Osteoregenerative capacities of dicalcium phosphate-rich calcium phosphate bone cement Chia-Ling Ko,1,2 Jian-Chih Chen,3 Yin-Chun Tien,3* Chun-Cheng Hung,1 Jen-Chyan Wang,1* Wen-Cheng Chen2 1
College of Dental Medicine, Kaohsiung Medical University; Department of Dentistry, Kaohsiung Medical University Hospital, Kaohsiung 807, Taiwan 2 Department of Fiber and Composite Materials, Feng Chia University, Taichung 407, Taiwan 3 College of Medicine, Kaohsiung Medical University; Department of Orthopaedics, Kaohsiung Medical University Hospital, Kaohsiung 807, Taiwan Received 25 November 2013; revised 27 February 2014; accepted 12 March 2014 Published online 00 Month 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35167 Abstract: Calcium phosphate cement (CPC) is a widely used bone substitute. However, CPC application is limited by poor bioresorption, which is attributed to apatite, the stable product. This study aims to systematically survey the biological performance of dicalcium phosphate (DCP)-rich CPC. DCPrich CPC exhibited a twofold, surface-modified DCP anhydrous (DCPA)-to-tetracalcium phosphate (TTCP) molar ratio, whereas conventional CPC (c-CPC) showed a onefold, surface unmodified DCPA-to-TTCP molar ratio. Cell adhesion, morphology, viability, and alkaline phosphatase (ALP) activity in the two CPCs were examined with bone cell progenitor D1 cultured in vitro. Microcomputed tomography and histological observation were conducted after CPC implantation in vivo to analyze the residual implant ratio and new bone
formation rate. D1 cells cultured on DCP-rich CPC surfaces exhibited higher cell viability, ALP activity, and ALP quantity than c-CPC. Histological evaluation indicated that DCP-rich CPC showed lesser residual implant and higher new bone formation rate than c-CPC. Therefore, DCP-rich CPC can improve bioresorption. The newly developed DCP-rich CPC exhibited potential therapeutic applications for bone reconC 2014 Wiley Periodicals, Inc. J Biomed Mater Res Part A: struction. V 00A:000–000, 2014.
Key Words: calcium phosphate cements, bioresorption, nanocrystallites, dicalcium phosphate anhydrous, dicalcium phosphate dihydrate
How to cite this article: Ko C-L, Chen J-C, Tien Y-C, Hung C-C, Wang J-C, Chen W-C. 2014. Osteoregenerative capacities of dicalcium phosphate-rich calcium phosphate bone cement. J Biomed Mater Res Part A 2014:00A:000–000.
INTRODUCTION
Calcium phosphate cements (CPCs) have been used as bone substitute materials because of their superior biocompatibility and osteoconductivity. Conventional CPC (c-CPC) is the known general formula of CPC prepared by mixing of hardening solution with powder consisting of an equimolar mixture of tetracalcium phosphate [TTCP, Ca4(PO4)2O] and dicalcium phosphate anhydrous (DCPA, CaHPO4, or DCPD and CaHPO42H2O).1 The hardened c-CPC produces a poor crystalline hydroxyapatite (HA) phase.1,2 which cannot be extensively resorbed and replaced by bone tissue.3 To activate the interactions between bone cells and materials and enhance the resorption rate of CPCs, incorporation of macroporous fabricating additives, such as water-soluble polymers, growth factors, and collagens, in apatite-forming CPCs is proposed.4–6 However, these additives may affect the physiochemical property of c-CPC, such as setting and working durations, viscosity, strength, toughness, and dispersibility.
Nevertheless, the factors that critically affect CPC bioresorption are associated with the solubility of their constitutive phases.3 Accordingly, the dissolution rate increased with the decrease in TTCP/DCPA molar ratio from 1/1 to 1/3 because the ion release rate was increased. The ions released from CPCs caused ion–dipole interaction and remineralized decomposition, a result indicating enhancement of cell differentiation into osteoblasts7 and stimulation of osteogenesis.8 This occurrence suggests that fast CPC resorption by biological response can be achieved with a low Ca/P formula in the CPC original reactants. Decreasing the TTCP/DCPA molar ratio can reduce the pH after reaction and thus lead to low biocompatibility.9 Increasing the DCPA molar ratio to enhance the resorption rate of CPCs accompanies the damage of mechanical property. Therefore, maintaining the mechanical property of CPCs through an increase in the DCPA molar ratio can significantly improve the clinical application of CPCs.
*These authors contributed equally to this work. Correspondence to: W.-C. Chen; e-mail: [email protected] or [email protected] Contract grant sponsor: National Science Council of the Executive Yuan, Taiwan; contract grant number: NSC 97–2221-E-037-006 and 102–2622E-035-018-CC2
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Nanocrystal pretreatment on the reactant surfaces of CPCs results in significant particle interlocking ability, particularly in medium-immersed situations.10–12 Our previous research11 indicated that 1/2 is the optimal experimental molar ratio of TTCP to surface-modified DCPA. However, the effects of this DCP-rich CPC characteristic on cell interactions in vitro and the improvement in bioresorption rate in vivo remain undetermined. This study aims to systematically investigate osteogenic abilities in interaction with the bone cell progenitor (D1), as well as the histological properties of DCP-rich CPC and c-CPC. The null hypothesis is that DCP-rich CPC exerts more positive effects on bioresorption in vivo than c-CPC does. MATERIALS AND METHODS
Specimen preparation TTCP powder was prepared by reaction, as suggested by Brown and Epstein.13 Nanocrystal pretreatment on DCPA was performed according to our modified method.11–13 Briefly, 5 g of DCPA (CaHPO4; Acros Organics, Geel, Belgium) and 8 g of TTCP were added to dilute acidic aqueous phosphate solution (25 mM, pH 1.96) and initiate the reaction in 20 min at a pH between 5 and 7. Particles were vacuumfiltered, washed with deionized water, and dried in an oven. DCP-rich CPC was prepared by mixing of the surfacemodified DCPA and TTCP (DCPA/TTCP molar ratio 5 2). cCPC (surface unmodified DCPA/TTCP molar ratio 5 1) with untreated DCPA was used for comparison. The hardening solution (pH 7.4) was a neutral sodium hydrogen phosphate solution (1 M) with a powder-to-liquid ratio of 3.0 g/mL. The CPC powders and hardening solution were g-ray sterilized at 15 kGy (China Biotech Co., Taiwan). For in vitro testing, the specimens were mixed for 2 min, and then the paste was molded (diameter 3 depth, 6 mm 3 3 mm) under 0.7 MPa pressure for 8 min. The cement was demolded after being mixed for 15 min. To enhance the mechanical strength of c-CPC, the demolded specimens were placed under 100% humidity at 37 C after they were mixed. This set-up was maintained for 24 h. DCP-rich CPC specimens were prepared according to the procedures used in c-CPC preparation, except that the DCP-rich CPC was not placed under 100% humidity at 37 C for 24 h and it was directly immersed in culture solution. In vitro testing D1 cells are a bone marrow mesenchymal stem cell (MSC) line cloned from Balb/C mice. These cells were purchased from the American Type Culture Collection and cultured in Dulbecco’s modified Eagle’s medium supplemented with 10% fetal bovine serum at 37 C under a humidified 5% CO2 atmosphere. The D1 cells were used before their eighth passage. Testing was performed at time intervals of 1, 2, 4, 8, and 16 days after initial seeding of 1 3 105 D1 cells on the specimen surface. AlamarBlueV assay (AbD Serotec) and p-nitrophenyl phosphate kits (Sigma-Aldrich) were used to determine cell viability and production of alkaline phosphatase (ALP). Each experiment was performed in triplicate (n 5 3). ALP staining was performed with the use of serum tartrate-resistant acid R
