Materials Science and Engineering C 35 (2014) 70–76
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Bioactive calcium sulfate/magnesium phosphate cement for bone substitute applications Guangyong Yang a,c,1, Jianli Liu b,d,1, Fan Li c, Zongyou Pan c, Xiao Ni c, Yue Shen c, Huazi Xu c,⁎, Qing Huang d,⁎ a
Department of Orthopaedics, Taizhou Hospital of Zhejiang Province, Linhai Zhejiang, 317000, China Trauma Center, Affiliated Hospital of Hainan Medical University, Haikou, Hainan, 570206, China Department of Orthopaedic Surgery, Second Affiliated Hospital of Wenzhou Medical College, Wenzhou, Zhejiang, 325000, China d Division of Functional Materials and Nanodevices, Ningbo Institute of Materials Technology and Engineering (NIMTE), Chinese Academy of Sciences (CAS), Ningbo, Zhejiang, 315201, China b c
a r t i c l e
i n f o
Article history: Received 6 April 2013 Received in revised form 29 September 2013 Accepted 19 October 2013 Available online 31 October 2013 Keywords: Calcium sulfate/magnesium phosphate cement (CSMPC) Bone substitutes Bioactivity Degradability Cytotoxicity
a b s t r a c t A novel calcium sulfate/magnesium phosphate cement (CSMPC) composite was prepared and studied in the present work. The physical properties including the phases, the microstructures, the setting properties and the compressive strengths of the CSMPCs were studied. The bio-performances of the CSMPCs were comprehensively evaluated using in vitro simulated body fluid (SBF) method and in vitro cell culture. The dependence of the physical and chemical properties of the CSMPC on its composition and microstructure was studied in detail. It is found that the CSMPC composites exhibited mediate setting times (6–12 min) compared to the calcium sulfate (CS) and the magnesium phosphate cement (MPC). They showed an encapsulation structure in which the unconverted hexagonal prism CSH particles were embedded in the xerogel-like MPC matrix. The phase compositions and the mechanical properties of the CSMPCs were closely related to the content of MPC and the hardening process. The CSMPCs exhibited excellent bioactivity and good biocompatibility to support the cells to attach and proliferate on the surface. The CSMPC composite has the potential to serve as bone grafts for the bone regeneration. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Bone grafting is the most common tissue transplantation except for blood transfusion [1]. The bone substitutes perform to provide support, fill voids, and stimulate bone-healing in orthopedics and dentistry. Autograft, harvested from the patient's own iliac crest, is considered to be the gold standard for the bone grafting. However, the problems of the autografting include the limited supply, the morbidity of the donor site, and the risk of infection and fracture [2,3]. Allograft is regarded as the second option to use the grafts from the other donors or animals [4]. However, the risks of transferring viral diseases, immunogenicity as well as rejection responds exist to restrict its application [5–7]. Over the past several decades, synthetic bone substitutes have attracted more and more interests. Plenty of biomaterials have been studied for the bone grafting applications [8–13], including the metals (stainless steels, titanium and its alloys and magnesium and its alloys), the ceramics (calcium hydroxyapatite, calcium phosphate and calcium sulfate) and the polymers (polylactic acid and polylactic–co-glycolic acid). More recently, the composite biomaterials are widely studied for the complex requirements for the clinical implantation [14].
⁎ Corresponding authors. E-mail addresses: [email protected] (H. Xu), [email protected] (Q. Huang). 1 Equal contribution authors. 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.10.016
Calcium sulfate (CS) has a long history to be used in orthopedics and dentistry due to its excellent biocompatibility, osteoconduction and complete resorption [15–17]. Despite these virtues, the CS has been criticized for several shortcomings. One is that the mechanical strength of the CS is low and not able to provide adequate long-term mechanical support for the osseous defect site [18]. The other problem is that the CS shows nearly no bioactivity which causes a poor bonding between the CS graft and the tissue [19–22]. The most important is that the CS resorbs too fast after implantation to match the bone regeneration. Magnesium phosphate cement (MPC) is a rapid-repair material with the characteristics of fast setting and high early setting strength [23]. Recently, much attention has been paid to use the MPC for the bone regeneration [24–27]. Studies have proved that the MPC had good biocompatibility without mutagenic and carcinogenicity responds after implantation [28]. It has been discovered that the MPC had good bioactivity to form apatite layer on the cement surface [29]. It is also found that the MPC was degradable after implantation and can release magnesium ion to promote the activity of the osteoblasts [30,31]. However, the setting of the MPC is too fast about 3 min, which restricts the application of the MPC as the binder for very small bone defects [32]. Therefore, it reasonable to develop new calcium sulfate/magnesium phosphate cements (CSMPCs) to obtain a bioactive bone substitute material with excellent biocompatibility, high early setting strength and controllable resorption rate. The aim of the present work is to prepare
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the CSMPC and investigate its physical and chemical properties as well as the in vitro bioperformances.
2. Materials and methods 2.1. Sample preparation and characterization All the starting chemicals were analytical grade and purchased from Sinopharm Chemical Reagent Co. Ltd. (Shanghai, China). α-Calcium sulfate hemihydrate (α-CSH, α-CaSO4 · 1/2H2O) powder was prepared using a salt solution method as described in our previous work [33]. In brief, 30 wt.% CaCl2 aqueous solutions with an ethanol/water volume ratio of 1 were prepared as the reaction mediums. The reaction solution was transferred into a three-neck flask and stirred using a magnetic stirring. Oil bathing of 90 °C was used to sustain the reaction temperature. In order to keep an invariant condition, a reflux condenser was connected to the flask for vapor reflux. Then, the calcium sulfate dihydrate (CSD, CaSO4· 2H2O) precursor was added into the reaction solution and reacted for 48 h. Finally, the α-CSH product was rapidly filtered and dried at 110 °C for 30 min in oven. The commercial magnesium oxide powders were pretreated at 1500 °C for 6 h to reduce the reactivity. After that, they were dry ball milled at 200 rpm for 2 h and sieved through a 200 mesh to obtain fine magnesium oxide powders. The MPC powders were then synthesized by mixing the treated magnesium oxide with monoammonium phosphate in a molar ratio of 3.8:1, in which magnesium oxide was the excess to serve as the second phase to reinforce the resulted MPC [23]. The CSMPC powders were prepared by mixing the CS powders and the MPC powders at various weight ratios as listed in Table 1. The pure CS and the MPC were used as the blank samples. The weighed CS and MPC powders were premixed using a pestle and mortar, and followed by dry ball milled at 200 rpm for 2 h to get homogeneous CSMPC powders. The obtained CS, MPC, and CSMPC powders were mixed with the distilled water at optimized ratios to form pastes. The setting times of the pastes were measured using a Vicat needle according to ISO95971989E. The time when the needle weighing 350 g, with 2.0 mm in diameter, could only penetrate no more than 1 mm into the sample was considered as the setting time. Each experiment was tested for six times, and the average value was calculated and expressed as means ± standard deviation (M ± SD).The setting time was listed in Table 1. The pastes were added into Teflon molds of Φ6 mm × 12 mm and Φ10 mm × 3 mm to prepare samples for the following compressive strength testing and the simulated body fluid (SBF) immersion, respectively. After taken out of the molds, the samples were maintained in a 100% humidity environment at 37 °C for 24 h to achieve sufficient strengths. A final hardening of the sample at 37 °C with 100% humidity for 28 days was further conducted on the samples for the complete crystallization. The changes of the phase, the morphology, and the compressive strength of the samples at various final hardening times of 7, 14, 21 and 28 days were analyzed by XRD (D8 Advance, Bruker Inc., Germany), FESEM (Hitachi S-4800, Hitachi Ltd., Japan) and Instron (Instron5985, USA). To test the compressive strength, six specimens were replicated for each group and the results were expressed as M ± SD.
