Materials Science and Engineering C 34 (2014) 110–114
Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Short communication
Low elastic modulus titanium–nickel scaffolds for bone implants Jing Li, Hailin Yang, Huifeng Wang, Jianming Ruan ⁎ State Key laboratory of Powder Metallurgy, Central South University, Changsha 410083, China
a r t i c l e
i n f o
Article history: Received 31 May 2013 Received in revised form 20 July 2013 Accepted 29 August 2013 Available online 7 September 2013 Keywords: Porous TiNi alloy Slurry immersing with polymer sponge and sintering method Pore structural properties Compressing properties Biological evaluation in vitro Cancellous bone substitute
a b s t r a c t The superelastic nature of repeating the human bones is crucial to the ideal artificial biomedical implants to ensure smooth load transfer and foster the ingrowth of new bone tissues. Three dimensional interconnected porous TiNi scaffolds, which have the tailorable porous structures with micro-hole, were fabricated by slurry immersing with polymer sponge and sintering method. The crystallinity and phase composition of scaffolds were studied by X-ray diffraction. The pore morphology, size and distribution in the scaffolds were characterized by scanning electron microscopy. The porosity ranged from 65 to 72%, pore size was 250–500 μm. Compressive strength and elastic modulus of the scaffolds were ~73 MPa and ~3GPa respectively. The above pore structural and mechanical properties are similar to those of cancellous bone. In the initial cell culture test, osteoblasts adhered well to the scaffold surface during a short time, and then grew smoothly into the interconnected pore channels. These results indicate that the porous TiNi scaffolds fabricated by this method could be bone substitute materials. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Bone tissue engineering has tremendous potential for bone regeneration. The osteogenic activity together with an appropriate artificial extracellular matrix can stimulate bone formation [1]. In the orthopedic field, scaffold material is very important for regeneration bone. The primary function of any orthopedic scaffold material is to achieve an initial stabilization of the construct, while inducing formation of bone from the surrounding tissue and acting as a carrier or template for implanted bone cells or other agents [2,3]. Near-equiatomic TiNi shape memory alloy can be regarded as a type of promising bone substitute for its unique shape memory effect, low elastic modulus, high strength, superior corrosion resistance and excellent biocompatibility such as new bone tissue ingrowth ability and vascularization [4–6], especially for the superelastic biomechanical properties, which are similar with some human hard tissues including bones and tendons [7–9]. However the high stiffness of dense metals compared with human bones often leads to heavy stress-shielding effect thereby increasing the healing time and even causing implant failure eventually [10]. Furthermore, in order to achieve a high cell density within the scaffold, a high specific surface area, pore connecting each other and pore size being large enough are desired. After cells having attached to the material surface there must be enough space for tissue ingrowth, as well as connecting channels to allow nutrient delivery, and waste removal, even cell proliferation and differentiation, etc. All of those could be carried out better on the scaffolds with suitable interconnected network of pores [11,12]. A number of techniques ⁎ Corresponding author. Tel.: +86 731 88836827. E-mail address: [email protected] (J. Ruan). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.08.043
have been developed to fabricate porous TiNi scaffolds, typically including self-propagating high-temperature synthesis (SHS), hot isostatic pressing (HIP), space-holder sintering process and other conventional sintering (CS) [13–18]. However, up to now, by the methods mentioned above, it is very difficult to construct a special micro-environment for tissue ingrowth, though the approaches to construct the microenvironment through controlling the pore sizes, pore structural and mechanical properties of metals and/or alloy scaffolds during material fabrication processing, are continually searched for. In the present study, TiNi porous scaffolds, with improved controllable pore structure and matched mechanical strength to bone, were processed with a novel technique for metallic biomaterials, in which the slurry immersing with polymer sponge and sintering method were integrated. The scaffold, in terms of its composition, morphology, mechanical properties and cell biological responses for extended clinical use, was mainly studied. 2. Material and methods A pre-alloyed, spherical equiatomic TiNi powder (Powder Metallurgy Research Institute of Central South University, China), with 99.8% purity and particles size between 5 and 15 μm, was selected to be the raw material for TiNi scaffold fabrication. A scanning electron micrograph (SEM) of TiNi powder is shown in Fig. 1. After mixing the prealloyed TiNi powder with Polyvinyl alcohol (PVA) (Tianjin Kemiou Chemical Reagent Co., China) solution (8 wt.% PVA + distilled water), the powder binder, the well-dispersed TiNi slurry was obtained. The open-cell polyurethane sponges (40 ppi) (Dongguan Inoac Polymer Co., Ltd., China), employed as the original scaffolds, were pretreated by washing in the 5 mol/L NaOH solution for 30 min at the
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
111
Fig. 1. SEM micrograph of TiNi pre-alloyed powder with the particle size of 5–15 μm.