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phosphatase and ALP double-stain kit (Takara Bio, Shiga, Japan). ALP-stained tested samples were washed prior to gross examination under a light microscope. In vivo testing The animal testing procedures used in this study were approved by the Institutional Animal Care and Use Committee of Kaohsiung Medical University. National Institutes of Health (NIH) guidelines for the care and use of laboratory animals (NIH Publication #85-23 Rev. 1985) were observed. A total of 12 rabbits were randomly divided into four groups on the basis of the following implantation periods: 3, 6, 9, and 24 weeks. The c-CPC and the DCP-rich CPC were implanted in the left and right femurs, respectively. The surgical procedure was the same as that in our previous study.14,15 Prior to setting, the mixing powders with hardening solution were stirred for 2 min. The paste was allowed to form for 1 min and then loaded into a 1-mL syringe for injection into the prepared bone cavity (4 mm 3 5 mm, diameter 3 depth). At 3, 6, 9, and 24 weeks after surgery, the rabbits were sacrificed, and the femur portions were immediately excised. The femurs were scanned at 35 mm intervals at 89 kV and 100 mA by microcomputed tomography scanner analysis (SkyScan1076; SkyScan, Kontich, Belgium). The residual ratios of the implants were calculated with CTAn Software (SkyScan). The retrieved bone was sectioned into three 1.5 mm-thick slices from the lateral to the medial parts to evaluate the resorption of the implants. The middle slice was used to measure the total new bone area (BA) ratio. Optical images of the unstained slices were analyzed with an automatic image analysis system (Matrox Inspector, Germany) to determine the residual ratios of the implants. The total BA (in %) was determined according to the following modified equation:14 Total bone ratio ðBA; in %Þ 5 ½1 2 ðcross-sectional residual area of the implantÞ=
(1)
ðcross-sectional area of the original implantÞ 3 100%:
The sectioned bones were thinned out to a final thickness of 60 lm, polished, and glued to slides with Permount (Fisher Scientific, Fair Lawn, NJ). These sections were stained with hematoxylin–eosin (H&E). Histology was observed under an optical microscope (BX51; OLYMPUS, Tokyo, Japan). Statistical analysis Statistical significance was evaluated by one-way analysis of variance (ANOVA). Significant differences were determined by Student’s t-test and two-way ANOVA (SAS Institute, Cary, NC). In all cases, p < 0.05 was considered statistically significant. RESULTS
Cell attachment Excellent biomaterials result in early cell attachment on the material surface and thus affect cell proliferation and mineralization. Figure 1 shows the morphologies of D1 cells cultured on specimen surfaces after 1 h, 1 day, and 2 days. After 1 h, the D1 cells on both CPC samples extended their filopodia, which are thin, actin-rich, plasma-membrane
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FIGURE 1. Typical morphologies of different CPC specimens with fixed D1 cells, as evaluated by SEM.
protrusions that function as antennae for cells.16 After 1 and 2 days of culture, cells with long cytoplasmic extensions were observed. These overlapping cells were visible in the c-CPC and DCP-rich CPC samples. The cell extensions are the regions of the cell plasma membrane containing a meshwork or bundles of actin-containing microfilaments, which allow migrating cells to move along a substratum.17 Overlapping cells became distinct on the DCP-rich CPC after 1 day of culture. This finding suggests that the D1 cells showed higher proliferation on the DCP-rich CPC than on cCPC. Cell viabilities after different culture time periods Figure 2 shows the cell viability on CPCs. Cell viability was higher on DCP-rich CPC than on c-CPC in the different
FIGURE 2. Cell viability of D1 cells seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary. com.]
culture time periods. The cell viability on c-CPC significantly increased after 2 and 4 days of culture and was maintained until day 8. Cell viability on DCP-rich CPC was slightly higher from day 2 to day 4 than that on c-CPC, but the difference was not significant. Cell viability on DCP-rich CPC remained almost the same from days 2 to 8. ALP quantification and staining ALP activity is an early-stage marker relevant in progenitor D1 cell differentiation.18 Because cell viability is directly associated with cell number, the quantified values shown in Figure 3 indicate the ALP activity per cell. These values were obtained by division of the absorbance values of the
FIGURE 3. OD and ALP activity of each D1 cell seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5, symbol * means at the 0.05 level, the difference of the population was significantly different). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 4. ALP staining of D1 cells seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
ALP measurements by the respective values of cell viability. The absorbance value of the ALP measurements and the quantified ALP activity per cell cultured on c-CPC and DCPrich CPC surfaces were the same. Maximum values were all reached at 8 days of incubation. The quantified ALP activity per D1 cell in DCP-rich CPC was markedly higher than that in c-CPC, and the quantity significantly increased as the incubation time was prolonged to 8 days (Fig. 4). DCP-rich CPC showed darker blue staining than c-CPC from days 2 to
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FIGURE 5. Microcomputed tomography views of rabbit tibiae after 3, 6, 9, and 24 weeks of implantation of different CPC specimens. The values indicated at the bottom of the image represent residual implant areas, expressed in %.