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2.2. In vitro SBF immersion test The bioactivity and the degradation of the CSMPCs were evaluated using the in vitro SBF immersion method. The SBF solution was prepared through the procedure as described by Kokubo [34]. The CS, the MPC and the CSMPC samples of Φ10 mm × 3 mm were immersed in the SBF solution with a sample surface area/SBF volume ratio of 0.1 cm−1 and kept at 37 °C. All the samples were pretreated at 37 °C with 100% humidity for 28 days for the final hardening. Six samples for each CS, MPC and CSMPC were taken out of the SBF solution every 2 days. The samples were gently rinsed using the deionized water and dried in air. The weights of the samples were recorded to trace the weight loss of the materials. The pH values of the SBF solution during the test were measured and recorded every other day using an electrolyte-type pH meter (FE20K, Mettler Toledo, Switzerland). The values obtained from six samples were statistically calculated and expressed as M ± SD. The sample surfaces were examined using an invia-Reflex Raman spectrometer (RS, Renichaw, England) and a FESEM. 2.3. In vitro cell culture The cytotoxicity assay was conducted by culturing the MG63 osteoblast-like cells in the extracts of the materials using the Cell Counting Kit-8 (CCK-8, Dojindo Laboratories, Kumamoto, Japan) proliferation assay. The MG63 cells were grown in Dulbecco's modified Eagle's medium (DMEM, Gibco, USA) supplemented with 10% inactivated fetal bovine serum (FBS, Gibco, USA), 100 U mL−1 penicillin and 100 U mL−1 streptomycin at 37 °C, 5% CO2. In brief, the CS, MPC and CSMPC samples were crushed to powders and sieved to 200-mesh for further experiments. The extracts were prepared by adding the powder to DMEM cell culture medium for 1 day at 37 °C with 5% CO2 atmosphere. Then the mixture was centrifuged and the supernatant collected. The cell suspension were seeded at a density of 1 × 104 per well in 96-well plates. After an attachment period of 24 h, the cells were cultured in the presence of 100 μL of extracts for 72 h. The DMEM with 10% FBS without the addition of extracts was used as negative control. After culturing for 1, 3 and 5 days, 10 μL of kit reagent was added to the cells in 96-well plates and incubated for 1 h at 37 °C. The optical density (OD) of each well at 450 nm was evaluated using a microplate reader (Infinite M200 Pro, Tecan Group Ltd., Switzerland). Six samples per group were tested for the statistical analysis. The morphology of MG63 cells seeded on the CSMPC60 samples was examined using SEM after culturing for 4 days and 7 days at 37 °C, 100% humidity with 5% CO2 atmosphere. The cell/sample constructs were washed twice with PBS and fixed with 3% glutaraldehyde solution for 2 h at 4 °C. Then they were dehydrated in a series of graded ethanol (50, 60, 70, 80, 90 and 100% v/v) for 3 min and air-dried in a desiccator overnight. Finally, the samples were sputtered with platinum for SEM observation. 2.4. Statistical analysis Statistical analysis was performed using one-way ANOVA followed by post hoc tests and expressed as M ± SD. Statistical significance was set at p b 0.05. 3. Results and discussion
Table 1 Composition of different cements and their setting times.
3.1. Setting properties and compression strength of the CSMPC
Sample name
MPC percent (wt.%)
CS percent (wt.%)
Setting time
CS CSMPC40 CSMPC50 CSMPC60 MPC
0 40 50 60 100
100 60 50 40 0
25 12 7 6 3
MPC: magnesium phosphate cement; CS: calcium sulfate.
± ± ± ± ±
0.5 0.4 0.3 0.5 0.5
Table 1 compared the setting times of the CSMPC samples with that of the CS and the MPC. The MPC showed very fast setting for about 3 min while the CS exhibited a relative long setting time of about 25 min. The CSMPCs had setting times between those of the two components. With the increase of the MPC content, the setting time of the CSMPC decreased, indicating that the function of the MPC in the setting
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Fig. 1. SEM images of (a) the obtained CSH powders and (b) the CS, (c) the MPC, and (d) the CSMPC60 cements after setting at 37 °C in 100% relative humidity for 24 h.
of the CSMPC became obvious. This result provides a methodology to prepare the CSMPC with controlled setting time by adjusting the MPC content in the composite. For the facility of the surgical operation, the setting time of the cement is required to be between 8 and 15 min [35]. The MPC cement was blamed for its too fast setting to leave enough time for operation while the CS cement has to wait before operation. In the present work, the CSMPC composites exhibited mediate setting times (6–12 min) between the CS and the MPC. Such a setting time is proper for a surgeon to inject the cement or modulate it to the required shape for the clinical application. Fig. 1 shows the SEM images of the obtained CSH powders and the CS, the MPC, and the CSMPC60 cements after setting for 24 h at 37 °C with 100% relative humidity. In Fig. 1a, the CSH powder showed well crystallized hexagonal prism crystals with a length of 30 μm and a length ratio of 1. After setting, the CS cement showed a microstructure with the interlocked flaky CSD crystals and high porosity (in Fig. 1b). The MPC cement was in the form of dense xerogel with dispersed fine particles as shown in Fig. 1c. In the CSMPC60 cement (in Fig. 1d), the unconverted hexagonal prism CSH particles were embedded in the
xerogel-like MPC matrix. It is clear that the MPC addition hindered the transformation of the CSH to the CSD. This can be explained by that the MPC cement phase wrapped the CSH crystals and caused the difficulty for the water to penetrate and react with the CSH. Fig. 2 presents the XRD patterns of the prepared CSH powders (Fig. 2c) and the CSMPC60 (Fig. 2a), the MPC (Fig. 2b) and the CS (Fig. 2d) cements after setting for 24 h at 37 °C with 100% relative humidity. In the spectrum of the prepared CSH powders, all peaks were related to α-CSH (JCPDS card 74-1433) with the monoclinic structure. The CS cement exhibited a pure calcium sulfate dihydrate phase indicating a complete hydration of the CSH crystals. The MPC cement showed two phases of struvite and MgO. The struvite phase was the hydration product of the MPC and the MgO phase was the excess precursor. It is worth noting that there were composite phases of α-CSH, struvite and MgO in the CSMPC60 cement. It shows the fact that the CSH component in the CSMPC60 did not convert to the CSD phase after setting for 24 h, which was well accordant with the result in Fig. 1d. Because of no or incomplete transformation of the CSH to the CSD in the CSMPC60 after a short time setting, a hardening process at the
Fig. 2. XRD patterns of (c) the prepared CSH powders and (a) the CSMPC60, (b) the MPC and (d) the CS cements after setting at 37 °C in 100% relative humidity for 24 h.
