CS composites The method used here for the preparation of CS powder has been described elsewhere [12]. In brief, reagent grade SiO2 (High Pure Chemicals, Saitama, Japan), CaO (Riedel-de Haen, Steinheim, Germany), Al2O3 (Sigma-Aldrich, St. Louis, MO) and ZnO (Wako, Osaka, Japan) powders were used as matrix materials (composition: 65% CaO, 25% SiO2, 5% Al2O3, and 5% ZnO) and mixed at 1500 rpm for 15 min using a hybrid defoaming mixer (ARE-250, Thinky, Tokyo, Japan). The oxide mixtures were then sintered at 1400 °C for 2 h using a high-temperature furnace and then ball-milled in ethyl alcohol using a centrifugal ball mill (S 100, Retsch, Haan, Germany) for 6 h. The β-TCP/CS composite material was obtained by mixing β-TCP (Sigma-Aldrich) and CS powder with composite weight ratios of 80:20 (C8T2), 60:40 (C6T4), 40:60 (C4T6), and 20:80 (C2T8) wt.% at 1500 rpm for 15 min using a hybriddefoaming mixer. The code of the pure CS and β-TCP was C10T0 and C0T10, respectively. The β-TCP/CS powder was mixed using a liquid/ powder ratio of 0.35 mL/g. After mixing with liquids (1 M Na2HPO4), the cements were molded in a Teflon mold (diameter: 6 mm, height: 3 mm). The cement quantities were enough to fully cover each well of the 24-well plate (GeneDireX, Las Vegas, NV) to a thickness of 2 mm for cell experiments. All samples were stored in an incubator at 100% relative humidity and 37 °C for 1 day of hydration.
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(ISO) 9917-1. The setting time was recorded when the Gilmore needle failed to create a 1-mm deep indentation in three separate areas. 2.4. In vitro soaking To evaluate the in vitro bioactivity, the cements were immersed in a 10 mL simulated body fluid (SBF) solution at 37 °C. The SBF solution, of which the ionic composition is similar to that of human blood plasma, consisted of 7.9949 g of NaCl, 0.3528 g of NaHCO3, 0.2235 g of KCl, 0.147 g of K2HPO4, 0.305 g of MgCl2·6H2O, 0.2775 g of CaCl2, and 0.071 g of Na2SO4 in 1000 mL of distilled H2O and was buffered to a pH of 7.4 with hydrochloric acid (HCl) and trishydroxymethyl aminomethane (Tris, CH2OH)3CNH2) [20,21]. All chemicals used were of reagent grade. The solution in the shaker water bath exhibited no change under static conditions. After soaking for different time durations (3 days to 3 months), specimens were removed from the tube and evaluated for several physicochemical properties. 2.5. Mechanical properties The composite specimens without/with immersion in SBF were dried at 60 °C for 3 days. Next, diametral tensile strength (DTS) testing was conducted on an EZ-Test machine (Shimadzu, Kyoto, Japan) at a loading rate of 1 mm/min. The maximal compression load at failure was obtained from the recorded load–deflection curves. At least 10 specimens from each group were tested. 2.6. Weight loss The degree of degradation was determined by monitoring the weight change of the specimens. After drying at 60 °C, the composites were weighed using a balance both before and after immersion. The ten specimens were examined at each time point for each of the materials being investigated. 2.7. Ion concentration The Ca, Si, P, Zn, and Al ion concentrations released from the cement on the SBF were determined using an inductively coupled plasmaatomic emission spectrometer (ICP-AES; Perkin-Elmer OPT 1MA 3000DV, Shelton, CT, USA) after the samples had been immersed for various periods of time. Three samples were measured for each data point and the results were obtained in triplicate from three separate samples for each test.
2.2. Phase composition and morphology
2.8. Dental pulp cell isolation and culture
The phase composition of the cements was analyzed using X-ray diffractometry (XRD; Bruker D8 SSS, Karlsruhe, Germany), and operated at 30 kV and 30 mA at a scanning speed of 1°/min. The morphology of the cement specimens was coated with gold and examined under a scanning electron microscope (SEM; JSM-6700F, JEOL) operated in the lower secondary electron image (LEI) mode at 3 kV accelerating voltage. Scanning electron microscopy associated with energy dispersive spectroscopy (EDS) was used to characterize the chemical analyses of the specimens.
The human dental pulp cells (hDPCs) were freshly derived from caries-free, intact premolars that had been extracted for orthodontic treatment purposes, as described previously [11]. The patient gave informed consent, and approval from the Ethics Committee of the Chung Shan Medicine University Hospital was obtained (CSMUH No. CS11187). A sagittal split was performed on each tooth using a chisel, and the pulp tissue was immersed in a PBS buffer solution. The pulp tissue was then cut into fragments, distributed onto plates and cultured in DMEM. These cultures were supplemented with 20% fetal bovine serum (FBS; Caisson), 1% penicillin (10,000 U/mL)/streptomycin (10,000 mg/mL) (PS, Caisson) and kept in a humidified atmosphere with 5% CO2 at 37 °C; the medium was changed every 3 days. The osteogenic differentiation medium was DMEM supplemented with 10−8 M dexamethasone (Sigma-Aldrich), 0.05 g/L L-ascorbic acid (Sigma-Aldrich) and 2.16 g/L glycerol 2-phosphate disodium salt hydrate (Sigma-Aldrich).
2.3. Setting time After the powder was mixed with H2O, the composites were placed into a cylindrical mold and stored in an incubator at 37 °C and 100% relative humidity for hydration. The setting time of the cements was tested according to standards set by the International Standards Organization
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2.9. Cell viability Cell suspensions at a density of 104 cells/mL were directly seeded over each specimen for various numbers of days. Cell cultures were incubated at 37 °C in a 5% CO2 atmosphere. After different culturing times, cell viability was evaluated using the PrestoBlue® (Invitrogen, Grand Island, NY) assay. At the end of the culture period, the medium was discarded and the wells were washed with cold PBS twice. Each well was then filled with a medium with a 1:9 ratio of PrestoBlue® in fresh DMEM and incubated at 37 °C for 30 min after which the solution in each well was transferred to a new 96-well plate. Plates were read in a multiwell spectrophotometer (Hitachi, Tokyo, Japan) at 570 nm with a reference wavelength of 600 nm. Cells cultured on a tissue culture plate without the cement were used as a control (Ctl). The results were obtained in triplicate from three separate experiments in terms of optical density (OD).
transferred to a new 96-well plate. Plates were then read in a multiwell spectrophotometer (Hitachi, Tokyo, Japan) at 570 nm with a reference wavelength of 600 nm. Cells cultured on the tissue culture plate without the cement were used as a control (Ctl). The results were obtained in triplicate from three separate experiments in terms of optical density (OD).
2.14. Statistical analysis A one-way analysis of the variance statistical data was used to evaluate the significance of the differences between the means in the measured data. Scheffe's multiple comparison test was used to determine the significance of the deviations in the data for each specimen. In all cases, the results were considered statistically significant with p value b 0.05.