8. Bone cell progenitor D1 seeded on the DCP-rich CPC surface initiated at an earlier stage of differentiation than on the c-CPC surface. Histological observations The residual implant area of DCP-rich CPC was significantly less than that of c-CPC at 3, 9, and 24 weeks after implantation (Fig. 5). The cross sections of the CPC implant sites indicated an increase in new BA (in %) over time [Fig. 6(a,b)]. The BA% of the DCP-rich CPC group was
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significantly higher than that of the c-CPC group even at 24 weeks (p < 0.05) postoperation. Regression between the DCP-rich effect of CPC composition and implantation time was reflected on the residual implant rate [Fig. 7(a)] and new bone formation rate [Fig. 7(b)]. The linear relationship of the residual implant rate was exhibited only in the DCPrich CPC and c-CPC implanted before 9 and 6 weeks, respectively. Compared with the residual implant rate, the linear regression between the new bone formation rate and implantation time was reflected in the DCP-rich CPC but not in the c-CPC. Figure 8 shows the typical histological morphologies of nondecalcified H&E-stained sections. Both CPC groups showed good tissue contact postimplantation, as indicated by the absence of inflammatory or other immune responses. DCP-rich CPC exhibited good bone remodeling, which increased the formation of new trabecular bone tissue around the implant sites (Fig. 8) over time. This increase was more prominent in DCP-rich CPC than in cCPC graft material at 9 and 24 weeks. DISCUSSION
The significant differences between DCP-rich CPC and c-CPC indicated that the null hypothesis of this study is consistent with the results. The cell viability, ALP activity in vitro, and bioresorption rate in vivo of DCP-rich CPC were higher than those of c-CPC. The early D1 cell behaviors of primary adhesion on the surfaces showed the quantity of surface cell attachment and influenced the succeeding mechanisms by spread morphology, proliferation, and differentiation.19 The chemical composition and ionic solubility of the biomaterial also significantly affect the interaction between cells.20,21 CPCs can release Ca21 and PO432 ions in the physiological environment. These ions attract cells to the material surface,
FIGURE 6. (a) Cross-sectional images of implantation with different bone cements showing color contrast images (orange for natural bone; gray for residual CPCs). (b) Total bone areas of different materials implanted in a rabbit femur for different lengths of time (n 5 3, symbol * means at the 0.05 level, the difference of the population was significantly different). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
FIGURE 7. Regression analysis based on the curves of (a) residual implant in ratios and (b) bone area (%) for different time periods. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 8. Nondecalcified histological analysis of the medial condyle of the rabbit femur. Representative H&E-stained sections of different CPC cements after 3, 6, 9, and 24 weeks of implantation. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary. com.]
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FIGURE 9. Schematic of the biological mechanisms of bone cement, such as ion release, cell adhesion, cell proliferation, cell differentiation, and new bone regeneration.
which is modulated by physicochemical interactions at the beginning of cell seeding, as well as adsorption of fibronectin and vitronectin from the serum.18,22 Sogo et al.23 observed a significantly higher ALP activity in osteogenically induced human MSC cultured on a fibronectin-coated HA ceramic than on an untreated material. Figure 9 shows the summary of the simulated biological mechanisms of bone cement with ion release and the consequences of cell adhesion, proliferation, and differentiation, as well as new bone regeneration. Decreasing the TTCP/DCPA molar ratio from the traditional ratio of 1 to 0.33 increased the CPC dissolution rate by 40%.24 When CPC reduced the TTCP/DCPA molar ratio from 1 to 0.33, the human umbilical cord MSCs showed excellent viability and proliferation, as well as high ALP activity, osteocalcin content, collagen I content, and osterix gene expression.24 Adjusting CPC composition to enhance Ca21 and PO432 ion release can promote adhesion of many growth factors and proteins to the surface of the implant substances, as well as the activity and gene expression of osteoblasts.25 This study adjusted the surface-modified DCPA/TTCP molar ratio to 2/1 to increase the CPC dissolution rate. The change in molar ratio of DCP-rich CPC resulted in higher cell viability, ALP activity, and ALP quality than c-CPC, as observed in the D1 cell cultures (Figs. 1–4). D1 cells cultured on the DCP-rich CPC showed limited proliferation, which induced early biological transition from a proliferating progenitor to a differentiated postmitotic cell. After implantation and contact with body fluids, the DCPA of DCPD-forming CPC converted to DCPD and showed the primary dissolution of the cement.26 A linear and rapid resorption rate, which was observed in vivo (20 lm/day) in the DCPD-forming CPC after implantation, was similar to the speed of bone growth.27 However, the DCPD resorption was too quick for tissue regeneration and may cause a small gap for soft tissue growth between the implant and the newly formed bones.28 To achieve an ideal CPC resorption