Fig. 3. XRD patterns of the CSMPC60 after it was hardened at 37 °C in 100% relative humidity for various times of 1, 2, 3 and 4 weeks.
G. Yang et al. / Materials Science and Engineering C 35 (2014) 70–76 Table 2 Phase percents of the CSD in the CSMPC60 at various hardening time of 1, 2, 3 and 4 weeks. Hardening time (week)
1
2
3
4
R
0.6557
0.6997
0.76997
0.782
R is the phase percent of the CSD calculated by dividing the intensities of the highest peaks of the CSD at 2θ = 11.6° and the total CSH at 2θ = 14.7° and the CSD at 2θ = 11.6°.
the hardening process, besides the phases of struvite and MgO from the MPC component, both the CSH and CSD phase existed in the CSMPC60. With the increase of the hardening time from 1 week to 4 weeks, the CSH peaks became weaker and the CSD peaks became sharper. The transformation degree of the CSH to the CSD in the CSMPC60 was evaluated by the phase percent of the CSD according to the following equation [36] as listed in Table 2: R¼
Fig. 4. Compressive strengths of the CS, the MPC and the CSMPC samples after these were hardened at 37 °C in 100% relative humidity for various times of 1 day and 1, 2, 3 and 4 weeks. Statistically significant difference between the compressive strengths of the CSMPC50, CSMPC60 and CS, MPC, and CSMPC40 after these were hardened for 4 weeks (**p b 0.01, *p b 0.05).
moisture atmosphere was conducted on the CSMPC for a complete transformation. The phase evolution of the CSMPC60 during the hardening process at the moisture atmosphere was shown in Fig. 3. After
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ICSD ICSD þ ICSH
ð1Þ
where R is the phase percent of the CSD. ICSD and ICSH represent the intensities of the highest peaks of the CSD at 2θ = 11.6° and the CSH at 2θ = 14.7° respectively. It can be seen that the percent of the CSD in the CSMPC increased from 65.57% at 1 week hardening to 78.2% at 4 week hardening. During the hardening process, the cements were put in the moisture atmosphere for a long time. It allowed the water vapor to penetrate into the cement and react with the unconverted CSH. The changes of the compressive strengths of the CS, the MPC and the CSMPC samples during the hardening process are shown in Fig. 4. It can be found that the compressive strengths of the CS, the MPC and the CSMPC samples were all enhanced with the increase of the hardening times, indicating that the hardening process was important for the cements to achieve enough mechanical strengths. It is worth noting that when the MPC content was as low as 40 wt.%, the compressive strengths of the CSMPC40 were slightly lower than those of the CS and the MPC. However, with increasing the MPC content to 50 wt.% and 60 wt.%, the compressive strengths of the CSMPC sample were significantly enhanced to 65 MPa for the CSMPC50 and 71 MPa for the CSMPC60 after 28 days of hardening, respectively. The compressive
Fig. 5. SEM images of the surface morphologies of the CS, the MPC and the CSMPC60 samples after soaking in SBF for 1 day and 3 days.
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strengths of the CSMPC50 and CSMPC60 incubated up to 28 days were significantly higher than those of CS, MPC and CSMPC40, and no significant difference between the CS and CSMPC40. It reveals that the strength of the CSMPC was closely related to the content of MPC. Two mechanisms may take the effect on the dependence of the strength of the CSMPC on the MPC addition. One is the negative effect of the MPC to retard the hydration of the CSH. It has been proved by the microstructure and the phase of the CSMPC60 cement in Figs. 1d and 2d. The other one is that the MPC serves as a glue to reinforce the composite. When the MPC content is low, the negative effect may be dominant. In the CSMPC system, two cement phases were formed during the setting as shown in the following equations: CaSO4 ·1=2H2 O þ 1:5H2 O→CaSO4 ·2H2 O þ Q
ð2Þ
NH4 H2 PO4 þ MgO þ 5H2 O→NH4 MgPO4 ·6H2 O þ Q
ð3Þ
Both of the hydration reactions were exothermic. As reported by Soudee and Pera, the hydration of the MPC is a highly exothermic reaction [37]. During the setting of the CSMPC, the hydration of the MPC would cause an enhancement in the temperature of the material, which is inevitable to influence the thermodynamic balance of the hydration of the CSH and cause the decrease of the strength. However, when the MPC content is higher than half of the composite, the glue function of the MPC becomes significant. The hydration product of the
MPC tends to fill the micropores between the CSD crystals and bind the CSD crystals together. 3.2. In vitro bioactivity and degradation and pH value change of the CSMPC Fig. 5 shows surface morphologies of the CS, the MPC and the CSMPC60 samples after soaking in SBF for 1 day and 3 days. There was no significant change on the surface of the CS samples after immersion (Fig. 5a and b). AS for MPC, hemispherical particles were deposited on the sample surface after 1 day immersion as shown in Fig. 5c, and the new formed hemispherical agglomerates grown bigger when MPC was incubated for 3 days (Fig. 5d). In the case of the CSMPC60 (Fig. 5e and f), a layer of spherical agglomerates were formed on the CSMPC60 sample surface. The spherical agglomerates were reported to be the typical morphology of hydroxyapatite [38]. The deposited layers on the sample surface were further identified using the Raman spectroscopy. The Raman spectra of the CSMPC60, the MPC and the CS before and after soaking in the SBF for 1 week were shown in Fig. 6. After soaking for 1 week, the CSMPC60 and the MPC both showed a new peak around 960 cm−1 in their Raman spectra (as shown in Fig. 6b and d), which was attributed to the P–O stretching vibration and assigned to the formation of crystalline hydroxyapatite [39–42]. This peak was not detected in the samples before soaking in the SBF (in Fig. 6a and c). It is noticed that the characteristic peak was not detected in the CS samples both before (Fig. 6e) and after soaking in the SBF (Fig. 6f), indicating that no crystalline hydroxyapatite formed.
Fig. 6. Raman spectra of the CSMPC60, the MPC and the CS. (a, c, e) Before soaking. (b, d, f) After soaking in SBF for 1 week.