Fig. 2. A flowchart of process steps for TiNi scaffold fabrication.
room temperature. Then the sponges were repeatedly immersed into the TiNi slurry and compressed to make it fill all of the pores. After being squeezed to get rid of the excess slurry, the impregnated sponges were dried at 40 °C for 24 h in vacuum. Subsequently, the green scaffolds were heated at 400 °C for 2 h to burn out the polyurethane sponges, and sintered at 1200 °C for 4 h in high vacuum (≦1.5 × 10−2 Pa). A flowchart of the process is given in Fig. 2. The crystallinity and phase composition of as-received TiNi prealloyed powder and the sintered samples were characterized by X-ray diffraction analysis at room temperature. The macrostructures and microstructures of TiNi scaffolds were observed by SEM.
Fig. 4. Macrographic morphology of TiNi scaffolds (a) and SEM microstructure of TiNi scaffolds (b) compared with that of human cancellous bone (c).
Fig. 3. XRD patterns for TiNi powder as-received (a); TiNi sintered at 1200 °C (b).
The porosity of the TiNi scaffolds was determined using the liquid displacement method [12]. A scaffold with the weight of W was immersed in a graduated cylinder containing a known volume (V1) of liquid paraffin of which the volume (V1) was known. In order to force
112
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
Real density (ρr) value of TiNi is 6.45 g·cm−3, so the porosity of the scaffold, ε was evaluated as ε ¼ ρ=ρr :
ð3Þ
The porosity of the open pores in the scaffold, εo was evaluated as ε o ¼ ðV 1 −V 3 Þ=ðV 2 −V 3 Þ:
Fig. 5. Stress–strain curves of TiNi scaffolds submitted at compression test.
the liquid paraffin into the pores of the scaffolds, the cylinder was placed in vacuum until no air bubble emerged from the scaffolds. The total volume of the liquid paraffin and scaffold was recorded as V2. So (V2 − V1) was the volume of the skeleton of the scaffold. The scaffold was removed from the liquid paraffin and the residual liquid paraffin volume was recorded as V3. The total volume of the scaffold, V, was evaluated as V ¼ V 2 −V 3 :
ð1Þ
The apparent density of the scaffold, ρ was evaluated as ρ ¼ W=ðV 2 −V 3 Þ:
ð2Þ
ð4Þ
Compression mechanical tests were conducted on the scaffolds (10 mm diameter, 25 mm height) at room temperature using the MTS testing machine with 10 kN load cell at a constant crosshead speed of 0.5 mm/min. The slope of the initial linear portion of the stress–strain curve was used to define the elastic modulus. The cross point of the two tangents on the stress–strain curve around the yield point was used to calculate the yield strength. The test disks with a gage diameter of 10 mm and a gage thickness of 2 mm were machined from the porous TiNi scaffolds to study the cell biological responses in vitro. The scaffold disks were sterilized in an autoclave prior to be seeded with the human fetal osteoblast cell line (hFOB) and then cultured in Dulbecco's Modified Eagle Medium (DMEM) with 10% Fetal Calf Serum (FCS) at 37 °C and 5% CO2-atmosphere for 5 days, 7 days, 9 days and 11 days. After seeding and fixed in Phosphate Buffered Solution (PBS) with 3% glutaraldehyde for 2 days at 4 °C, the scaffold disks were washed twice by PBS. The samples were then dehydrated using a graded ethanol series from 10 to 100%, with three times 10 min incubation at each step. Dehydration was completed by critical point drying. The scaffolds were sputtered and cell adherence was qualitatively analyzed by SEM.