rate, a technique was investigated to increase the molar ratio of DCPA and thus enhance the formation of product phases of DCPAs in the DCP-rich CPC. The residual implant rate indicated that a twofold DCPA molar ratio showed a deep resorption slope in early implantation, and resorption/ mineralization periods lasted for 24 weeks compared with c-CPC. The new bone formation rate of DCP-rich CPC continued to increase up to 24 weeks of implantation compared with that of c-CPC. The varied c-CPC showed a dt trend after 9 weeks. The new mechanism of bone formation in DCP-rich CPC continued for 24 weeks, but the new bone formation rate of c-CPC was almost the same during 9–24 weeks of implantation (Figs. 5 and 6). DCP-rich CPC and cCPC showed a linear regression for new bone formation and degradation rates before 9 and 6 weeks of implantation, respectively. In contrast to DCPD, apatite degraded very slowly because the bioresorb mechanism was solely facilitated by osteoclast activity rather than macrophage phagocytosis.29,30 After CPC implantation, the central area of the implanted DCP-rich CPC was converted into apatite over time. This phase was difficult to dissolve by osteoclast activity; thus, the linear relationship of bioresorption rate is not shown (Fig. 7). These findings indicate that DCP-rich CPC had a better dissolution rate than c-CPC in releasing large quantities of Ca21 and PO432 ions. Large amounts of surface-modified DCP formed a Ca21- and PO432-rich environment, which improved protein adhesion and enhanced biological performance on the specimen after implantation, as illustrated in Figure 9. The results of the in vitro test were consistent with those of the in vivo test. DCP-rich CPC performed better than c-CPC regardless of the complex interactions between different cell types and the implanted materials in the biological mechanisms in the in vivo test. In vitro test results may confirm the conclusions derived in vivo. In vitro analyses of cellular behavior, such as cell adhesion and proliferation, as well as mineralization on CPC substances in an
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in vitro environment, are low cost and ethical. In vitro examinations can help avoid controversial issues on the use of animals in experiments. New materials developed at an early stage can be evaluated with in vitro techniques. CONCLUSION
The formula of DCP-rich CPC contains a twofold higher surface-modified DCPA-to-TTCP molar ratio than c-CPC. DCPrich CPC showed improved cell attachment, viability, and ALP activity compared with c-CPC in vitro. Histological evaluation in vivo also confirmed that DCP-rich CPC exhibited higher bone regeneration efficiency than c-CPC. DCP-rich CPC dissolved faster than c-CPC, which resulted in a Ca21 and PO432 ion-rich environment. The enhanced protein adhesion on specimen surfaces and biological performance can be attributed to the large amount of ions in the environment. Improving bioresorption by adjustment of the composition of DCP-rich CPC can lead to high osteoconductivity and excellent biocompatibility for application in bone regeneration. ACKNOWLEDGMENTS
The authors sincerely acknowledge the assistance of Ms. HuiYu Wu and Ya-Shun Chen in this research. REFERENCES 1. Brown WE, Chow LC. A new calcium phosphate water setting cement. In: Brown PW, editor. Cements Research Progress. Westerville, OH: Am Ceram Soc; 1986. p 352–379. 2. Friedman CD, Costantino PD, Takagi S, Chow LC. BoneSourceTM hydroxyapatite cement: A novel biomaterial for craniofacial skeletal tissue engineering and reconstruction. J Biomed Mater Res 1998;43:428–432. 3. Dorozhkin SV. Calcium orthophosphate cements and concretes. Materials 2009;2:221–291. 4. Almirall A, Larrecq G, Delgado JA, Martinez S, Planell JA, Ginebra MP. Fabrication of low temperature macroporous hydroxyapatite scaffolds by foaming and hydrolysis of an alphaTCP paste. Biomaterials 2004;25:3671–3680. 5. Sun L, Xu HH, Takagi S, Chow LC. Fast setting calcium phosphate cement–chitosan composite: Mechanical properties and dissolution rates. J Biomater Appl 2007;21:299–315. 6. Chen JC, Ko CL, Shih CJ, Tien YC, Chen WC. Calcium phosphate bone cement with 10 wt% platelet-rich plasma in vitro and in vivo. J Dent 2012;40:114–122. 7. Perez RA, Ginebra MP, Spector M. Cell response to collagencalcium phosphate cement scaffolds investigated for nonviral gene delivery. J Mater Sci Mater Med 2011;22:887–897. 8. Tsigkou O, Jones JR, Polak JM, Stevens MM. Differentiation of fetal osteoblasts and formation of mineralized bone nodules by V 45S5 Bioglass conditioned medium in the absence of osteogenic supplements. Biomaterials 2009;30: 3542–3550. 9. Hirayama S, Takagi S, Markovic M, Chow LC. Properties of calcium phosphate cements with different tetracalcium phosphate and dicalcium phosphate anhydrous molar ratios. J Res Natl Inst Stand Technol 2008;113:311–320. 10. Wang JC, Ko CL, Hung CC, Tyan YC, Lai CH, Chen WC, Wang CK. Deriving fast setting properties of tetracalcium phosphate/dicalcium phosphate anhydrous bone cement with nanocrystallites on the reactant surfaces. J Dent 2010;38:158–165. 11. Ko CL, Chen JC, Tien YC, Hung CC, Wang JC, Chen WC. Biphasic products of dicalcium phosphate-rich cement with injectability and nondispersibility. Mater Sci Eng C Mater Biol Appl 2014;39: 40–46. doi: 10.1016/j.msec.2014.02.033. R
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12. Chen WC, Hung CC, Ko CL. Calcium phosphate bone cement, precursor thereof and fabrication method thereof. US Patent Application No: 2010/0313,791; 2010. 13. Brown WE, Epstein EF. Crystallography of tetracalcium phosphate. J Res Natl Bur Stand 1965;69A:547–551. 14. Chen WC, Ju CP, Tien YC, Chern Lin JH. In vivo testing of nanoparticle-treated TTCP/DCPA-based ceramic surfaces. Acta Biomater 2009;5:1767–1774. 15. Ko CL, Chen WC, Chen JC, Wu HY, Shih CJ, Tien YC, Hung CC, Wang JC. Properties of osteoconductive biomaterials: Calcium phosphate cement with different ratios of platelet-rich plasma as identifiers. Mater Sci Eng C Mater Biol Appl 2013;33:3537–3544. 16. Mattila PK, Lappalainen P. Filopodia: Molecular architecture and cellular functions. Nat Rev Mol Cell Biol 2008;9:446–454. 17. Chen W, Zhou H, Weir MD, Tang M, Bao C, Xu HH. Human embryonic stem cell-derived mesenchymal stem cell seeding on calcium phosphate cement-chitosan-RGD scaffold for bone repair. Tissue Eng Part A 2013;19:915–927. 18. Bernhardt A, Lode A, Peters F, Gelinsky M. Comparative evaluation of different calcium phosphate-based bone graft granules— An in vitro study with osteoblast-like cells. Clin Oral Implants Res 2013;24:441–449. 19. Mandal BB, Das S, Choudhury K, Kundu SC. Implication of silk film RGD availability and surface roughness on cytoskeletal organization and proliferation of primary rat bone marrow cells. Tissue Eng Part A 2010;16:2391–2403. 20. Wu C, Ramaswamy Y, Zhu Y, Zheng R, Appleyard R, Howard A, Zreiqat H. The effect of mesoporous bioactive glass on the physiochemical, biological and drug-release properties of poly( D , L -lactide-co-glycolide) films. Biomaterials 2009;30: 2199–2208. 