2.10. Osteogenesis assay 3. Results and discussion The level of ALP activity was determined after the cells had been seeded for 3 and 7 days. The process was as follows: the cells were lysed from disks using 0.2% NP-40 and centrifuged for 10 min at 2000 rpm after washing with PBS. ALP activity was determined using p-nitrophenyl phosphate (pNPP, Sigma) as the substrate. Each sample was mixed with pNPP in 1 M diethanolamine buffer for 15 min, after which the reaction was stopped by the addition of 5 N NaOH and quantified by absorbance at 405 nm. All experiments were done in triplicate. The OC protein released from the pulp cells was cultured on different substrates for 7 and 14 days after cell seeding. Following the manufacturer's instructions, an osteocalcin enzyme-linked immunosorbent assay kit (Invitrogen) was used to determine the OC protein content. The OC protein concentration was measured by correlation with a standard curve. The analyzed blank disks were treated as controls. All experiments were done in triplicate. 2.11. Alizarin Red S stain Accumulated calcium deposition was observed after 7 and 14 days using Alizarin Red S staining as described in a previous study [3]. To summarize briefly, the cells were fixed with 4% paraformaldehyde (Sigma-Aldrich) for 15 min and then incubated in 0.5% Alizarin Red S (Sigma-Aldrich) at pH 4.0 for 15 min at room temperature in an orbital shaker (25 rpm). To quantify the stained calcified nodules after staining, samples were immersed with 1.5 mL of 5% SDS in 0.5 N HCl for 30 min at room temperature, following which the tubes were centrifuged at 5000 rpm for 10 min after which the supernatant was transferred to the new 96-well plate (GeneDireX). At this time, absorbance was measured at 405 nm (Hitachi). In addition, we also quantify the stained specimens without culturing cells (0 day).
3.1. Characterization of β-TCP/CS biocomposites Fig. 1 shows the XRD patterns of the composite powder after hydration for 1 day. The results indicate both that the composite samples consist of β-TCP and CS and that no chemical reaction occurred between the CS and TCP. The analysis of the specimens containing CS reveals two key points: first, an obvious diffraction peak near 2θ = 29.4°, which corresponds to the calcium silicate hydrate (CSH) gel, and second, incompletely reacted inorganic component phases of the βdicalcium silicate (β-Ca2SiO4) at 2θ between 32° and 34° [8]. It is clear that the addition of CS results in lower peak intensities of the CSH and β-Ca2SiO4 phases. Previously, Ni et al. performed a detailed study of the phase diagram of the system Ca2SiO4/Ca3(PO4)2 and their results indicate that the system is a true binary [22]. Our results are similar. Fig. 2 shows that with an increase in the amount of CS the setting time of the cement becomes shorter, going from 37 min (β-TCP group) all the way down to 16 min (CS group), a significant difference (p b 0.05). Under clinical conditions the setting time is a critical factor and a long setting period can lead to clinical problems. Fernández et al. suggest that 10–15 min is a suitable setting time interval [23]. In the present study, the setting time of calcium-silicate based cements was proportional to the CSH amount [7]. CSH is able to reduce the setting time of the calcium-based cement. Our results show that a setting
2.12. Intracellular Ang-1 and vWF measurement The production of Ang-1 and vWF was quantified using ELISA kits (Abcam, catalog nos. ab99970 and ab108918) according to the manufacturer's instructions. The hDPCs were cultured on substrates for 3 and 7 days, and proteins from whole cell lysates were collected and quantified using the ELISA kit. 2.13. Antibacterial property To investigate the anti-bacterial effect of the β-TCP/CS composite, the hydration composites were mixed with 1 mL Staphylococcus aureus in LB culture media (4.0 × 104 bacteria per mL) and cultured for 3 and 12 h. Aliquots of 0.1 mL from each group were then mixed with 0.9 mL PrestoBlue® for 30 min after which the solution in each well was
Fig. 1. XRD patterns of β-TCP/CS composites have different ratios after hydration.
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Dissolubility is a key factor in biodegradation [22] and must be considered when developing a material which has a degradation rate most appropriate to hastening and easing the process of tissue repair. With this in mind, the degradation rates of the β-TCP/CS composites in SBF solution have been recorded for various periods of time ranging from 3 to 84 days, as recorded in Fig. 3. After soaking for 3 days, the CS cement shows a relatively modest weight loss (~ 11%), whereas the β-TCP-containing cement lost considerably more weight (13–18%). All the specimens display an increased weight loss as the immersion time is increased. The pure CS cement (T0C10) has the lowest dissolution rate and solubility compared with other samples over the whole soaking period, reaching 24% after 12 weeks. By contrast, the overall amount of β-TCP dissolution reached 45% after 12 weeks. At the end of the immersion period, weight losses of approximately 32%, 34%, 38%, and 41%, were observed for the C2T8, C4T6, C6T4, and C8T2 cement
mixtures, respectively, indicating significant differences (p b 0.05). As expected, the β-TCP/CS biphasic composites show an intermediate dissolution behavior between those of pure β-TCP and CS. Moreover, as the CS content is increased, the dissolution rate also decreases steadily. Hence, the degradation rate of the composite cement may be controlled to a certain extent by varying the CS content in the composite [22]. Changes in the strength values of cement samples after immersion in SBF are shown in Fig. 4. The DTS values of hydration cements that are not immersed range from 2.1–3.5 MPa, indicating a significant (p b 0.05) decrease in the strength as the amount of β-TCP is increased. The decrease in the strength of the composite cement is probably due to the addition of the inherently weak β-TCP to CS. In addition, there is no reaction or chemical bonding between the CS and β-TCP in the cement; the two substances only stay together because they are mixed in the composite cements; they do not otherwise interact. After immersion in SBF, the mechanical properties change into biodegradable biomaterials. These results demonstrate that the CS-containing cement increases the DTS after 4 weeks of immersion, and thereafter decreases. When cement specimens are immersed in solution for 4 weeks, the CS hydration reaction dramatically changes in the CSH phase and increases in strength [13,26]. After 12 weeks of soaking, the strengths of C0T10, C2T8, and C4T6 are 1.4, 1.4, 2.2 MPa, respectively. This indicates that the strength is lower than cements that are not soaked. In contrast, the strength of the higher CS content composites (C6T4, C8T2 and C10T0) is similar to as-hardened cements. Moreover, the DTS of the higher β-TCP content cements declines due to the degradation, consistent with the results of weight loss. The results of an examination of the surface microstructure of the specimens before and after immersion in SBF after periods of 7 days are shown in Fig. 5. It is readily visible that the β-TCP cement has a looser and rougher surface texture, with irregular pores (Fig. 5F). In contrast, the pure CS cement exhibits a dense and smooth surface containing particle entanglement and micro-pores (Fig. 5A). In a bone graft, it is believed that when bonded to living bone, an apatite layer will form on the surface [4]. The formation of the bone-like apatite in SBF has proven to be useful in predicting the bone-bonding ability of material in vitro. It can be seen that the surfaces of specimens with 70% and 100% CS content are covered by a dense apatite layer after being immersed for 7 days. After immersion, the surface of the two samples is completely covered by an apatite layer with tiny spherical shapes on the surface. However, it is clear that no precipitate has formed on
Fig. 3. Weight loss of various cements after immersion in SBF for predetermined time durations. “*” indicates a significant difference (p b 0.05) compared to C10T0.