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The peaks located at around 412 cm−1 and 492 cm−1 are attributed to the ν2 (SO4) symmetric bend mode, while peaks that appeared at 621 cm− and 670 cm− are associated with the ν2 (SO4) antisymmetric bend mode. Two other peaks at 1008 cm−1 and 1137 cm−1 are related to the ν1 (SO4) symmetric stretch and the ν3 (SO4) antisymmetric stretch, respectively [43,44]. It is a common notion that bioactive materials can bond to living bone through a new apatite layer that is deposited on their surfaces in contact with body fluid [45,46]. Bone-like hydroxyapatite plays an important role in the formation, growth and maintenance of the bone tissue–biomaterial interface [47]. As multiple in vivo studies have pointed out, bioactive materials were also proved to accelerate bone formation at implantation sites [48–51]. However, CS has been criticized for its no new apatite layer formed in the tissue–biomaterial interface because of its poor bioactivity [19–23], which was also proved by the fact that no apatite was deposited on the CS surface after incubating in SBF in our study. On the contrary, in the present study, the results of SEM and Raman analysis verify that the CSMPC can induce the formation of apatite in SBF within 3 days, indicating that the composite had excellent bioactivity. With the addition of the MPC, PO3− and OH− are released 4 during the dissolution of the composite. Local increases in Ca2+, PO43−, and OH− ions concentration on the CSMPC surface may promote the ionic activity, affect the nucleation of apatite crystals genesis and act as a stimulus to form newly spherical apatite deposition on the composite surface through a dissolution–deposition process. Thus, the CSMPC composite with good bioactivity would form a stronger bond between the implanted material and the surrounding bone tissue compared with the traditional CS material. Fig. 7 shows the weight loss of the CS, CSMPC40, CSMPC50, CSMPC60 and MPC samples after soaking in SBF for various time periods. It can be seen that the CS degraded very fast and the mass of 95.2% dissolved in the SBF for 28 days. While the degradation of the MPC was very slow and only the mass of 19.2% dissolved after 84 days of soaking. The weight losses of the CSMPC composites were remarkably slowed down than that of the CS due to the addition of the MPC. It also can be found that the weight losses of the CSMPC composites became slower with the higher MPC contents. Therefore, it is efficient to tune the degradation of the CSMPC by changing the MPC addition. The changes of pH values of the SBF solution during the sample immersion were shown in Fig. 8. During the SBF immersion, the CS caused the pH of the environment solution to decrease from 7.36 to 6.49 to form an acid environment. The MPC caused slight increase in the pH of the SBF to 7.42 after 15 days of immersion. The pH of the CSMPC60 immersed solution increased continuously and was alkaline around 9.0 after 15 days. It is reported that alkaline environments were helpful to formation of apatite and the nucleation rate of the hydroxyapatite
Fig. 7. Weight losses of the CS, the CSMPC40, the CSMPC50, the CSMPC60 and the MPC samples during soaking in SBF. Statistically significant difference between the degradation rates of CSMPCs and CS (**p b 0.01, *p b 0.05).
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Fig. 8. The changes of pH values of the SBF solution during the sample immersion for 15 days.
crystals is faster at pH 9 than at pH 7 [52]. These could further explain the phenomena that the apatite formed on the MPC and the CSMPC instead of the CS. 3.3. Cell proliferation and cell morphology Fig. 9 shows the proliferation of MG63 cells cultured on the CS, the MPC and the CSMPC60 samples after culturing for 1, 3 and 5 days. The optical density (OD) values from the CCK-8 assay provide an indicator of the cell growth and proliferation on various materials [53,54]. It can be seen that the cell proliferation on all three materials increased with time, which indicates that three materials cause no significant cytotoxicity against cells. Moreover, the proliferation of cell for the CSMPC60 was significantly higher than the CS after incubating for 3 and 5 days; no significant difference appeared after 1 day of culture. The above results indicate that the CSMPC60 can increase the cell viability in comparison to the CS. Fig. 10 shows the SEM micrographs of morphologies of MG63 cells cultured on the CSMPC60 surfaces. After culturing for 4 days (Fig. 10a), the cells firmly attached and spread well on the surface of CSMPC60. After culturing for 7 days as shown in Fig. 10b it was revealed that the cells reached a confluent layer on material surface and had contact
Fig. 9. CCK-8 assay showing the optical density (OD) of the MG63 cells on the CS, the MPC and the CSMPC60 samples after culturing for 1, 3 and 5 days. Significant difference (**p b 0.01).
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Fig. 10. SEM micrographs showing the attachment of the MG63 cells cultured on the CSMPC60 surfaces for (a) 4 days and (b) 7 days.
with each other. Thus, it can be concluded that CSMPC60 did not inhibit the cell proliferation and had a stimulatory effect on cell growth. 4. Conclusions In the present work, a novel bioactive calcium sulfate/magnesium phosphate cement (CSMPC) was fabricated to combine the advantages of the calcium sulfate cement and the magnesium phosphate cement.
[13] [14] [15] [16] [17] [18] [19] [20] [21]
• The prepared CSMPC composites featured the controllable setting time mechanical strength and degradation rate by adjusting the MPC content, which would provide the facility to fulfill the requirement for the clinical application. • The CSMPC composites exhibited good bioactivity to form bone-like apatite in the in vitro SBF test. • The CSMPC composites had good biocompatibility to support the cells to attach and proliferate on the surface. In summary, the CSMPC composite is a promising alternative of bone graft for the bone regeneration. Acknowledgement This work was supported by the Zhejiang Provincial Open Foundation of the Most Important Subjects (2011GK002), Zhejiang Provincial Science and Technology Program of China (No. 2011C23113) and Ningbo Natural Science Fund (2012A610101). References [1] A. Van Heest, M. Swiontkowski, Lancet 353 (Suppl. 1) (1999) SI28–SI29. [2] C.G. Finkemeier, J. Bone Joint Surg. Am. 84-A (2002) 454–464. [3] E. Ahlmann, M. Patzakis, N. Roidis, L. Shepherd, P. Holtom, J. Bone Joint Surg. Am. 84-A (2002) 716–720. [4] G. Carter, AORN J. 70 (1999) 660–670(quiz 672-666). [5] W.W. Tomford, J. Bone Joint Surg. (Am. Vol.) 77 (1995) 1742–1754. [6] P.V. Giannoudis, H. Dinopoulos, E. Tsiridis, Injury 36 (Suppl. 3) (2005) S20–S27. [7] D.M. Strong, G.E. Friedlaender, W.W. Tomford, D.S. Springfield, T.C. Shives, H. Burchardt, W.F. Enneking, H.J. Mankin, Clin. Orthop. Relat. Res. (1996) 107–114. [8] E.M. Ooms, E.A. Egglezos, J.G. Wolke, J.A. Jansen, Biomaterials 24 (2003) 749–757. [9] K. Ohura, M. Bohner, P. Hardouin, J. Lemaitre, G. Pasquier, B. Flautre, J. Biomed. Mater. Res. 30 (1996) 193–200. [10] A. Ogose, T. Hotta, H. Hatano, H. Kawashima, K. Tokunaga, N. Endo, H. Umezu, J. Biomed. Mater. Res. 63 (2002) 601–604. [11] N. Kondo, A. Ogose, K. Tokunaga, T. Ito, K. Arai, N. Kudo, H. Inoue, H. Irie, N. Endo, Biomaterials 26 (2005) 5600–5608. [12] S.K. Nandi, S. Roy, P. Mukherjee, B. Kundu, D.K. De, D. Basu, Indian J. Med. Res. 132 (2010) 15–30.