21. Varanasi VG, Saiz E, Loomer PM, Ancheta B, Uritani N, Ho SP, Tomsia AP, Marshall SJ, Marshall GW. Enhanced osteocalcin expression by osteoblast-like cells (MC3T3-E1) exposed to bioactive coating glass (SiO2–CaO–P2O5–MgO–K2O–Na2O system) ions. Acta Biomater 2009;5:3536–3547. 22. Anselme K, Bigerelle M, Noel B. Dufresne E. Judas D, Iost A, Hardouin P, Qualitative and quantitative study of human osteoblast adhesion on materials with various surface roughnesses. J Biomed Mater Res 2000;49:155–166. 23. Sogo Y, Ito A, Matsuno T, Oyane A, Tamazawa G, Satoh T, Yamazaki A, Uchimura E, Ohno T. Fibronectin-calcium phosphate composite layer on hydroxyapatite to enhance adhesion, cell spread and osteogenic differentiation of human mesenchymal stem cells in vitro. Biomed Mater 2007;2:116–123. 24. Xu HH, Zhao L, Detamore MS, Takagi S, Chow LC. Umbilical cord stem cell seeding on fast-resorbable calcium phosphate bone cement. Tissue Eng Part A 2010;16:2743–2753. 25. Chow LC. Next generation calcium phosphate-based biomaterials. Dent Mater J 2009;28:1–10. 26. Bohner M, Theiss F, Apelt D, Hirsiger W, Houriet R, Rizzoli G, Gnos E, Frei C, Auer JA, von Rechenberg B. Compositional changes of a dicalcium phosphate dihydrate cement after implantation in sheep. Biomaterials 2003;24:3463–3474. 27. Bohner M, Gbureck U, Barralet JE. Technological issues for the development of more efficient calcium phosphate bone cements: A critical assessment. Biomaterials 2005;26:6423–6429. 28. Ohura K, Bohner M, Hardouin P, Lemaitre J, Pasquier G, Flautre B. Resorption of, and bone formation from, new beta-tricalcium phosphate-monocalcium phosphate cements: An in vivo study. J Biomed Mater Res 1996;30:193–200. € rbs T, 29. Schneider G, Blechschmidt K, Linde D, Litschko P, Ko Beleites E. Bone regeneration with glass ceramic implants and calcium phosphate cements in a rabbit cranial defect model. J Mater Sci Mater Med 2010;21:2853–2859. 30. Grossardt C, Ewald A, Grover LM, Barralet JE, Gbureck U. Passive and active in vitro resorption of calcium and magnesium phosphate cements by osteoclastic cells. Tissue Eng Part A 2010;16: 3687–3695.
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College of Dental Medicine, Kaohsiung Medical University; Department of Dentistry, Kaohsiung Medical University Hospital, Kaohsiung 807, Taiwan 2 Department of Fiber and Composite Materials, Feng Chia University, Taichung 407, Taiwan 3 College of Medicine, Kaohsiung Medical University; Department of Orthopaedics, Kaohsiung Medical University Hospital, Kaohsiung 807, Taiwan Received 25 November 2013; revised 27 February 2014; accepted 12 March 2014 Published online 00 Month 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35167 Abstract: Calcium phosphate cement (CPC) is a widely used bone substitute. However, CPC application is limited by poor bioresorption, which is attributed to apatite, the stable product. This study aims to systematically survey the biological performance of dicalcium phosphate (DCP)-rich CPC. DCPrich CPC exhibited a twofold, surface-modified DCP anhydrous (DCPA)-to-tetracalcium phosphate (TTCP) molar ratio, whereas conventional CPC (c-CPC) showed a onefold, surface unmodified DCPA-to-TTCP molar ratio. Cell adhesion, morphology, viability, and alkaline phosphatase (ALP) activity in the two CPCs were examined with bone cell progenitor D1 cultured in vitro. Microcomputed tomography and histological observation were conducted after CPC implantation in vivo to analyze the residual implant ratio and new bone
formation rate. D1 cells cultured on DCP-rich CPC surfaces exhibited higher cell viability, ALP activity, and ALP quantity than c-CPC. Histological evaluation indicated that DCP-rich CPC showed lesser residual implant and higher new bone formation rate than c-CPC. Therefore, DCP-rich CPC can improve bioresorption. The newly developed DCP-rich CPC exhibited potential therapeutic applications for bone reconC 2014 Wiley Periodicals, Inc. J Biomed Mater Res Part A: struction. V 00A:000–000, 2014.
Key Words: calcium phosphate cements, bioresorption, nanocrystallites, dicalcium phosphate anhydrous, dicalcium phosphate dihydrate
How to cite this article: Ko C-L, Chen J-C, Tien Y-C, Hung C-C, Wang J-C, Chen W-C. 2014. Osteoregenerative capacities of dicalcium phosphate-rich calcium phosphate bone cement. J Biomed Mater Res Part A 2014:00A:000–000.
INTRODUCTION
Calcium phosphate cements (CPCs) have been used as bone substitute materials because of their superior biocompatibility and osteoconductivity. Conventional CPC (c-CPC) is the known general formula of CPC prepared by mixing of hardening solution with powder consisting of an equimolar mixture of tetracalcium phosphate [TTCP, Ca4(PO4)2O] and dicalcium phosphate anhydrous (DCPA, CaHPO4, or DCPD and CaHPO42H2O).1 The hardened c-CPC produces a poor crystalline hydroxyapatite (HA) phase.1,2 which cannot be extensively resorbed and replaced by bone tissue.3 To activate the interactions between bone cells and materials and enhance the resorption rate of CPCs, incorporation of macroporous fabricating additives, such as water-soluble polymers, growth factors, and collagens, in apatite-forming CPCs is proposed.4–6 However, these additives may affect the physiochemical property of c-CPC, such as setting and working durations, viscosity, strength, toughness, and dispersibility.
Nevertheless, the factors that critically affect CPC bioresorption are associated with the solubility of their constitutive phases.3 Accordingly, the dissolution rate increased with the decrease in TTCP/DCPA molar ratio from 1/1 to 1/3 because the ion release rate was increased. The ions released from CPCs caused ion–dipole interaction and remineralized decomposition, a result indicating enhancement of cell differentiation into osteoblasts7 and stimulation of osteogenesis.8 This occurrence suggests that fast CPC resorption by biological response can be achieved with a low Ca/P formula in the CPC original reactants. Decreasing the TTCP/DCPA molar ratio can reduce the pH after reaction and thus lead to low biocompatibility.9 Increasing the DCPA molar ratio to enhance the resorption rate of CPCs accompanies the damage of mechanical property. Therefore, maintaining the mechanical property of CPCs through an increase in the DCPA molar ratio can significantly improve the clinical application of CPCs.