Fig. 4. Diametral tensile strength of various cements after immersion in SBF for predetermined time durations.
Fig. 2. The setting time of various β-TCP/CS composites hydrated. Statistical data are presented as means ± standard deviation for n = 6. Values not sharing a common letter are significantly different at p b 0.05.
time of approximately 20 min is required for injectable bone cements for clinical use [7,24,25]. 3.2. Immersion studies of β-TCP/CS biocomposites
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Fig. 5. SEM micrographs of the β-TCP/CS composites surfaces before and after immersion in SBF for 7 days. The white arrow indicated apatite sphere.
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the surfaces of the β-TCP-rich cement specimens (T10C0 and T7C3). For the T5C5 composite, after 7 days of immersion spherical granules that have precipitated on the surface of the composite become evident, and the morphology reveals an early stage of apatite precipitation. In addition, the Ca/P ratios of the C10T0 specimen after 1 and 4 weeks of immersion were 2.13 and 1.68, respectively. This precipitate has been identified to be apatite, although a Ca/P ratio of 1.68, which was appropriated Ca/P ratio (the stoichiometric Ca/P ratio, 1.67, of apatite) on 4week immersion surfaces. Similar apatite spherules have been observed
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to precipitate on CS materials surface after immersion in SBF [25]. These results are similar to reports by other researchers [14,22], who elucidate the lack of apatite formation observed in the β-TCP samples after soaking in SBF for several days. The in vitro bioactivity of the calciumions in the silicate based materials indicates that the presence of PO3− 4 composition is not an essential requirement for the formation of an apdepletes atite layer. This is noteworthy because it is known that PO3− 4 ions originate from the calcium and phosphate ions because the PO3− 4 in vitro assay solution [7]. Most importantly, the Si-OH functional
Fig. 6. (A) Ca, (B) Si, and (C) P ion concentrations of SBF after immersion for different times. Data represent means ± SD (n = 6). “*”, statistically significant difference from C10T0.
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group on CS materials has been proven to act as the nucleation center for apatite formation [4]. This may explain why the β-TCP cement had no bioactivity. Thus, the CS-rich cements will develop a stronger bond with the surrounding bone tissue compared with the β-TCP-rich cement. Variations of SBF Ca, Si, P, Zn and Al ion concentrations after being cultured for different time periods are shown in Fig. 6. After 12 weeks, the Ca ion concentration in the CS cement medium has decreased to approximately 1.2 mM, which is lower than the baseline Ca concentration of SBF (2.5 mM) (p b 0.05). As for the P ion concentration of SBF, it decreases after 1 day of immersion. The bioactivity of silicate-based materials indicates that the presence of PO43− ions in the composition is not an essential requirement for the development of an apatite layer, which consumes calcium and phosphate ions [4,24,27]. Si concentrations increase with increasing the duration of incubation. Si ion concentrations are about 1.5, 2.0, 2.3, 2.7, and 3.4 mM at C2T8, C4T6, C6T4, C8T2, and C10T0 after 12 weeks, respectively. However, the Zn and Al ions released from specimens were stable after immersion for 4 weeks. Calcium silicate cement undergoes hydrolysis on immersion, and Si ion dissolves during incubation [7,28]. Taking cell functions into account, the certain amounts of Si released from silicate-based materials may promote desired cell behaviors [29–31]. The current CS gives the ability to control the rate at which soluble Si ions affect cell adhesion and proliferation [7,32].
3.3. hDPC proliferation Cell viability and functions associated with a bone graft are closely related to the physical, chemical, and biological characteristics of the materials used. To consider the effects of CS on osteogenic activity, the biological functions of hDPCs cultured on various cement specimens have been evaluated after different periods of time (Fig. 7). Cell proliferation gradually increases as the amount of CS added to β-TCP is increased, which is notably different (p b 0.05) from the behavior of β-TCP cement. For example, on day 14 the C4T6 cement elucidated an increase of approximately 22% in the OD value compared to the pure β-TCP cement. Similarly, the β-TCP cement elicited a significant decrease of 12%, 17%, 30%, and 24% in comparison with the CS cement on days 1, 3, 7 and 14 of cell seeding, respectively.
Fig. 7. Proliferation of hDPCs treated with various specimens for 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
Fig. 8. (A) ALP activity of hDPCs cultured on various specimens for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
3.4. Osteogenesis protein secretion Osteogenesis differentiation of osteoblast-like cells is one of the key steps in determining whether bone formation is successful. The ALP expression of hDPCs cultured on different composites has also been examined, and Fig. 8 shows the analysis of quantitative examination data and the ALP activity of cells cultured on the different composites after 3 and 7 days. Increasing the CS content led to increased ALP levels in the cements at all incubation times. A significant 29% and 35% increase (p b 0.05) in the ALP level was measured for C4T6 and C6T4 in comparison with the β-TCP cement after 14 days. ALP enzyme activity is also associated with bone formation, and it is produced in high levels during the bone formation phase [7]. Similarly, the OC secretion by cells cultured on T0C10 is significantly higher (p b 0.05) than the levels secreted by cells on other substances (Fig. 9). After a 14 day period, the OC secretion was found to have been increased by 14% and 26% in the cells cultured on C4T6 and C6T4 in comparison with the β-TCP cement. Recent studies also show that CS promotes hDPC proliferation and differentiation [5,7]. This stimulatory effect may be attributed to the
Fig. 9. (A) OC amount of hDPCs cultured on various specimens for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
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Fig. 10. Quantification of calcium mineral deposits by Alizarin Red S assay of hDPCs cultured on various cement for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