Table 1 Porosity and compressive properties of the porous NiTi scaffolds. Total porosity (%) Open porosity (%) Fraction of open porosity (%) Compressive strength (MPa) Elastic modulus (GPa)
55.93 ± 0.89 54.19 ± 0.52 96.89 97.60 ± 1.06 3.99 ± 0.04
66.56 ± 0.76 66.23 ± 0.61 99.50 78.11 ± 1.59 3.67 ± 0.07
69.24 ± 0.76 68.92 ± 0.59 99.54 74.82 ± 1.44 3.41 ± 0.06
71.58 ± 0.89 71.33 ± 0.56 99.65 67.36 ± 1.57 2.86 ± 0.08
Fig. 6. Compressive strength and elastic modulus of TiNi scaffolds compared with that of cancellous bone and cortical bone.
79.33 ± 1.24 79.17 ± 0.75 99.79 26.37 ± 2.03 0.81 ± 0.09
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
113
Fig. 7. SEM micrographs illustrating bone cell morphology after 5 days (a), 7 days (b), 9 days (c) and 11 days (d) of culture on the porous TiNi scaffold.
3. Results and discussions The crystallographic composition of TiNi pre-alloyed powder and the porous TiNi alloy scaffolds are shown in Fig. 3. After sintering at 1200 °C, the porous TiNi alloy scaffolds were found to be in the fully austenitic state, being similar with the TiNi powder as-received, while oxidation was found to have been prevented, and secondary intermetallics, TiC, were inconspicuously identified, as revealed by the results of the XRD analysis (Fig. 3). In Fig. 3, the peaks of both diffraction patterns were matched well with those of standard TiNi in the powder diffraction file (Card No. 65-4572), but X-ray diffraction line broadening was obvious in the pattern of sintered TiNi scaffold. Macrographic morphology from the TiNi scaffolds after sintering is shown in Fig. 4(a). This photograph presents that continuous open macroporous TiNi scaffolds have been successfully produced, which duplicate basic structural characteristics of polyurethane sponge. As shown in Fig. 4(b), a three-dimensionally interconnected, permeable structure with pore sizes varying from 250 to 500 μm, and the microporous structure on the macroporous skeleton was observed by SEM. Those characteristics are quite similar to human cancellous bone (Fig. 4(c)). Flatly et al. [19] reported that the optimum pore size for
osteoconduction was 500 μm. Itin et al. [20] showed that a pore size of 100–500 μm is compatible for new bone ingrowth. The appropriate pore morphology of TiNi scaffolds, the optimum pore size and interconnected macropore with micropore structure, like the ones shown in Fig. 4(b) should be beneficial for cell adhesion, proliferation and differentiation, rapid vascularization or bone ingrowth and remodeling [8,21–24]. The connected pores in the scaffolds allow also biomolecules, degraded substances and biofluid to freely flow into and out of the scaffold [6]. Therefore, porous TiNi scaffolds with 3D interconnected pore morphology and similar macro- and microarchitecture to cancellous bone, can be fabricated by slurry immersing with polymer sponge and sintering method. The open porosity of TiNi scaffolds ranges from 65% to 72%, which share over 99.5% total porosity, as the liquid displacement method determined. The strength, permeability, and structural defects of a porous scaffold highly relate to its porosity [25]. The high porosity, especially open porosity, which would lead to high specific surface area, promotes cell adhesion to the scaffold and promotes bone tissue regeneration. As the scaffolds duplicate the morphology of polyurethane sponges, shown in Fig. 4, the polyurethane sponge type influences the porosity of TiNi scaffolds greatly. The TiNi slurry concentration in the sponge is an
114
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