*These authors contributed equally to this work. Correspondence to: W.-C. Chen; e-mail: [email protected] or [email protected] Contract grant sponsor: National Science Council of the Executive Yuan, Taiwan; contract grant number: NSC 97–2221-E-037-006 and 102–2622E-035-018-CC2
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Nanocrystal pretreatment on the reactant surfaces of CPCs results in significant particle interlocking ability, particularly in medium-immersed situations.10–12 Our previous research11 indicated that 1/2 is the optimal experimental molar ratio of TTCP to surface-modified DCPA. However, the effects of this DCP-rich CPC characteristic on cell interactions in vitro and the improvement in bioresorption rate in vivo remain undetermined. This study aims to systematically investigate osteogenic abilities in interaction with the bone cell progenitor (D1), as well as the histological properties of DCP-rich CPC and c-CPC. The null hypothesis is that DCP-rich CPC exerts more positive effects on bioresorption in vivo than c-CPC does. MATERIALS AND METHODS
Specimen preparation TTCP powder was prepared by reaction, as suggested by Brown and Epstein.13 Nanocrystal pretreatment on DCPA was performed according to our modified method.11–13 Briefly, 5 g of DCPA (CaHPO4; Acros Organics, Geel, Belgium) and 8 g of TTCP were added to dilute acidic aqueous phosphate solution (25 mM, pH 1.96) and initiate the reaction in 20 min at a pH between 5 and 7. Particles were vacuumfiltered, washed with deionized water, and dried in an oven. DCP-rich CPC was prepared by mixing of the surfacemodified DCPA and TTCP (DCPA/TTCP molar ratio 5 2). cCPC (surface unmodified DCPA/TTCP molar ratio 5 1) with untreated DCPA was used for comparison. The hardening solution (pH 7.4) was a neutral sodium hydrogen phosphate solution (1 M) with a powder-to-liquid ratio of 3.0 g/mL. The CPC powders and hardening solution were g-ray sterilized at 15 kGy (China Biotech Co., Taiwan). For in vitro testing, the specimens were mixed for 2 min, and then the paste was molded (diameter 3 depth, 6 mm 3 3 mm) under 0.7 MPa pressure for 8 min. The cement was demolded after being mixed for 15 min. To enhance the mechanical strength of c-CPC, the demolded specimens were placed under 100% humidity at 37 C after they were mixed. This set-up was maintained for 24 h. DCP-rich CPC specimens were prepared according to the procedures used in c-CPC preparation, except that the DCP-rich CPC was not placed under 100% humidity at 37 C for 24 h and it was directly immersed in culture solution. In vitro testing D1 cells are a bone marrow mesenchymal stem cell (MSC) line cloned from Balb/C mice. These cells were purchased from the American Type Culture Collection and cultured in Dulbecco’s modified Eagle’s medium supplemented with 10% fetal bovine serum at 37 C under a humidified 5% CO2 atmosphere. The D1 cells were used before their eighth passage. Testing was performed at time intervals of 1, 2, 4, 8, and 16 days after initial seeding of 1 3 105 D1 cells on the specimen surface. AlamarBlueV assay (AbD Serotec) and p-nitrophenyl phosphate kits (Sigma-Aldrich) were used to determine cell viability and production of alkaline phosphatase (ALP). Each experiment was performed in triplicate (n 5 3). ALP staining was performed with the use of serum tartrate-resistant acid R
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phosphatase and ALP double-stain kit (Takara Bio, Shiga, Japan). ALP-stained tested samples were washed prior to gross examination under a light microscope. In vivo testing The animal testing procedures used in this study were approved by the Institutional Animal Care and Use Committee of Kaohsiung Medical University. National Institutes of Health (NIH) guidelines for the care and use of laboratory animals (NIH Publication #85-23 Rev. 1985) were observed. A total of 12 rabbits were randomly divided into four groups on the basis of the following implantation periods: 3, 6, 9, and 24 weeks. The c-CPC and the DCP-rich CPC were implanted in the left and right femurs, respectively. The surgical procedure was the same as that in our previous study.14,15 Prior to setting, the mixing powders with hardening solution were stirred for 2 min. The paste was allowed to form for 1 min and then loaded into a 1-mL syringe for injection into the prepared bone cavity (4 mm 3 5 mm, diameter 3 depth). At 3, 6, 9, and 24 weeks after surgery, the rabbits were sacrificed, and the femur portions were immediately excised. The femurs were scanned at 35 mm intervals at 89 kV and 100 mA by microcomputed tomography scanner analysis (SkyScan1076; SkyScan, Kontich, Belgium). The residual ratios of the implants were calculated with CTAn Software (SkyScan). The retrieved bone was sectioned into three 1.5 mm-thick slices from the lateral to the medial parts to evaluate the resorption of the implants. The middle slice was used to measure the total new bone area (BA) ratio. Optical images of the unstained slices were analyzed with an automatic image analysis system (Matrox Inspector, Germany) to determine the residual ratios of the implants. The total BA (in %) was determined according to the following modified equation:14 Total bone ratio ðBA; in %Þ 5 ½1 2 ðcross-sectional residual area of the implantÞ=
(1)
ðcross-sectional area of the original implantÞ 3 100%:
The sectioned bones were thinned out to a final thickness of 60 lm, polished, and glued to slides with Permount (Fisher Scientific, Fair Lawn, NJ). These sections were stained with hematoxylin–eosin (H&E). Histology was observed under an optical microscope (BX51; OLYMPUS, Tokyo, Japan). Statistical analysis Statistical significance was evaluated by one-way analysis of variance (ANOVA). Significant differences were determined by Student’s t-test and two-way ANOVA (SAS Institute, Cary, NC). In all cases, p < 0.05 was considered statistically significant. RESULTS
Cell attachment Excellent biomaterials result in early cell attachment on the material surface and thus affect cell proliferation and mineralization. Figure 1 shows the morphologies of D1 cells cultured on specimen surfaces after 1 h, 1 day, and 2 days. After 1 h, the D1 cells on both CPC samples extended their filopodia, which are thin, actin-rich, plasma-membrane
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FIGURE 1. Typical morphologies of different CPC specimens with fixed D1 cells, as evaluated by SEM.