dissolution of silicate ions [3,8]. It is worth noting that the β-TCP-rich cements with FGF-2 stimulate significantly more (p b 0.05) osteogenesis protein secretion than the pure cements. 3.5. Mineralization The aim of this assay is to determine and show the effects of various ratios of β-TCP/CS cement on bone matrix formation following analysis using Alizarin Red S staining to identify calcium deposition, as seen in Fig. 10. First, the results of quantification of all specimens without cell (0 day) were similar between all specimens (p N 0.05). Therefore, we consider that the method is suitable in this study. In the case of the C0T10 and C2T8 composites, no significant differences (p N 0.05) in the quantities of the calcium mineral matrix deposition were detected after either 7 or 14 days. By contrast, on day 14 a significant (1.18-, 1.32-, and 1.37-fold on C4T6, C6T4, and C8T2, respectively) enhancement (p b 0.05) of calcium content has been observed compared with β-TCP cement. According to the literature, tricalcium phosphate cement always fails to form a chemical bond with bone tissue during the early stages of therapy because of its poor bioactivity [22]. Taken together, our result indicates that the composite cement may possess a greater ability to interact with the hDPCs and to promote osteogenesis and matrix mineralization than the β-TCP cement. Such a system would combine the benefits of composite properties, such as resorbability (the ability to be replaced by native bone tissue), and osteoconductivity (the ability to induce bone growth). 3.6. Angiogenesis The protein expression levels of Ang-1 and vWF in hDPCs cultured on various specimens were evaluated on days 3 and 7 (Fig. 11). The Ang-1 expression in C4T6, C6T4, and C8T2 on day 7 was enhanced 1.54, 1.78, and 1.99 times, respectively, as compared to that of C0T10 (Fig. 11A). The vWF results shown in Fig. 11B can be seen to be similar to Ang-1, both showing a dose-dependent up-regulation in the specimens in which the CS levels have been increased (p b 0.05). vWF is an important protein involved in coagulation and thrombus formation. Following synthesis, it can be found in secretary granules called Weibel–Palade bodies and in vessels, and is released both constitutively and in a regulated manner [33,34]. Ang-1 is another family of growth
Fig. 11. The protein expression of (A) Ang-1 and (B) vWF on hDPCs cultured on β-TCP/CS cement for different days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
factors that plays an important role in vascular development [35,36]. Our data shows that cement with a higher CS content has a higher potential to induce vWF and Ang-1 expression than the β-TCP cement during angiogenesis. In addition, Si ions have been shown to increase the osmolality required to facilitate the angiogenic differentiation of hDPCs via the MAPK/p38 pathway [11]. 3.7. Antibacterial property Besides the osteostimulation and angiogenesis properties of the CS/βTCP composite, these biocomposites have been shown to have significant
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Our results suggest that the incorporation of CS into β-TCP is a useful approach for obtaining composites with improved properties in regards to setting time, degradation, osteogenesis, angiogenesis, and antibacterial activity. Thus, we suggest 60 wt.% CS-containing β-TCP cement may be the best choice for hard tissue repair applications. Acknowledgments The authors acknowledge receipt of a grant from the National Science Council Taiwan grants (NSC 101-2314-B-040-011-MY3). The authors declare that they have no conflicts of interest. References
Fig. 12. The anti-bacterial effect of β-TCP/CS cement for 3 and 12 h. “*” indicates a significant difference (p b 0.05) compared to Ctl. “#” indicates a significant difference (p b 0.05) compared to Ca(OH)2.
antibacterial effects, as seen in Fig. 12. The Staphylococcus aureus cultures with Ctl and C0T10 show the highest absorbance at both the times observed (3 and 12 h) (p b 0.05). However, the antibacterial activity of CS/β-TCP composite increases in proportion with an increase in the CS-content. No significant difference was detected in the absorbance readings for C6T4, C8T2 and C10T0 groups compared to the Ca(OH)2 groups (p N 0.05). There may be two potential reasons for the antibacterial effect of the CS/β-TCP composite. One is that CS can release Ca and Si ions, leading to a weak alkaline microenvironment with a higher pH value which may restrain bacterial growth [37]; the other is that the CS particles themselves may permeate the bacteria [38] with possible antibacterial consequences. Whatever the cause, the CS/β-TCP composite may prove to be an excellent antibacterial agent. 4. Conclusions In this study, degradable and highly bioactive calcium-silicate based composite cement containing CS and β-TCP in different ratios was prepared and analyzed. Our research reveals that the mechanical properties of biocomposites increase steadily with an increase in CS portion in the ratio of various biocomposites. The dissolution rate of β-TCP/CS is strongly dependent on the β-TCP content, and β-TCP/CS composites with a lower CS content show higher dissolution rates. When the concentration of CS in the composite increases to 60%, the ability to form bone-like apatite is about the same as that for the pure CS. The results obtained in this study may be useful for designing calcium-based biocomposites with optimal biological and degradation properties.
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Using calcium silicate to regulate the physicochemical and biological properties when using β-tricalcium phosphate as bone cement Chia-Tze Kao a,b, Tsui-Hsien Huang a,b, Yi-Jyun Chen a,b,c, Chi-Jr Hung a,b, Chi-Chang Lin d,⁎,1, Ming-You Shie d,⁎⁎,1 a
School of Dentistry, Chung Shan Medical University, Taichung, Taiwan Department of Dentistry, Chung Shan Medical University Hospital, Taichung, Taiwan Dental Department, Taichung Hospital, Ministry of Health and Welfare, Taichung City, Taiwan d Department of Chemical and Materials Engineering, Tunghai University, Taichung, Taiwan b c
a r t i c l e
i n f o
Article history: Received 1 December 2013 Received in revised form 19 May 2014 Accepted 30 June 2014 Available online 6 July 2014 Keywords: β-Tricalcium phosphate Calcium silicate Biocomposites Osteogenic Angiogenic Antibacterial
a b s t r a c t β-Tricalcium phosphate (β-TCP) is an osteoconductive material. For this research we have combined it with a low degradation calcium silicate (CS) to enhance its bioactive and osteostimulative properties. To check its effectiveness, a series of β-TCP/CS composites with different ratios were prepared to make new bioactive and biodegradable biocomposites for bone repair. Regarding the formation of bone-like apatite, the diametral tensile strength as well as the ion release and weight loss of composites were compared both before and after immersions in simulated body fluid (SBF). In addition, we also examined the behavior of human dental pulp cells (hDPCs) cultured on β-TCP/CS composites. The results show that the apatite deposition ability of the β-TCP/CS composites improves as the CS content is increased. For composites with more than a 60% CS content, the samples become completely covered by a dense bone-like apatite layer. At the end of the immersion period, weight losses of 24%, 32%, 34%, 38%, 41%, and 45% were observed for the composites containing 0%, 20%, 40%, 80%, 80% and 100% β-TCP cements, respectively. In addition, the antibacterial activity of CS/β-TCP composite improves as the CS-content is increased. In vitro cell experiments show that the CS-rich composites promote human dental pulp cell (hDPC) proliferation and differentiation. However, when the CS quantity in the composite is less than 60%, the quantity of cells and osteogenesis protein of hDPCs is stimulated by Si released from the β-TCP/CS composites. The degradation of β-TCP and the osteogenesis of CS give strong reason to believe that these calcium-based composite cements will prove to be effective bone repair materials. © 2014 Elsevier B.V. All rights reserved.