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Bioactive calcium sulfate/magnesium phosphate cement for bone substitute applications Guangyong Yang a,c,1, Jianli Liu b,d,1, Fan Li c, Zongyou Pan c, Xiao Ni c, Yue Shen c, Huazi Xu c,⁎, Qing Huang d,⁎ a
Department of Orthopaedics, Taizhou Hospital of Zhejiang Province, Linhai Zhejiang, 317000, China Trauma Center, Affiliated Hospital of Hainan Medical University, Haikou, Hainan, 570206, China Department of Orthopaedic Surgery, Second Affiliated Hospital of Wenzhou Medical College, Wenzhou, Zhejiang, 325000, China d Division of Functional Materials and Nanodevices, Ningbo Institute of Materials Technology and Engineering (NIMTE), Chinese Academy of Sciences (CAS), Ningbo, Zhejiang, 315201, China b c
a r t i c l e
i n f o
Article history: Received 6 April 2013 Received in revised form 29 September 2013 Accepted 19 October 2013 Available online 31 October 2013 Keywords: Calcium sulfate/magnesium phosphate cement (CSMPC) Bone substitutes Bioactivity Degradability Cytotoxicity
a b s t r a c t A novel calcium sulfate/magnesium phosphate cement (CSMPC) composite was prepared and studied in the present work. The physical properties including the phases, the microstructures, the setting properties and the compressive strengths of the CSMPCs were studied. The bio-performances of the CSMPCs were comprehensively evaluated using in vitro simulated body fluid (SBF) method and in vitro cell culture. The dependence of the physical and chemical properties of the CSMPC on its composition and microstructure was studied in detail. It is found that the CSMPC composites exhibited mediate setting times (6–12 min) compared to the calcium sulfate (CS) and the magnesium phosphate cement (MPC). They showed an encapsulation structure in which the unconverted hexagonal prism CSH particles were embedded in the xerogel-like MPC matrix. The phase compositions and the mechanical properties of the CSMPCs were closely related to the content of MPC and the hardening process. The CSMPCs exhibited excellent bioactivity and good biocompatibility to support the cells to attach and proliferate on the surface. The CSMPC composite has the potential to serve as bone grafts for the bone regeneration. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Bone grafting is the most common tissue transplantation except for blood transfusion [1]. The bone substitutes perform to provide support, fill voids, and stimulate bone-healing in orthopedics and dentistry. Autograft, harvested from the patient's own iliac crest, is considered to be the gold standard for the bone grafting. However, the problems of the autografting include the limited supply, the morbidity of the donor site, and the risk of infection and fracture [2,3]. Allograft is regarded as the second option to use the grafts from the other donors or animals [4]. However, the risks of transferring viral diseases, immunogenicity as well as rejection responds exist to restrict its application [5–7]. Over the past several decades, synthetic bone substitutes have attracted more and more interests. Plenty of biomaterials have been studied for the bone grafting applications [8–13], including the metals (stainless steels, titanium and its alloys and magnesium and its alloys), the ceramics (calcium hydroxyapatite, calcium phosphate and calcium sulfate) and the polymers (polylactic acid and polylactic–co-glycolic acid). More recently, the composite biomaterials are widely studied for the complex requirements for the clinical implantation [14].
⁎ Corresponding authors. E-mail addresses: [email protected] (H. Xu), [email protected] (Q. Huang). 1 Equal contribution authors. 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.10.016
Calcium sulfate (CS) has a long history to be used in orthopedics and dentistry due to its excellent biocompatibility, osteoconduction and complete resorption [15–17]. Despite these virtues, the CS has been criticized for several shortcomings. One is that the mechanical strength of the CS is low and not able to provide adequate long-term mechanical support for the osseous defect site [18]. The other problem is that the CS shows nearly no bioactivity which causes a poor bonding between the CS graft and the tissue [19–22]. The most important is that the CS resorbs too fast after implantation to match the bone regeneration. Magnesium phosphate cement (MPC) is a rapid-repair material with the characteristics of fast setting and high early setting strength [23]. Recently, much attention has been paid to use the MPC for the bone regeneration [24–27]. Studies have proved that the MPC had good biocompatibility without mutagenic and carcinogenicity responds after implantation [28]. It has been discovered that the MPC had good bioactivity to form apatite layer on the cement surface [29]. It is also found that the MPC was degradable after implantation and can release magnesium ion to promote the activity of the osteoblasts [30,31]. However, the setting of the MPC is too fast about 3 min, which restricts the application of the MPC as the binder for very small bone defects [32]. Therefore, it reasonable to develop new calcium sulfate/magnesium phosphate cements (CSMPCs) to obtain a bioactive bone substitute material with excellent biocompatibility, high early setting strength and controllable resorption rate. The aim of the present work is to prepare
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the CSMPC and investigate its physical and chemical properties as well as the in vitro bioperformances.
2. Materials and methods 2.1. Sample preparation and characterization All the starting chemicals were analytical grade and purchased from Sinopharm Chemical Reagent Co. Ltd. (Shanghai, China). α-Calcium sulfate hemihydrate (α-CSH, α-CaSO4 · 1/2H2O) powder was prepared using a salt solution method as described in our previous work [33]. In brief, 30 wt.% CaCl2 aqueous solutions with an ethanol/water volume ratio of 1 were prepared as the reaction mediums. The reaction solution was transferred into a three-neck flask and stirred using a magnetic stirring. Oil bathing of 90 °C was used to sustain the reaction temperature. In order to keep an invariant condition, a reflux condenser was connected to the flask for vapor reflux. Then, the calcium sulfate dihydrate (CSD, CaSO4· 2H2O) precursor was added into the reaction solution and reacted for 48 h. Finally, the α-CSH product was rapidly filtered and dried at 110 °C for 30 min in oven. The commercial magnesium oxide powders were pretreated at 1500 °C for 6 h to reduce the reactivity. After that, they were dry ball milled at 200 rpm for 2 h and sieved through a 200 mesh to obtain fine magnesium oxide powders. The MPC powders were then synthesized by mixing the treated magnesium oxide with monoammonium phosphate in a molar ratio of 3.8:1, in which magnesium oxide was the excess to serve as the second phase to reinforce the resulted MPC [23]. The CSMPC powders were prepared by mixing the CS powders and the MPC powders at various weight ratios as listed in Table 1. The pure CS and the MPC were used as the blank samples. The weighed CS and MPC powders were premixed using a pestle and mortar, and followed by dry ball milled at 200 rpm for 2 h to get homogeneous CSMPC powders. The obtained CS, MPC, and CSMPC powders were mixed with the distilled water at optimized ratios to form pastes. The setting times of the pastes were measured using a Vicat needle according to ISO95971989E. The time when the needle weighing 350 g, with 2.0 mm in diameter, could only penetrate no more than 1 mm into the sample was considered as the setting time. Each experiment was tested for six times, and the average value was calculated and expressed as means ± standard deviation (M ± SD).The setting time was listed in Table 1. The pastes were added into Teflon molds of Φ6 mm × 12 mm and Φ10 mm × 3 mm to prepare samples for the following compressive strength testing and the simulated body fluid (SBF) immersion, respectively. After taken out of the molds, the samples were maintained in a 100% humidity environment at 37 °C for 24 h to achieve sufficient strengths. A final hardening of the sample at 37 °C with 100% humidity for 28 days was further conducted on the samples for the complete crystallization. The changes of the phase, the morphology, and the compressive strength of the samples at various final hardening times of 7, 14, 21 and 28 days were analyzed by XRD (D8 Advance, Bruker Inc., Germany), FESEM (Hitachi S-4800, Hitachi Ltd., Japan) and Instron (Instron5985, USA). To test the compressive strength, six specimens were replicated for each group and the results were expressed as M ± SD.