protrusions that function as antennae for cells.16 After 1 and 2 days of culture, cells with long cytoplasmic extensions were observed. These overlapping cells were visible in the c-CPC and DCP-rich CPC samples. The cell extensions are the regions of the cell plasma membrane containing a meshwork or bundles of actin-containing microfilaments, which allow migrating cells to move along a substratum.17 Overlapping cells became distinct on the DCP-rich CPC after 1 day of culture. This finding suggests that the D1 cells showed higher proliferation on the DCP-rich CPC than on cCPC. Cell viabilities after different culture time periods Figure 2 shows the cell viability on CPCs. Cell viability was higher on DCP-rich CPC than on c-CPC in the different
FIGURE 2. Cell viability of D1 cells seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary. com.]
culture time periods. The cell viability on c-CPC significantly increased after 2 and 4 days of culture and was maintained until day 8. Cell viability on DCP-rich CPC was slightly higher from day 2 to day 4 than that on c-CPC, but the difference was not significant. Cell viability on DCP-rich CPC remained almost the same from days 2 to 8. ALP quantification and staining ALP activity is an early-stage marker relevant in progenitor D1 cell differentiation.18 Because cell viability is directly associated with cell number, the quantified values shown in Figure 3 indicate the ALP activity per cell. These values were obtained by division of the absorbance values of the
FIGURE 3. OD and ALP activity of each D1 cell seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5, symbol * means at the 0.05 level, the difference of the population was significantly different). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 4. ALP staining of D1 cells seeded on different CPC specimens after 1, 4, 8, and16 days of incubation (n 5 5). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
ALP measurements by the respective values of cell viability. The absorbance value of the ALP measurements and the quantified ALP activity per cell cultured on c-CPC and DCPrich CPC surfaces were the same. Maximum values were all reached at 8 days of incubation. The quantified ALP activity per D1 cell in DCP-rich CPC was markedly higher than that in c-CPC, and the quantity significantly increased as the incubation time was prolonged to 8 days (Fig. 4). DCP-rich CPC showed darker blue staining than c-CPC from days 2 to
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FIGURE 5. Microcomputed tomography views of rabbit tibiae after 3, 6, 9, and 24 weeks of implantation of different CPC specimens. The values indicated at the bottom of the image represent residual implant areas, expressed in %.
8. Bone cell progenitor D1 seeded on the DCP-rich CPC surface initiated at an earlier stage of differentiation than on the c-CPC surface. Histological observations The residual implant area of DCP-rich CPC was significantly less than that of c-CPC at 3, 9, and 24 weeks after implantation (Fig. 5). The cross sections of the CPC implant sites indicated an increase in new BA (in %) over time [Fig. 6(a,b)]. The BA% of the DCP-rich CPC group was
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significantly higher than that of the c-CPC group even at 24 weeks (p < 0.05) postoperation. Regression between the DCP-rich effect of CPC composition and implantation time was reflected on the residual implant rate [Fig. 7(a)] and new bone formation rate [Fig. 7(b)]. The linear relationship of the residual implant rate was exhibited only in the DCPrich CPC and c-CPC implanted before 9 and 6 weeks, respectively. Compared with the residual implant rate, the linear regression between the new bone formation rate and implantation time was reflected in the DCP-rich CPC but not in the c-CPC. Figure 8 shows the typical histological morphologies of nondecalcified H&E-stained sections. Both CPC groups showed good tissue contact postimplantation, as indicated by the absence of inflammatory or other immune responses. DCP-rich CPC exhibited good bone remodeling, which increased the formation of new trabecular bone tissue around the implant sites (Fig. 8) over time. This increase was more prominent in DCP-rich CPC than in cCPC graft material at 9 and 24 weeks. DISCUSSION
The significant differences between DCP-rich CPC and c-CPC indicated that the null hypothesis of this study is consistent with the results. The cell viability, ALP activity in vitro, and bioresorption rate in vivo of DCP-rich CPC were higher than those of c-CPC. The early D1 cell behaviors of primary adhesion on the surfaces showed the quantity of surface cell attachment and influenced the succeeding mechanisms by spread morphology, proliferation, and differentiation.19 The chemical composition and ionic solubility of the biomaterial also significantly affect the interaction between cells.20,21 CPCs can release Ca21 and PO432 ions in the physiological environment. These ions attract cells to the material surface,
FIGURE 6. (a) Cross-sectional images of implantation with different bone cements showing color contrast images (orange for natural bone; gray for residual CPCs). (b) Total bone areas of different materials implanted in a rabbit femur for different lengths of time (n 5 3, symbol * means at the 0.05 level, the difference of the population was significantly different). [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
FIGURE 7. Regression analysis based on the curves of (a) residual implant in ratios and (b) bone area (%) for different time periods. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 8. Nondecalcified histological analysis of the medial condyle of the rabbit femur. Representative H&E-stained sections of different CPC cements after 3, 6, 9, and 24 weeks of implantation. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary. com.]
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FIGURE 9. Schematic of the biological mechanisms of bone cement, such as ion release, cell adhesion, cell proliferation, cell differentiation, and new bone regeneration.