1. Introduction Materials with varying degrees of bioactivity and degradation are required so as to conform to differing clinical requirements for hard tissue repair. Autograft possesses all the characteristics indispensable for new bone formation, namely, osteogenesis, osteoconductivity, and osteoinductivity, and it is currently considered the gold standard in this field. In spite of its strengths, however, there are still various disadvantages when using this material for certain medical clinical applications. Therefore, the design of an even better composite material would be very beneficial for controlling material physicochemical properties. Silicate-based ceramics have received a considerable amount of positive attention in recent years as these materials have better
⁎ Corresponding author. Tel.: +886 4 23590262x210. ⁎⁎ Corresponding author. Tel.: +886 4 24718668x55511; fax: +886 4 24759065. E-mail addresses: [email protected] (C.-C. Lin), [email protected] (M.-Y. Shie). 1 Both authors contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2014.06.030 0928-4931/© 2014 Elsevier B.V. All rights reserved.
bioactivity than calcium phosphate-based materials [1,2]. Recently, several studies have indicated that calcium silicate (CS) can play an important role in bone formation, at least based upon these materials' patterns of Si ion release [3] and its fast apatite formation ability [4]. CS can promote the proliferation of human mesenchymal stem cells (hMSCs), human dental pulp cells (hDPCs), and osteoblast-like cell adhesion, proliferation, and differentiation [5–9]. A previous study demonstrates that CS cement is biocompatible with cultured lipopolysaccharide-treated dental pulp cells and inhibits inflammation in lipopolysaccharide-treated dental pulp cells, leading to higher IL-1β expression and apoptosis [10]. Furthermore, this CS cement is able to activate angiogenesis [11] and anti-osteoclastogenesis [12]. These studies demonstrate that CS has the ability to promote the apatite layer's formation in vitro [13] and also show that CS is better at promoting bone regeneration in vivo [14,15]. However, the slow degradation rate of CS may result in a decrease in osteoconductivity, which may limit its clinical application [13]. In order to ameliorate its relative disadvantage in regard to material degradation, we used β-tricalcium phosphate (β-TCP) as an additive to see how it would affect CS's rate of decay. βTCP is a bioceramic material that is widely used for hard tissue repair.
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It has a chemical composition similar to apatite present in bone tissue, and has been applied extensively as a bone grafting material [16–18]. Previous studies have shown that wollastonite-doped TCP bioceramics have a higher degradation rate and promote new bone formation better than TCP in vivo [13,14]. Wang et al. assert that composite scaffolds containing 50% and 80% CS not only have good osteoconductivity, but also promote rapid bone formation compared with scaffolding comprised of pure β-TCP and CS [19]. Thus, to obtain both osteostimulation and osteoconductivity by taking advantage of the favorable bioactivity of calcium silicate and the high degradability of β-TCP, β-TCP/CS composite materials have been produced in the hopes that the right mixture can both help to control the degradation rate and improve interactions of the material with human tissue. In this study, β-TCP/CS composite cements with varied ratios were prepared so that we could observe the changes in physiochemical properties, bioactivity, in vitro degradation behavior, osteogenesis, angiogenesis and anti-bacterial activity in composites with different β-TCP/CS ratios. It is our hope that this knowledge may prove useful in the design of optimal biomaterials for bone regeneration. 2. Materials and methods 2.1. Preparation of β-TCP/CS composites The method used here for the preparation of CS powder has been described elsewhere [12]. In brief, reagent grade SiO2 (High Pure Chemicals, Saitama, Japan), CaO (Riedel-de Haen, Steinheim, Germany), Al2O3 (Sigma-Aldrich, St. Louis, MO) and ZnO (Wako, Osaka, Japan) powders were used as matrix materials (composition: 65% CaO, 25% SiO2, 5% Al2O3, and 5% ZnO) and mixed at 1500 rpm for 15 min using a hybrid defoaming mixer (ARE-250, Thinky, Tokyo, Japan). The oxide mixtures were then sintered at 1400 °C for 2 h using a high-temperature furnace and then ball-milled in ethyl alcohol using a centrifugal ball mill (S 100, Retsch, Haan, Germany) for 6 h. The β-TCP/CS composite material was obtained by mixing β-TCP (Sigma-Aldrich) and CS powder with composite weight ratios of 80:20 (C8T2), 60:40 (C6T4), 40:60 (C4T6), and 20:80 (C2T8) wt.% at 1500 rpm for 15 min using a hybriddefoaming mixer. The code of the pure CS and β-TCP was C10T0 and C0T10, respectively. The β-TCP/CS powder was mixed using a liquid/ powder ratio of 0.35 mL/g. After mixing with liquids (1 M Na2HPO4), the cements were molded in a Teflon mold (diameter: 6 mm, height: 3 mm). The cement quantities were enough to fully cover each well of the 24-well plate (GeneDireX, Las Vegas, NV) to a thickness of 2 mm for cell experiments. All samples were stored in an incubator at 100% relative humidity and 37 °C for 1 day of hydration.
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(ISO) 9917-1. The setting time was recorded when the Gilmore needle failed to create a 1-mm deep indentation in three separate areas. 2.4. In vitro soaking To evaluate the in vitro bioactivity, the cements were immersed in a 10 mL simulated body fluid (SBF) solution at 37 °C. The SBF solution, of which the ionic composition is similar to that of human blood plasma, consisted of 7.9949 g of NaCl, 0.3528 g of NaHCO3, 0.2235 g of KCl, 0.147 g of K2HPO4, 0.305 g of MgCl2·6H2O, 0.2775 g of CaCl2, and 0.071 g of Na2SO4 in 1000 mL of distilled H2O and was buffered to a pH of 7.4 with hydrochloric acid (HCl) and trishydroxymethyl aminomethane (Tris, CH2OH)3CNH2) [20,21]. All chemicals used were of reagent grade. The solution in the shaker water bath exhibited no change under static conditions. After soaking for different time durations (3 days to 3 months), specimens were removed from the tube and evaluated for several physicochemical properties. 2.5. Mechanical properties The composite specimens without/with immersion in SBF were dried at 60 °C for 3 days. Next, diametral tensile strength (DTS) testing was conducted on an EZ-Test machine (Shimadzu, Kyoto, Japan) at a loading rate of 1 mm/min. The maximal compression load at failure was obtained from the recorded load–deflection curves. At least 10 specimens from each group were tested. 2.6. Weight loss The degree of degradation was determined by monitoring the weight change of the specimens. After drying at 60 °C, the composites were weighed using a balance both before and after immersion. The ten specimens were examined at each time point for each of the materials being investigated. 2.7. Ion concentration The Ca, Si, P, Zn, and Al ion concentrations released from the cement on the SBF were determined using an inductively coupled plasmaatomic emission spectrometer (ICP-AES; Perkin-Elmer OPT 1MA 3000DV, Shelton, CT, USA) after the samples had been immersed for various periods of time. Three samples were measured for each data point and the results were obtained in triplicate from three separate samples for each test.
2.2. Phase composition and morphology
2.8. Dental pulp cell isolation and culture
The phase composition of the cements was analyzed using X-ray diffractometry (XRD; Bruker D8 SSS, Karlsruhe, Germany), and operated at 30 kV and 30 mA at a scanning speed of 1°/min. The morphology of the cement specimens was coated with gold and examined under a scanning electron microscope (SEM; JSM-6700F, JEOL) operated in the lower secondary electron image (LEI) mode at 3 kV accelerating voltage. Scanning electron microscopy associated with energy dispersive spectroscopy (EDS) was used to characterize the chemical analyses of the specimens.