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2.2. In vitro SBF immersion test The bioactivity and the degradation of the CSMPCs were evaluated using the in vitro SBF immersion method. The SBF solution was prepared through the procedure as described by Kokubo [34]. The CS, the MPC and the CSMPC samples of Φ10 mm × 3 mm were immersed in the SBF solution with a sample surface area/SBF volume ratio of 0.1 cm−1 and kept at 37 °C. All the samples were pretreated at 37 °C with 100% humidity for 28 days for the final hardening. Six samples for each CS, MPC and CSMPC were taken out of the SBF solution every 2 days. The samples were gently rinsed using the deionized water and dried in air. The weights of the samples were recorded to trace the weight loss of the materials. The pH values of the SBF solution during the test were measured and recorded every other day using an electrolyte-type pH meter (FE20K, Mettler Toledo, Switzerland). The values obtained from six samples were statistically calculated and expressed as M ± SD. The sample surfaces were examined using an invia-Reflex Raman spectrometer (RS, Renichaw, England) and a FESEM. 2.3. In vitro cell culture The cytotoxicity assay was conducted by culturing the MG63 osteoblast-like cells in the extracts of the materials using the Cell Counting Kit-8 (CCK-8, Dojindo Laboratories, Kumamoto, Japan) proliferation assay. The MG63 cells were grown in Dulbecco's modified Eagle's medium (DMEM, Gibco, USA) supplemented with 10% inactivated fetal bovine serum (FBS, Gibco, USA), 100 U mL−1 penicillin and 100 U mL−1 streptomycin at 37 °C, 5% CO2. In brief, the CS, MPC and CSMPC samples were crushed to powders and sieved to 200-mesh for further experiments. The extracts were prepared by adding the powder to DMEM cell culture medium for 1 day at 37 °C with 5% CO2 atmosphere. Then the mixture was centrifuged and the supernatant collected. The cell suspension were seeded at a density of 1 × 104 per well in 96-well plates. After an attachment period of 24 h, the cells were cultured in the presence of 100 μL of extracts for 72 h. The DMEM with 10% FBS without the addition of extracts was used as negative control. After culturing for 1, 3 and 5 days, 10 μL of kit reagent was added to the cells in 96-well plates and incubated for 1 h at 37 °C. The optical density (OD) of each well at 450 nm was evaluated using a microplate reader (Infinite M200 Pro, Tecan Group Ltd., Switzerland). Six samples per group were tested for the statistical analysis. The morphology of MG63 cells seeded on the CSMPC60 samples was examined using SEM after culturing for 4 days and 7 days at 37 °C, 100% humidity with 5% CO2 atmosphere. The cell/sample constructs were washed twice with PBS and fixed with 3% glutaraldehyde solution for 2 h at 4 °C. Then they were dehydrated in a series of graded ethanol (50, 60, 70, 80, 90 and 100% v/v) for 3 min and air-dried in a desiccator overnight. Finally, the samples were sputtered with platinum for SEM observation. 2.4. Statistical analysis Statistical analysis was performed using one-way ANOVA followed by post hoc tests and expressed as M ± SD. Statistical significance was set at p b 0.05. 3. Results and discussion
Table 1 Composition of different cements and their setting times.
3.1. Setting properties and compression strength of the CSMPC
Sample name
MPC percent (wt.%)
CS percent (wt.%)
Setting time
CS CSMPC40 CSMPC50 CSMPC60 MPC
0 40 50 60 100
100 60 50 40 0
25 12 7 6 3
MPC: magnesium phosphate cement; CS: calcium sulfate.
± ± ± ± ±
0.5 0.4 0.3 0.5 0.5
Table 1 compared the setting times of the CSMPC samples with that of the CS and the MPC. The MPC showed very fast setting for about 3 min while the CS exhibited a relative long setting time of about 25 min. The CSMPCs had setting times between those of the two components. With the increase of the MPC content, the setting time of the CSMPC decreased, indicating that the function of the MPC in the setting
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Fig. 1. SEM images of (a) the obtained CSH powders and (b) the CS, (c) the MPC, and (d) the CSMPC60 cements after setting at 37 °C in 100% relative humidity for 24 h.
of the CSMPC became obvious. This result provides a methodology to prepare the CSMPC with controlled setting time by adjusting the MPC content in the composite. For the facility of the surgical operation, the setting time of the cement is required to be between 8 and 15 min [35]. The MPC cement was blamed for its too fast setting to leave enough time for operation while the CS cement has to wait before operation. In the present work, the CSMPC composites exhibited mediate setting times (6–12 min) between the CS and the MPC. Such a setting time is proper for a surgeon to inject the cement or modulate it to the required shape for the clinical application. Fig. 1 shows the SEM images of the obtained CSH powders and the CS, the MPC, and the CSMPC60 cements after setting for 24 h at 37 °C with 100% relative humidity. In Fig. 1a, the CSH powder showed well crystallized hexagonal prism crystals with a length of 30 μm and a length ratio of 1. After setting, the CS cement showed a microstructure with the interlocked flaky CSD crystals and high porosity (in Fig. 1b). The MPC cement was in the form of dense xerogel with dispersed fine particles as shown in Fig. 1c. In the CSMPC60 cement (in Fig. 1d), the unconverted hexagonal prism CSH particles were embedded in the
xerogel-like MPC matrix. It is clear that the MPC addition hindered the transformation of the CSH to the CSD. This can be explained by that the MPC cement phase wrapped the CSH crystals and caused the difficulty for the water to penetrate and react with the CSH. Fig. 2 presents the XRD patterns of the prepared CSH powders (Fig. 2c) and the CSMPC60 (Fig. 2a), the MPC (Fig. 2b) and the CS (Fig. 2d) cements after setting for 24 h at 37 °C with 100% relative humidity. In the spectrum of the prepared CSH powders, all peaks were related to α-CSH (JCPDS card 74-1433) with the monoclinic structure. The CS cement exhibited a pure calcium sulfate dihydrate phase indicating a complete hydration of the CSH crystals. The MPC cement showed two phases of struvite and MgO. The struvite phase was the hydration product of the MPC and the MgO phase was the excess precursor. It is worth noting that there were composite phases of α-CSH, struvite and MgO in the CSMPC60 cement. It shows the fact that the CSH component in the CSMPC60 did not convert to the CSD phase after setting for 24 h, which was well accordant with the result in Fig. 1d. Because of no or incomplete transformation of the CSH to the CSD in the CSMPC60 after a short time setting, a hardening process at the
Fig. 2. XRD patterns of (c) the prepared CSH powders and (a) the CSMPC60, (b) the MPC and (d) the CS cements after setting at 37 °C in 100% relative humidity for 24 h.