which is modulated by physicochemical interactions at the beginning of cell seeding, as well as adsorption of fibronectin and vitronectin from the serum.18,22 Sogo et al.23 observed a significantly higher ALP activity in osteogenically induced human MSC cultured on a fibronectin-coated HA ceramic than on an untreated material. Figure 9 shows the summary of the simulated biological mechanisms of bone cement with ion release and the consequences of cell adhesion, proliferation, and differentiation, as well as new bone regeneration. Decreasing the TTCP/DCPA molar ratio from the traditional ratio of 1 to 0.33 increased the CPC dissolution rate by 40%.24 When CPC reduced the TTCP/DCPA molar ratio from 1 to 0.33, the human umbilical cord MSCs showed excellent viability and proliferation, as well as high ALP activity, osteocalcin content, collagen I content, and osterix gene expression.24 Adjusting CPC composition to enhance Ca21 and PO432 ion release can promote adhesion of many growth factors and proteins to the surface of the implant substances, as well as the activity and gene expression of osteoblasts.25 This study adjusted the surface-modified DCPA/TTCP molar ratio to 2/1 to increase the CPC dissolution rate. The change in molar ratio of DCP-rich CPC resulted in higher cell viability, ALP activity, and ALP quality than c-CPC, as observed in the D1 cell cultures (Figs. 1–4). D1 cells cultured on the DCP-rich CPC showed limited proliferation, which induced early biological transition from a proliferating progenitor to a differentiated postmitotic cell. After implantation and contact with body fluids, the DCPA of DCPD-forming CPC converted to DCPD and showed the primary dissolution of the cement.26 A linear and rapid resorption rate, which was observed in vivo (20 lm/day) in the DCPD-forming CPC after implantation, was similar to the speed of bone growth.27 However, the DCPD resorption was too quick for tissue regeneration and may cause a small gap for soft tissue growth between the implant and the newly formed bones.28 To achieve an ideal CPC resorption
rate, a technique was investigated to increase the molar ratio of DCPA and thus enhance the formation of product phases of DCPAs in the DCP-rich CPC. The residual implant rate indicated that a twofold DCPA molar ratio showed a deep resorption slope in early implantation, and resorption/ mineralization periods lasted for 24 weeks compared with c-CPC. The new bone formation rate of DCP-rich CPC continued to increase up to 24 weeks of implantation compared with that of c-CPC. The varied c-CPC showed a dt trend after 9 weeks. The new mechanism of bone formation in DCP-rich CPC continued for 24 weeks, but the new bone formation rate of c-CPC was almost the same during 9–24 weeks of implantation (Figs. 5 and 6). DCP-rich CPC and cCPC showed a linear regression for new bone formation and degradation rates before 9 and 6 weeks of implantation, respectively. In contrast to DCPD, apatite degraded very slowly because the bioresorb mechanism was solely facilitated by osteoclast activity rather than macrophage phagocytosis.29,30 After CPC implantation, the central area of the implanted DCP-rich CPC was converted into apatite over time. This phase was difficult to dissolve by osteoclast activity; thus, the linear relationship of bioresorption rate is not shown (Fig. 7). These findings indicate that DCP-rich CPC had a better dissolution rate than c-CPC in releasing large quantities of Ca21 and PO432 ions. Large amounts of surface-modified DCP formed a Ca21- and PO432-rich environment, which improved protein adhesion and enhanced biological performance on the specimen after implantation, as illustrated in Figure 9. The results of the in vitro test were consistent with those of the in vivo test. DCP-rich CPC performed better than c-CPC regardless of the complex interactions between different cell types and the implanted materials in the biological mechanisms in the in vivo test. In vitro test results may confirm the conclusions derived in vivo. In vitro analyses of cellular behavior, such as cell adhesion and proliferation, as well as mineralization on CPC substances in an
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in vitro environment, are low cost and ethical. In vitro examinations can help avoid controversial issues on the use of animals in experiments. New materials developed at an early stage can be evaluated with in vitro techniques. CONCLUSION
The formula of DCP-rich CPC contains a twofold higher surface-modified DCPA-to-TTCP molar ratio than c-CPC. DCPrich CPC showed improved cell attachment, viability, and ALP activity compared with c-CPC in vitro. Histological evaluation in vivo also confirmed that DCP-rich CPC exhibited higher bone regeneration efficiency than c-CPC. DCP-rich CPC dissolved faster than c-CPC, which resulted in a Ca21 and PO432 ion-rich environment. The enhanced protein adhesion on specimen surfaces and biological performance can be attributed to the large amount of ions in the environment. Improving bioresorption by adjustment of the composition of DCP-rich CPC can lead to high osteoconductivity and excellent biocompatibility for application in bone regeneration. ACKNOWLEDGMENTS
The authors sincerely acknowledge the assistance of Ms. HuiYu Wu and Ya-Shun Chen in this research. REFERENCES 1. Brown WE, Chow LC. A new calcium phosphate water setting cement. In: Brown PW, editor. Cements Research Progress. Westerville, OH: Am Ceram Soc; 1986. p 352–379. 2. Friedman CD, Costantino PD, Takagi S, Chow LC. BoneSourceTM hydroxyapatite cement: A novel biomaterial for craniofacial skeletal tissue engineering and reconstruction. J Biomed Mater Res 1998;43:428–432. 3. Dorozhkin SV. Calcium orthophosphate cements and concretes. Materials 2009;2:221–291. 4. Almirall A, Larrecq G, Delgado JA, Martinez S, Planell JA, Ginebra MP. Fabrication of low temperature macroporous hydroxyapatite scaffolds by foaming and hydrolysis of an alphaTCP paste. Biomaterials 2004;25:3671–3680. 5. Sun L, Xu HH, Takagi S, Chow LC. Fast setting calcium phosphate cement–chitosan composite: Mechanical properties and dissolution rates. J Biomater Appl 2007;21:299–315. 6. Chen JC, Ko CL, Shih CJ, Tien YC, Chen WC. Calcium phosphate bone cement with 10 wt% platelet-rich plasma in vitro and in vivo. J Dent 2012;40:114–122. 7. Perez RA, Ginebra MP, Spector M. Cell response to collagencalcium phosphate cement scaffolds investigated for nonviral gene delivery. J Mater Sci Mater Med 2011;22:887–897. 8. Tsigkou O, Jones JR, Polak JM, Stevens MM. Differentiation of fetal osteoblasts and formation of mineralized bone nodules by V 45S5 Bioglass conditioned medium in the absence of osteogenic supplements. Biomaterials 2009;30: 3542–3550. 9. Hirayama S, Takagi S, Markovic M, Chow LC. Properties of calcium phosphate cements with different tetracalcium phosphate and dicalcium phosphate anhydrous molar ratios. J Res Natl Inst Stand Technol 2008;113:311–320. 10. Wang JC, Ko CL, Hung CC, Tyan YC, Lai CH, Chen WC, Wang CK. Deriving fast setting properties of tetracalcium phosphate/dicalcium phosphate anhydrous bone cement with nanocrystallites on the reactant surfaces. J Dent 2010;38:158–165. 11. Ko CL, Chen JC, Tien YC, Hung CC, Wang JC, Chen WC. Biphasic products of dicalcium phosphate-rich cement with injectability and nondispersibility. Mater Sci Eng C Mater Biol Appl 2014;39: 40–46. doi: 10.1016/j.msec.2014.02.033. R
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