The human dental pulp cells (hDPCs) were freshly derived from caries-free, intact premolars that had been extracted for orthodontic treatment purposes, as described previously [11]. The patient gave informed consent, and approval from the Ethics Committee of the Chung Shan Medicine University Hospital was obtained (CSMUH No. CS11187). A sagittal split was performed on each tooth using a chisel, and the pulp tissue was immersed in a PBS buffer solution. The pulp tissue was then cut into fragments, distributed onto plates and cultured in DMEM. These cultures were supplemented with 20% fetal bovine serum (FBS; Caisson), 1% penicillin (10,000 U/mL)/streptomycin (10,000 mg/mL) (PS, Caisson) and kept in a humidified atmosphere with 5% CO2 at 37 °C; the medium was changed every 3 days. The osteogenic differentiation medium was DMEM supplemented with 10−8 M dexamethasone (Sigma-Aldrich), 0.05 g/L L-ascorbic acid (Sigma-Aldrich) and 2.16 g/L glycerol 2-phosphate disodium salt hydrate (Sigma-Aldrich).
2.3. Setting time After the powder was mixed with H2O, the composites were placed into a cylindrical mold and stored in an incubator at 37 °C and 100% relative humidity for hydration. The setting time of the cements was tested according to standards set by the International Standards Organization
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2.9. Cell viability Cell suspensions at a density of 104 cells/mL were directly seeded over each specimen for various numbers of days. Cell cultures were incubated at 37 °C in a 5% CO2 atmosphere. After different culturing times, cell viability was evaluated using the PrestoBlue® (Invitrogen, Grand Island, NY) assay. At the end of the culture period, the medium was discarded and the wells were washed with cold PBS twice. Each well was then filled with a medium with a 1:9 ratio of PrestoBlue® in fresh DMEM and incubated at 37 °C for 30 min after which the solution in each well was transferred to a new 96-well plate. Plates were read in a multiwell spectrophotometer (Hitachi, Tokyo, Japan) at 570 nm with a reference wavelength of 600 nm. Cells cultured on a tissue culture plate without the cement were used as a control (Ctl). The results were obtained in triplicate from three separate experiments in terms of optical density (OD).
transferred to a new 96-well plate. Plates were then read in a multiwell spectrophotometer (Hitachi, Tokyo, Japan) at 570 nm with a reference wavelength of 600 nm. Cells cultured on the tissue culture plate without the cement were used as a control (Ctl). The results were obtained in triplicate from three separate experiments in terms of optical density (OD).
2.14. Statistical analysis A one-way analysis of the variance statistical data was used to evaluate the significance of the differences between the means in the measured data. Scheffe's multiple comparison test was used to determine the significance of the deviations in the data for each specimen. In all cases, the results were considered statistically significant with p value b 0.05.
2.10. Osteogenesis assay 3. Results and discussion The level of ALP activity was determined after the cells had been seeded for 3 and 7 days. The process was as follows: the cells were lysed from disks using 0.2% NP-40 and centrifuged for 10 min at 2000 rpm after washing with PBS. ALP activity was determined using p-nitrophenyl phosphate (pNPP, Sigma) as the substrate. Each sample was mixed with pNPP in 1 M diethanolamine buffer for 15 min, after which the reaction was stopped by the addition of 5 N NaOH and quantified by absorbance at 405 nm. All experiments were done in triplicate. The OC protein released from the pulp cells was cultured on different substrates for 7 and 14 days after cell seeding. Following the manufacturer's instructions, an osteocalcin enzyme-linked immunosorbent assay kit (Invitrogen) was used to determine the OC protein content. The OC protein concentration was measured by correlation with a standard curve. The analyzed blank disks were treated as controls. All experiments were done in triplicate. 2.11. Alizarin Red S stain Accumulated calcium deposition was observed after 7 and 14 days using Alizarin Red S staining as described in a previous study [3]. To summarize briefly, the cells were fixed with 4% paraformaldehyde (Sigma-Aldrich) for 15 min and then incubated in 0.5% Alizarin Red S (Sigma-Aldrich) at pH 4.0 for 15 min at room temperature in an orbital shaker (25 rpm). To quantify the stained calcified nodules after staining, samples were immersed with 1.5 mL of 5% SDS in 0.5 N HCl for 30 min at room temperature, following which the tubes were centrifuged at 5000 rpm for 10 min after which the supernatant was transferred to the new 96-well plate (GeneDireX). At this time, absorbance was measured at 405 nm (Hitachi). In addition, we also quantify the stained specimens without culturing cells (0 day).
3.1. Characterization of β-TCP/CS biocomposites Fig. 1 shows the XRD patterns of the composite powder after hydration for 1 day. The results indicate both that the composite samples consist of β-TCP and CS and that no chemical reaction occurred between the CS and TCP. The analysis of the specimens containing CS reveals two key points: first, an obvious diffraction peak near 2θ = 29.4°, which corresponds to the calcium silicate hydrate (CSH) gel, and second, incompletely reacted inorganic component phases of the βdicalcium silicate (β-Ca2SiO4) at 2θ between 32° and 34° [8]. It is clear that the addition of CS results in lower peak intensities of the CSH and β-Ca2SiO4 phases. Previously, Ni et al. performed a detailed study of the phase diagram of the system Ca2SiO4/Ca3(PO4)2 and their results indicate that the system is a true binary [22]. Our results are similar. Fig. 2 shows that with an increase in the amount of CS the setting time of the cement becomes shorter, going from 37 min (β-TCP group) all the way down to 16 min (CS group), a significant difference (p b 0.05). Under clinical conditions the setting time is a critical factor and a long setting period can lead to clinical problems. Fernández et al. suggest that 10–15 min is a suitable setting time interval [23]. In the present study, the setting time of calcium-silicate based cements was proportional to the CSH amount [7]. CSH is able to reduce the setting time of the calcium-based cement. Our results show that a setting
2.12. Intracellular Ang-1 and vWF measurement The production of Ang-1 and vWF was quantified using ELISA kits (Abcam, catalog nos. ab99970 and ab108918) according to the manufacturer's instructions. The hDPCs were cultured on substrates for 3 and 7 days, and proteins from whole cell lysates were collected and quantified using the ELISA kit. 2.13. Antibacterial property To investigate the anti-bacterial effect of the β-TCP/CS composite, the hydration composites were mixed with 1 mL Staphylococcus aureus in LB culture media (4.0 × 104 bacteria per mL) and cultured for 3 and 12 h. Aliquots of 0.1 mL from each group were then mixed with 0.9 mL PrestoBlue® for 30 min after which the solution in each well was
Fig. 1. XRD patterns of β-TCP/CS composites have different ratios after hydration.