Fig. 3. XRD patterns of the CSMPC60 after it was hardened at 37 °C in 100% relative humidity for various times of 1, 2, 3 and 4 weeks.
G. Yang et al. / Materials Science and Engineering C 35 (2014) 70–76 Table 2 Phase percents of the CSD in the CSMPC60 at various hardening time of 1, 2, 3 and 4 weeks. Hardening time (week)
1
2
3
4
R
0.6557
0.6997
0.76997
0.782
R is the phase percent of the CSD calculated by dividing the intensities of the highest peaks of the CSD at 2θ = 11.6° and the total CSH at 2θ = 14.7° and the CSD at 2θ = 11.6°.
the hardening process, besides the phases of struvite and MgO from the MPC component, both the CSH and CSD phase existed in the CSMPC60. With the increase of the hardening time from 1 week to 4 weeks, the CSH peaks became weaker and the CSD peaks became sharper. The transformation degree of the CSH to the CSD in the CSMPC60 was evaluated by the phase percent of the CSD according to the following equation [36] as listed in Table 2: R¼
Fig. 4. Compressive strengths of the CS, the MPC and the CSMPC samples after these were hardened at 37 °C in 100% relative humidity for various times of 1 day and 1, 2, 3 and 4 weeks. Statistically significant difference between the compressive strengths of the CSMPC50, CSMPC60 and CS, MPC, and CSMPC40 after these were hardened for 4 weeks (**p b 0.01, *p b 0.05).
moisture atmosphere was conducted on the CSMPC for a complete transformation. The phase evolution of the CSMPC60 during the hardening process at the moisture atmosphere was shown in Fig. 3. After
73
ICSD ICSD þ ICSH
ð1Þ
where R is the phase percent of the CSD. ICSD and ICSH represent the intensities of the highest peaks of the CSD at 2θ = 11.6° and the CSH at 2θ = 14.7° respectively. It can be seen that the percent of the CSD in the CSMPC increased from 65.57% at 1 week hardening to 78.2% at 4 week hardening. During the hardening process, the cements were put in the moisture atmosphere for a long time. It allowed the water vapor to penetrate into the cement and react with the unconverted CSH. The changes of the compressive strengths of the CS, the MPC and the CSMPC samples during the hardening process are shown in Fig. 4. It can be found that the compressive strengths of the CS, the MPC and the CSMPC samples were all enhanced with the increase of the hardening times, indicating that the hardening process was important for the cements to achieve enough mechanical strengths. It is worth noting that when the MPC content was as low as 40 wt.%, the compressive strengths of the CSMPC40 were slightly lower than those of the CS and the MPC. However, with increasing the MPC content to 50 wt.% and 60 wt.%, the compressive strengths of the CSMPC sample were significantly enhanced to 65 MPa for the CSMPC50 and 71 MPa for the CSMPC60 after 28 days of hardening, respectively. The compressive
Fig. 5. SEM images of the surface morphologies of the CS, the MPC and the CSMPC60 samples after soaking in SBF for 1 day and 3 days.
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strengths of the CSMPC50 and CSMPC60 incubated up to 28 days were significantly higher than those of CS, MPC and CSMPC40, and no significant difference between the CS and CSMPC40. It reveals that the strength of the CSMPC was closely related to the content of MPC. Two mechanisms may take the effect on the dependence of the strength of the CSMPC on the MPC addition. One is the negative effect of the MPC to retard the hydration of the CSH. It has been proved by the microstructure and the phase of the CSMPC60 cement in Figs. 1d and 2d. The other one is that the MPC serves as a glue to reinforce the composite. When the MPC content is low, the negative effect may be dominant. In the CSMPC system, two cement phases were formed during the setting as shown in the following equations: CaSO4 ·1=2H2 O þ 1:5H2 O→CaSO4 ·2H2 O þ Q
ð2Þ
NH4 H2 PO4 þ MgO þ 5H2 O→NH4 MgPO4 ·6H2 O þ Q
ð3Þ
Both of the hydration reactions were exothermic. As reported by Soudee and Pera, the hydration of the MPC is a highly exothermic reaction [37]. During the setting of the CSMPC, the hydration of the MPC would cause an enhancement in the temperature of the material, which is inevitable to influence the thermodynamic balance of the hydration of the CSH and cause the decrease of the strength. However, when the MPC content is higher than half of the composite, the glue function of the MPC becomes significant. The hydration product of the
MPC tends to fill the micropores between the CSD crystals and bind the CSD crystals together. 3.2. In vitro bioactivity and degradation and pH value change of the CSMPC Fig. 5 shows surface morphologies of the CS, the MPC and the CSMPC60 samples after soaking in SBF for 1 day and 3 days. There was no significant change on the surface of the CS samples after immersion (Fig. 5a and b). AS for MPC, hemispherical particles were deposited on the sample surface after 1 day immersion as shown in Fig. 5c, and the new formed hemispherical agglomerates grown bigger when MPC was incubated for 3 days (Fig. 5d). In the case of the CSMPC60 (Fig. 5e and f), a layer of spherical agglomerates were formed on the CSMPC60 sample surface. The spherical agglomerates were reported to be the typical morphology of hydroxyapatite [38]. The deposited layers on the sample surface were further identified using the Raman spectroscopy. The Raman spectra of the CSMPC60, the MPC and the CS before and after soaking in the SBF for 1 week were shown in Fig. 6. After soaking for 1 week, the CSMPC60 and the MPC both showed a new peak around 960 cm−1 in their Raman spectra (as shown in Fig. 6b and d), which was attributed to the P–O stretching vibration and assigned to the formation of crystalline hydroxyapatite [39–42]. This peak was not detected in the samples before soaking in the SBF (in Fig. 6a and c). It is noticed that the characteristic peak was not detected in the CS samples both before (Fig. 6e) and after soaking in the SBF (Fig. 6f), indicating that no crystalline hydroxyapatite formed.
Fig. 6. Raman spectra of the CSMPC60, the MPC and the CS. (a, c, e) Before soaking. (b, d, f) After soaking in SBF for 1 week.