important influencing factor as well. With the increase of TiNi content, the open porosity decreased, which would lead to higher mechanical strength and less favorable biological environment, so the appropriate TiNi content must be found for the specific application. The compressive strength and elastic modulus of TiNi scaffolds were studied using the compression test. The compressive stress–strain curve of TiNi scaffolds is shown in Fig. 5. This curve is similar to the typical porous materials with open pores showing the three regimes of behavior [26,27], as linear elasticity, non-linear elasticity and densification, and plastic collapse and densification. While plateau is relatively short and densification has not occurred in the third regime of TiNi's curve, the scaffolds are collapsed when the moment exerted on the pore walls exceeds the fully plastic moment. These results mentioned above can attribute to the excessive porosity, and the existence of TiC, a brittle phase, together with the lack of excellent plasticity of TiNi, which is common in intermetallics. The mechanical properties of porous metallic materials are mainly determined by porosity and the morphology, size and distribution of the pores [23,28]. The relationship between porosity and compressive properties is shown in Table 1. The compressive strength and elastic modulus of TiNi scaffolds decrease with increase of the porosity. The comparison of elastic modulus and compressive strength of the TiNi scaffolds with those of cancellous bone and cortical bone is shown in Fig. 6. The compressive strength and compressive elastic modulus of the TiNi scaffolds are 73.52 MPa and 3.36 GPa, respectively, which are located between the cortical bone (130–150 MPa, 4.4–28.8 GPa) and cancellous bone (10–50 MPa, 0.01–3.0 GPa) [29], and met the mechanical properties of requirement of trabecular bone well. This fact is very attractive for its potential to minimize the stress shielding effect and to optimize short term and long term fixation. Fig. 7 shows the cell growth on the surface of the porous TiNi scaffolds after 5, 7, 9, 11 days of incubation period. Within 5 days, bone cells successfully anchored onto the porous surface, and proliferated well (Fig. 7(a)). After 7 days of culturing, the osteoblasts had migrated into the pores and were bridged to the substrate in addition to neighboring cells (Fig. 7(b)). There was a large quantity of pseudopodia or microvilli, as the arrow in Fig. 7(b) showed. Cells appeared to be more elongated and connected with each other then grew in a monolayer and had a flat, well-spread-out morphology, as shown in Fig. 7(c). Fig. 7(d) shows that cells spread to almost the entire surface, including inner surface, and developed into a multilayer, which indicates that the inner surface of pores also possesses good osteoconductivity, and bone cells can fully link and grow on the inner surface of the pores. The scaffold's structural parameters lead to direct osteogenesis, which play crucial roles not only in the cellular adhesion properties, but also in cell viability, ingrowth, distribution and the formation of an extracellular matrix [8,30]. Porous TiNi scaffold does not inhibit osteoblast proliferation and even accelerates cell adhesion and growth. 3D structure and rough surface created by slurry immersing with polymer sponge and sintering method can promote cell adhesion, trophic circulation and substance metabolism. However, other methods, such as conventional sintering, HIP and SHS, have limitations in formation of the three-dimensional structure.