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Dissolubility is a key factor in biodegradation [22] and must be considered when developing a material which has a degradation rate most appropriate to hastening and easing the process of tissue repair. With this in mind, the degradation rates of the β-TCP/CS composites in SBF solution have been recorded for various periods of time ranging from 3 to 84 days, as recorded in Fig. 3. After soaking for 3 days, the CS cement shows a relatively modest weight loss (~ 11%), whereas the β-TCP-containing cement lost considerably more weight (13–18%). All the specimens display an increased weight loss as the immersion time is increased. The pure CS cement (T0C10) has the lowest dissolution rate and solubility compared with other samples over the whole soaking period, reaching 24% after 12 weeks. By contrast, the overall amount of β-TCP dissolution reached 45% after 12 weeks. At the end of the immersion period, weight losses of approximately 32%, 34%, 38%, and 41%, were observed for the C2T8, C4T6, C6T4, and C8T2 cement
mixtures, respectively, indicating significant differences (p b 0.05). As expected, the β-TCP/CS biphasic composites show an intermediate dissolution behavior between those of pure β-TCP and CS. Moreover, as the CS content is increased, the dissolution rate also decreases steadily. Hence, the degradation rate of the composite cement may be controlled to a certain extent by varying the CS content in the composite [22]. Changes in the strength values of cement samples after immersion in SBF are shown in Fig. 4. The DTS values of hydration cements that are not immersed range from 2.1–3.5 MPa, indicating a significant (p b 0.05) decrease in the strength as the amount of β-TCP is increased. The decrease in the strength of the composite cement is probably due to the addition of the inherently weak β-TCP to CS. In addition, there is no reaction or chemical bonding between the CS and β-TCP in the cement; the two substances only stay together because they are mixed in the composite cements; they do not otherwise interact. After immersion in SBF, the mechanical properties change into biodegradable biomaterials. These results demonstrate that the CS-containing cement increases the DTS after 4 weeks of immersion, and thereafter decreases. When cement specimens are immersed in solution for 4 weeks, the CS hydration reaction dramatically changes in the CSH phase and increases in strength [13,26]. After 12 weeks of soaking, the strengths of C0T10, C2T8, and C4T6 are 1.4, 1.4, 2.2 MPa, respectively. This indicates that the strength is lower than cements that are not soaked. In contrast, the strength of the higher CS content composites (C6T4, C8T2 and C10T0) is similar to as-hardened cements. Moreover, the DTS of the higher β-TCP content cements declines due to the degradation, consistent with the results of weight loss. The results of an examination of the surface microstructure of the specimens before and after immersion in SBF after periods of 7 days are shown in Fig. 5. It is readily visible that the β-TCP cement has a looser and rougher surface texture, with irregular pores (Fig. 5F). In contrast, the pure CS cement exhibits a dense and smooth surface containing particle entanglement and micro-pores (Fig. 5A). In a bone graft, it is believed that when bonded to living bone, an apatite layer will form on the surface [4]. The formation of the bone-like apatite in SBF has proven to be useful in predicting the bone-bonding ability of material in vitro. It can be seen that the surfaces of specimens with 70% and 100% CS content are covered by a dense apatite layer after being immersed for 7 days. After immersion, the surface of the two samples is completely covered by an apatite layer with tiny spherical shapes on the surface. However, it is clear that no precipitate has formed on
Fig. 3. Weight loss of various cements after immersion in SBF for predetermined time durations. “*” indicates a significant difference (p b 0.05) compared to C10T0.
Fig. 4. Diametral tensile strength of various cements after immersion in SBF for predetermined time durations.
Fig. 2. The setting time of various β-TCP/CS composites hydrated. Statistical data are presented as means ± standard deviation for n = 6. Values not sharing a common letter are significantly different at p b 0.05.
time of approximately 20 min is required for injectable bone cements for clinical use [7,24,25]. 3.2. Immersion studies of β-TCP/CS biocomposites
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Fig. 5. SEM micrographs of the β-TCP/CS composites surfaces before and after immersion in SBF for 7 days. The white arrow indicated apatite sphere.
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the surfaces of the β-TCP-rich cement specimens (T10C0 and T7C3). For the T5C5 composite, after 7 days of immersion spherical granules that have precipitated on the surface of the composite become evident, and the morphology reveals an early stage of apatite precipitation. In addition, the Ca/P ratios of the C10T0 specimen after 1 and 4 weeks of immersion were 2.13 and 1.68, respectively. This precipitate has been identified to be apatite, although a Ca/P ratio of 1.68, which was appropriated Ca/P ratio (the stoichiometric Ca/P ratio, 1.67, of apatite) on 4week immersion surfaces. Similar apatite spherules have been observed
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to precipitate on CS materials surface after immersion in SBF [25]. These results are similar to reports by other researchers [14,22], who elucidate the lack of apatite formation observed in the β-TCP samples after soaking in SBF for several days. The in vitro bioactivity of the calciumions in the silicate based materials indicates that the presence of PO3− 4 composition is not an essential requirement for the formation of an apdepletes atite layer. This is noteworthy because it is known that PO3− 4 ions originate from the calcium and phosphate ions because the PO3− 4 in vitro assay solution [7]. Most importantly, the Si-OH functional
Fig. 6. (A) Ca, (B) Si, and (C) P ion concentrations of SBF after immersion for different times. Data represent means ± SD (n = 6). “*”, statistically significant difference from C10T0.
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group on CS materials has been proven to act as the nucleation center for apatite formation [4]. This may explain why the β-TCP cement had no bioactivity. Thus, the CS-rich cements will develop a stronger bond with the surrounding bone tissue compared with the β-TCP-rich cement. Variations of SBF Ca, Si, P, Zn and Al ion concentrations after being cultured for different time periods are shown in Fig. 6. After 12 weeks, the Ca ion concentration in the CS cement medium has decreased to approximately 1.2 mM, which is lower than the baseline Ca concentration of SBF (2.5 mM) (p b 0.05). As for the P ion concentration of SBF, it decreases after 1 day of immersion. The bioactivity of silicate-based materials indicates that the presence of PO43− ions in the composition is not an essential requirement for the development of an apatite layer, which consumes calcium and phosphate ions [4,24,27]. Si concentrations increase with increasing the duration of incubation. Si ion concentrations are about 1.5, 2.0, 2.3, 2.7, and 3.4 mM at C2T8, C4T6, C6T4, C8T2, and C10T0 after 12 weeks, respectively. However, the Zn and Al ions released from specimens were stable after immersion for 4 weeks. Calcium silicate cement undergoes hydrolysis on immersion, and Si ion dissolves during incubation [7,28]. Taking cell functions into account, the certain amounts of Si released from silicate-based materials may promote desired cell behaviors [29–31]. The current CS gives the ability to control the rate at which soluble Si ions affect cell adhesion and proliferation [7,32].