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The peaks located at around 412 cm−1 and 492 cm−1 are attributed to the ν2 (SO4) symmetric bend mode, while peaks that appeared at 621 cm− and 670 cm− are associated with the ν2 (SO4) antisymmetric bend mode. Two other peaks at 1008 cm−1 and 1137 cm−1 are related to the ν1 (SO4) symmetric stretch and the ν3 (SO4) antisymmetric stretch, respectively [43,44]. It is a common notion that bioactive materials can bond to living bone through a new apatite layer that is deposited on their surfaces in contact with body fluid [45,46]. Bone-like hydroxyapatite plays an important role in the formation, growth and maintenance of the bone tissue–biomaterial interface [47]. As multiple in vivo studies have pointed out, bioactive materials were also proved to accelerate bone formation at implantation sites [48–51]. However, CS has been criticized for its no new apatite layer formed in the tissue–biomaterial interface because of its poor bioactivity [19–23], which was also proved by the fact that no apatite was deposited on the CS surface after incubating in SBF in our study. On the contrary, in the present study, the results of SEM and Raman analysis verify that the CSMPC can induce the formation of apatite in SBF within 3 days, indicating that the composite had excellent bioactivity. With the addition of the MPC, PO3− and OH− are released 4 during the dissolution of the composite. Local increases in Ca2+, PO43−, and OH− ions concentration on the CSMPC surface may promote the ionic activity, affect the nucleation of apatite crystals genesis and act as a stimulus to form newly spherical apatite deposition on the composite surface through a dissolution–deposition process. Thus, the CSMPC composite with good bioactivity would form a stronger bond between the implanted material and the surrounding bone tissue compared with the traditional CS material. Fig. 7 shows the weight loss of the CS, CSMPC40, CSMPC50, CSMPC60 and MPC samples after soaking in SBF for various time periods. It can be seen that the CS degraded very fast and the mass of 95.2% dissolved in the SBF for 28 days. While the degradation of the MPC was very slow and only the mass of 19.2% dissolved after 84 days of soaking. The weight losses of the CSMPC composites were remarkably slowed down than that of the CS due to the addition of the MPC. It also can be found that the weight losses of the CSMPC composites became slower with the higher MPC contents. Therefore, it is efficient to tune the degradation of the CSMPC by changing the MPC addition. The changes of pH values of the SBF solution during the sample immersion were shown in Fig. 8. During the SBF immersion, the CS caused the pH of the environment solution to decrease from 7.36 to 6.49 to form an acid environment. The MPC caused slight increase in the pH of the SBF to 7.42 after 15 days of immersion. The pH of the CSMPC60 immersed solution increased continuously and was alkaline around 9.0 after 15 days. It is reported that alkaline environments were helpful to formation of apatite and the nucleation rate of the hydroxyapatite
Fig. 7. Weight losses of the CS, the CSMPC40, the CSMPC50, the CSMPC60 and the MPC samples during soaking in SBF. Statistically significant difference between the degradation rates of CSMPCs and CS (**p b 0.01, *p b 0.05).
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Fig. 8. The changes of pH values of the SBF solution during the sample immersion for 15 days.
crystals is faster at pH 9 than at pH 7 [52]. These could further explain the phenomena that the apatite formed on the MPC and the CSMPC instead of the CS. 3.3. Cell proliferation and cell morphology Fig. 9 shows the proliferation of MG63 cells cultured on the CS, the MPC and the CSMPC60 samples after culturing for 1, 3 and 5 days. The optical density (OD) values from the CCK-8 assay provide an indicator of the cell growth and proliferation on various materials [53,54]. It can be seen that the cell proliferation on all three materials increased with time, which indicates that three materials cause no significant cytotoxicity against cells. Moreover, the proliferation of cell for the CSMPC60 was significantly higher than the CS after incubating for 3 and 5 days; no significant difference appeared after 1 day of culture. The above results indicate that the CSMPC60 can increase the cell viability in comparison to the CS. Fig. 10 shows the SEM micrographs of morphologies of MG63 cells cultured on the CSMPC60 surfaces. After culturing for 4 days (Fig. 10a), the cells firmly attached and spread well on the surface of CSMPC60. After culturing for 7 days as shown in Fig. 10b it was revealed that the cells reached a confluent layer on material surface and had contact
Fig. 9. CCK-8 assay showing the optical density (OD) of the MG63 cells on the CS, the MPC and the CSMPC60 samples after culturing for 1, 3 and 5 days. Significant difference (**p b 0.01).
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Fig. 10. SEM micrographs showing the attachment of the MG63 cells cultured on the CSMPC60 surfaces for (a) 4 days and (b) 7 days.
with each other. Thus, it can be concluded that CSMPC60 did not inhibit the cell proliferation and had a stimulatory effect on cell growth. 4. Conclusions In the present work, a novel bioactive calcium sulfate/magnesium phosphate cement (CSMPC) was fabricated to combine the advantages of the calcium sulfate cement and the magnesium phosphate cement.
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• The prepared CSMPC composites featured the controllable setting time mechanical strength and degradation rate by adjusting the MPC content, which would provide the facility to fulfill the requirement for the clinical application. • The CSMPC composites exhibited good bioactivity to form bone-like apatite in the in vitro SBF test. • The CSMPC composites had good biocompatibility to support the cells to attach and proliferate on the surface. In summary, the CSMPC composite is a promising alternative of bone graft for the bone regeneration. Acknowledgement This work was supported by the Zhejiang Provincial Open Foundation of the Most Important Subjects (2011GK002), Zhejiang Provincial Science and Technology Program of China (No. 2011C23113) and Ningbo Natural Science Fund (2012A610101). References [1] A. Van Heest, M. Swiontkowski, Lancet 353 (Suppl. 1) (1999) SI28–SI29. [2] C.G. Finkemeier, J. Bone Joint Surg. Am. 84-A (2002) 454–464. [3] E. Ahlmann, M. Patzakis, N. Roidis, L. Shepherd, P. Holtom, J. Bone Joint Surg. Am. 84-A (2002) 716–720. [4] G. Carter, AORN J. 70 (1999) 660–670(quiz 672-666). [5] W.W. Tomford, J. Bone Joint Surg. (Am. Vol.) 77 (1995) 1742–1754. [6] P.V. Giannoudis, H. Dinopoulos, E. Tsiridis, Injury 36 (Suppl. 3) (2005) S20–S27. [7] D.M. Strong, G.E. Friedlaender, W.W. Tomford, D.S. Springfield, T.C. Shives, H. Burchardt, W.F. Enneking, H.J. Mankin, Clin. Orthop. Relat. Res. (1996) 107–114. [8] E.M. Ooms, E.A. Egglezos, J.G. Wolke, J.A. Jansen, Biomaterials 24 (2003) 749–757. [9] K. Ohura, M. Bohner, P. Hardouin, J. Lemaitre, G. Pasquier, B. Flautre, J. Biomed. Mater. Res. 30 (1996) 193–200. [10] A. Ogose, T. Hotta, H. Hatano, H. Kawashima, K. Tokunaga, N. Endo, H. Umezu, J. Biomed. Mater. Res. 63 (2002) 601–604. [11] N. Kondo, A. Ogose, K. Tokunaga, T. Ito, K. Arai, N. Kudo, H. Inoue, H. Irie, N. Endo, Biomaterials 26 (2005) 5600–5608. [12] S.K. Nandi, S. Roy, P. Mukherjee, B. Kundu, D.K. De, D. Basu, Indian J. Med. Res. 132 (2010) 15–30.
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