4. Conclusions Porous TiNi scaffolds with suitable pore property and good biocompatibility were successfully fabricated by a novel approach combining slurry immersing with polymer sponge and sintering methods. The composition of TiNi scaffolds remains the same as that of the initial
powder as-received. A compressive strength of 73.52 MPa and modulus of 3.36 GPa for the scaffold agree well with the mechanical properties of requirement of trabecular bone. The TiNi scaffolds with a 3D interconnected macroporous structure have an open porosity and a pore size ranging from 65 to 72% and from 250 to 500 μm, respectively. The pore size and shape of the produced scaffold are controllable, as its macroporous structure is the replicate of the polymer sponge template. Furthermore, the relationships between porous structure of TiNi scaffolds and osteoblast ingrowth as well as formation of bone tissues are investigated. The in vitro results reveal that osteoblasts can adhere to and proliferate well on the entire surface of the 3D interconnected macroporous TiNi scaffold. Moreover, osteoblasts penetrate into the scaffold through the pores and interconnected channels, which can enhance biological fixation. These results suggest that the porous TiNi scaffolds fabricated by slurry immersing with polymer sponge and sintering method have broad potential for clinical applications. Acknowledgments The authors would like to acknowledge the Fundamental Research Funds for the Central Universities of Central South University (No. 2012zztsa012), the National Natural Science Foundation of China (No. 51274247) and the National High Technology Research and Development Program (No. 2012BAE06B00) for their financial support. References [1] T.L. Livingston, Bioactive Foam for Bone Tissue Engineering: An In Vivo Study, 1999. [2] J.D. Bobyn, G.J. Stackpool, S.A. Hacking, M. Tanzer, J.J. Krygier, J. Bone Joint Surg. Br. 81B (1999) 907–914. [3] C.A. Vacanti, L.J. Bonassar, Clin. Orthop. Relat. Res. (1999) S375–S381. [4] D.J. Hoh, B.L. Hoh, A.P. Amar, M.Y. Wang, Neurosurgery 64 (2009) 199–214. [5] A. Bansiddhi, T.D. Sargeant, S.I. Stupp, D.C. Dunand, Acta Biomater. 4 (2008) 773–782. [6] O. Prymak, D. Bogdanski, M. Koller, S.A. Esenwein, G. Muhr, F. Beckmann, T. Donath, M. Assad, M. Epple, Biomaterials 26 (2005) 5801–5807. [7] C. Greiner, S.M. Oppenheimer, D.C. Dunand, Acta Biomater. 1 (2005) 705–716. [8] X.M. Liu, S.L. Wu, K.W.K. Yeung, Y.L. Chan, T. Hu, Z.S. Xu, X.Y. Liu, J.C.Y. Chung, K.M.C. Cheung, P.K. Chu, Biomaterials 32 (2011) 330–338. [9] C.H. Turner, Osteoporos. Int. 13 (2002) 97–104. [10] S. Sahin, M.C. Cehreli, E. Yalcin, J. Dent. 30 (2002) 271–282. [11] M.C. Peters, D.J. Mooney, Mater. Sci. Forum 250 (1997) 43–52. [12] H.R. Ramay, M.Q. Zhang, Biomaterials 24 (2003) 3293–3302. [13] G. Tosun, N. Orhan, L. Ozler, Mater. Lett. 66 (2012) 138–140. [14] S.L. Wu, C.Y. Chung, X.M. Liu, P.K. Chu, J.P.Y. Ho, C.L. Chu, Y.L. Chan, K.W.K. Yeung, W.W. Lu, K.M.C. Cheung, K.D.K. Luk, Acta Mater. 55 (2007) 3437–3451. [15] G.I. Nakas, A.F. Dericioglu, S. Bor, J. Mech. Behav. Biomed. 4 (2011) 2017–2023. [16] S.A. Hosseini, S.K. Sadrnezhaad, A. Ekrami, Mater. Sci. Eng. C Mater. 29 (2009) 2203–2207. [17] M.H. Ismail, R. Goodall, H.A. Davies, I. Todd, Mater. Sci. Eng. C Mater. 32 (2012) 1480–1485. [18] M. Kohl, T. Habijan, M. Bram, H.P. Buchkremer, D. Stover, M. Koller, Adv. Eng. Mater. 11 (2009) 959–968. [19] T.J. Flatly, K.L. Lynch, M. Benson, Clin. Orthop. Relat. Res. 179 (1983) 246–252. [20] V.I. Itin, V.E. Gjunter, S.A. Shabalovskaya, R.L.C. Sachdeva, Mater. Charact. 32 (1994) 179–187. [21] G. Tripathi, B. Basu, Ceram. Int. 38 (2012) 341–349. [22] F. Likibi, M. Assad, C. Coillard, G. Chabot, C.H. Rivard, Ann. Chir. 130 (2005) 235–241. [23] V. Karageorgiou, D. Kaplan, Biomaterials 26 (2005) 5474–5491. [24] B.S. Chang, C.K. Lee, K.S. Hong, H.J. Youn, H.S. Ryu, S.S. Chung, K.W. Park, Biomaterials 21 (2000) 1291–1298. [25] P. Sepulveda, J.G.P. Binner, J. Eur. Ceram. Soc. 19 (1999) 2059–2066. [26] F. Linde, P. Norgaard, I. Hvid, A. Odgaard, K. Soballe, J. Biomech. 24 (1991) 803–809. [27] D.A. Shimko, V.F. Shimko, E.A. Sander, K.F. Dickson, E.A. Nauman, J. Biomed. Mater. Res. B 73B (2005) 315–324. [28] L.D. Zardiackas, D.E. Parsell, L.D. Dillon, D.W. Mitchell, L.A. Nunnery, R. Poggie, J. Biomed. Mater. Res. 58 (2001) 180–187. [29] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Prog. Mater. Sci. 54 (2009) 397–425. [30] H. Haugen, J. Will, A. Kohler, U. Hopfner, J. Aigner, E. Wintermantel, J. Eur. Ceram. Soc. 24 (2004) 661–668.
Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Short communication
Low elastic modulus titanium–nickel scaffolds for bone implants Jing Li, Hailin Yang, Huifeng Wang, Jianming Ruan ⁎ State Key laboratory of Powder Metallurgy, Central South University, Changsha 410083, China
a r t i c l e
i n f o
Article history: Received 31 May 2013 Received in revised form 20 July 2013 Accepted 29 August 2013 Available online 7 September 2013 Keywords: Porous TiNi alloy Slurry immersing with polymer sponge and sintering method Pore structural properties Compressing properties Biological evaluation in vitro Cancellous bone substitute
a b s t r a c t The superelastic nature of repeating the human bones is crucial to the ideal artificial biomedical implants to ensure smooth load transfer and foster the ingrowth of new bone tissues. Three dimensional interconnected porous TiNi scaffolds, which have the tailorable porous structures with micro-hole, were fabricated by slurry immersing with polymer sponge and sintering method. The crystallinity and phase composition of scaffolds were studied by X-ray diffraction. The pore morphology, size and distribution in the scaffolds were characterized by scanning electron microscopy. The porosity ranged from 65 to 72%, pore size was 250–500 μm. Compressive strength and elastic modulus of the scaffolds were ~73 MPa and ~3GPa respectively. The above pore structural and mechanical properties are similar to those of cancellous bone. In the initial cell culture test, osteoblasts adhered well to the scaffold surface during a short time, and then grew smoothly into the interconnected pore channels. These results indicate that the porous TiNi scaffolds fabricated by this method could be bone substitute materials. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Bone tissue engineering has tremendous potential for bone regeneration. The osteogenic activity together with an appropriate artificial extracellular matrix can stimulate bone formation [1]. In the orthopedic field, scaffold material is very important for regeneration bone. The primary function of any orthopedic scaffold material is to achieve an initial stabilization of the construct, while inducing formation of bone from the surrounding tissue and acting as a carrier or template for implanted bone cells or other agents [2,3]. Near-equiatomic TiNi shape memory alloy can be regarded as a type of promising bone substitute for its unique shape memory effect, low elastic modulus, high strength, superior corrosion resistance and excellent biocompatibility such as new bone tissue ingrowth ability and vascularization [4–6], especially for the superelastic biomechanical properties, which are similar with some human hard tissues including bones and tendons [7–9]. However the high stiffness of dense metals compared with human bones often leads to heavy stress-shielding effect thereby increasing the healing time and even causing implant failure eventually [10]. Furthermore, in order to achieve a high cell density within the scaffold, a high specific surface area, pore connecting each other and pore size being large enough are desired. After cells having attached to the material surface there must be enough space for tissue ingrowth, as well as connecting channels to allow nutrient delivery, and waste removal, even cell proliferation and differentiation, etc. All of those could be carried out better on the scaffolds with suitable interconnected network of pores [11,12]. A number of techniques ⁎ Corresponding author. Tel.: +86 731 88836827. E-mail address: [email protected] (J. Ruan). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.08.043