3.3. hDPC proliferation Cell viability and functions associated with a bone graft are closely related to the physical, chemical, and biological characteristics of the materials used. To consider the effects of CS on osteogenic activity, the biological functions of hDPCs cultured on various cement specimens have been evaluated after different periods of time (Fig. 7). Cell proliferation gradually increases as the amount of CS added to β-TCP is increased, which is notably different (p b 0.05) from the behavior of β-TCP cement. For example, on day 14 the C4T6 cement elucidated an increase of approximately 22% in the OD value compared to the pure β-TCP cement. Similarly, the β-TCP cement elicited a significant decrease of 12%, 17%, 30%, and 24% in comparison with the CS cement on days 1, 3, 7 and 14 of cell seeding, respectively.
Fig. 7. Proliferation of hDPCs treated with various specimens for 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
Fig. 8. (A) ALP activity of hDPCs cultured on various specimens for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
3.4. Osteogenesis protein secretion Osteogenesis differentiation of osteoblast-like cells is one of the key steps in determining whether bone formation is successful. The ALP expression of hDPCs cultured on different composites has also been examined, and Fig. 8 shows the analysis of quantitative examination data and the ALP activity of cells cultured on the different composites after 3 and 7 days. Increasing the CS content led to increased ALP levels in the cements at all incubation times. A significant 29% and 35% increase (p b 0.05) in the ALP level was measured for C4T6 and C6T4 in comparison with the β-TCP cement after 14 days. ALP enzyme activity is also associated with bone formation, and it is produced in high levels during the bone formation phase [7]. Similarly, the OC secretion by cells cultured on T0C10 is significantly higher (p b 0.05) than the levels secreted by cells on other substances (Fig. 9). After a 14 day period, the OC secretion was found to have been increased by 14% and 26% in the cells cultured on C4T6 and C6T4 in comparison with the β-TCP cement. Recent studies also show that CS promotes hDPC proliferation and differentiation [5,7]. This stimulatory effect may be attributed to the
Fig. 9. (A) OC amount of hDPCs cultured on various specimens for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
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Fig. 10. Quantification of calcium mineral deposits by Alizarin Red S assay of hDPCs cultured on various cement for 7 and 14 days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
dissolution of silicate ions [3,8]. It is worth noting that the β-TCP-rich cements with FGF-2 stimulate significantly more (p b 0.05) osteogenesis protein secretion than the pure cements. 3.5. Mineralization The aim of this assay is to determine and show the effects of various ratios of β-TCP/CS cement on bone matrix formation following analysis using Alizarin Red S staining to identify calcium deposition, as seen in Fig. 10. First, the results of quantification of all specimens without cell (0 day) were similar between all specimens (p N 0.05). Therefore, we consider that the method is suitable in this study. In the case of the C0T10 and C2T8 composites, no significant differences (p N 0.05) in the quantities of the calcium mineral matrix deposition were detected after either 7 or 14 days. By contrast, on day 14 a significant (1.18-, 1.32-, and 1.37-fold on C4T6, C6T4, and C8T2, respectively) enhancement (p b 0.05) of calcium content has been observed compared with β-TCP cement. According to the literature, tricalcium phosphate cement always fails to form a chemical bond with bone tissue during the early stages of therapy because of its poor bioactivity [22]. Taken together, our result indicates that the composite cement may possess a greater ability to interact with the hDPCs and to promote osteogenesis and matrix mineralization than the β-TCP cement. Such a system would combine the benefits of composite properties, such as resorbability (the ability to be replaced by native bone tissue), and osteoconductivity (the ability to induce bone growth). 3.6. Angiogenesis The protein expression levels of Ang-1 and vWF in hDPCs cultured on various specimens were evaluated on days 3 and 7 (Fig. 11). The Ang-1 expression in C4T6, C6T4, and C8T2 on day 7 was enhanced 1.54, 1.78, and 1.99 times, respectively, as compared to that of C0T10 (Fig. 11A). The vWF results shown in Fig. 11B can be seen to be similar to Ang-1, both showing a dose-dependent up-regulation in the specimens in which the CS levels have been increased (p b 0.05). vWF is an important protein involved in coagulation and thrombus formation. Following synthesis, it can be found in secretary granules called Weibel–Palade bodies and in vessels, and is released both constitutively and in a regulated manner [33,34]. Ang-1 is another family of growth
Fig. 11. The protein expression of (A) Ang-1 and (B) vWF on hDPCs cultured on β-TCP/CS cement for different days. “*” indicates a significant difference (p b 0.05) compared to C0T10.
factors that plays an important role in vascular development [35,36]. Our data shows that cement with a higher CS content has a higher potential to induce vWF and Ang-1 expression than the β-TCP cement during angiogenesis. In addition, Si ions have been shown to increase the osmolality required to facilitate the angiogenic differentiation of hDPCs via the MAPK/p38 pathway [11]. 3.7. Antibacterial property Besides the osteostimulation and angiogenesis properties of the CS/βTCP composite, these biocomposites have been shown to have significant
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Our results suggest that the incorporation of CS into β-TCP is a useful approach for obtaining composites with improved properties in regards to setting time, degradation, osteogenesis, angiogenesis, and antibacterial activity. Thus, we suggest 60 wt.% CS-containing β-TCP cement may be the best choice for hard tissue repair applications. Acknowledgments The authors acknowledge receipt of a grant from the National Science Council Taiwan grants (NSC 101-2314-B-040-011-MY3). The authors declare that they have no conflicts of interest. References
Fig. 12. The anti-bacterial effect of β-TCP/CS cement for 3 and 12 h. “*” indicates a significant difference (p b 0.05) compared to Ctl. “#” indicates a significant difference (p b 0.05) compared to Ca(OH)2.
antibacterial effects, as seen in Fig. 12. The Staphylococcus aureus cultures with Ctl and C0T10 show the highest absorbance at both the times observed (3 and 12 h) (p b 0.05). However, the antibacterial activity of CS/β-TCP composite increases in proportion with an increase in the CS-content. No significant difference was detected in the absorbance readings for C6T4, C8T2 and C10T0 groups compared to the Ca(OH)2 groups (p N 0.05). There may be two potential reasons for the antibacterial effect of the CS/β-TCP composite. One is that CS can release Ca and Si ions, leading to a weak alkaline microenvironment with a higher pH value which may restrain bacterial growth [37]; the other is that the CS particles themselves may permeate the bacteria [38] with possible antibacterial consequences. Whatever the cause, the CS/β-TCP composite may prove to be an excellent antibacterial agent. 4. Conclusions In this study, degradable and highly bioactive calcium-silicate based composite cement containing CS and β-TCP in different ratios was prepared and analyzed. Our research reveals that the mechanical properties of biocomposites increase steadily with an increase in CS portion in the ratio of various biocomposites. The dissolution rate of β-TCP/CS is strongly dependent on the β-TCP content, and β-TCP/CS composites with a lower CS content show higher dissolution rates. When the concentration of CS in the composite increases to 60%, the ability to form bone-like apatite is about the same as that for the pure CS. The results obtained in this study may be useful for designing calcium-based biocomposites with optimal biological and degradation properties.
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