have been developed to fabricate porous TiNi scaffolds, typically including self-propagating high-temperature synthesis (SHS), hot isostatic pressing (HIP), space-holder sintering process and other conventional sintering (CS) [13–18]. However, up to now, by the methods mentioned above, it is very difficult to construct a special micro-environment for tissue ingrowth, though the approaches to construct the microenvironment through controlling the pore sizes, pore structural and mechanical properties of metals and/or alloy scaffolds during material fabrication processing, are continually searched for. In the present study, TiNi porous scaffolds, with improved controllable pore structure and matched mechanical strength to bone, were processed with a novel technique for metallic biomaterials, in which the slurry immersing with polymer sponge and sintering method were integrated. The scaffold, in terms of its composition, morphology, mechanical properties and cell biological responses for extended clinical use, was mainly studied. 2. Material and methods A pre-alloyed, spherical equiatomic TiNi powder (Powder Metallurgy Research Institute of Central South University, China), with 99.8% purity and particles size between 5 and 15 μm, was selected to be the raw material for TiNi scaffold fabrication. A scanning electron micrograph (SEM) of TiNi powder is shown in Fig. 1. After mixing the prealloyed TiNi powder with Polyvinyl alcohol (PVA) (Tianjin Kemiou Chemical Reagent Co., China) solution (8 wt.% PVA + distilled water), the powder binder, the well-dispersed TiNi slurry was obtained. The open-cell polyurethane sponges (40 ppi) (Dongguan Inoac Polymer Co., Ltd., China), employed as the original scaffolds, were pretreated by washing in the 5 mol/L NaOH solution for 30 min at the
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
111
Fig. 1. SEM micrograph of TiNi pre-alloyed powder with the particle size of 5–15 μm.
Fig. 2. A flowchart of process steps for TiNi scaffold fabrication.
room temperature. Then the sponges were repeatedly immersed into the TiNi slurry and compressed to make it fill all of the pores. After being squeezed to get rid of the excess slurry, the impregnated sponges were dried at 40 °C for 24 h in vacuum. Subsequently, the green scaffolds were heated at 400 °C for 2 h to burn out the polyurethane sponges, and sintered at 1200 °C for 4 h in high vacuum (≦1.5 × 10−2 Pa). A flowchart of the process is given in Fig. 2. The crystallinity and phase composition of as-received TiNi prealloyed powder and the sintered samples were characterized by X-ray diffraction analysis at room temperature. The macrostructures and microstructures of TiNi scaffolds were observed by SEM.
Fig. 4. Macrographic morphology of TiNi scaffolds (a) and SEM microstructure of TiNi scaffolds (b) compared with that of human cancellous bone (c).
Fig. 3. XRD patterns for TiNi powder as-received (a); TiNi sintered at 1200 °C (b).
The porosity of the TiNi scaffolds was determined using the liquid displacement method [12]. A scaffold with the weight of W was immersed in a graduated cylinder containing a known volume (V1) of liquid paraffin of which the volume (V1) was known. In order to force
112
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
Real density (ρr) value of TiNi is 6.45 g·cm−3, so the porosity of the scaffold, ε was evaluated as ε ¼ ρ=ρr :
ð3Þ
The porosity of the open pores in the scaffold, εo was evaluated as ε o ¼ ðV 1 −V 3 Þ=ðV 2 −V 3 Þ:
Fig. 5. Stress–strain curves of TiNi scaffolds submitted at compression test.
the liquid paraffin into the pores of the scaffolds, the cylinder was placed in vacuum until no air bubble emerged from the scaffolds. The total volume of the liquid paraffin and scaffold was recorded as V2. So (V2 − V1) was the volume of the skeleton of the scaffold. The scaffold was removed from the liquid paraffin and the residual liquid paraffin volume was recorded as V3. The total volume of the scaffold, V, was evaluated as V ¼ V 2 −V 3 :
ð1Þ
The apparent density of the scaffold, ρ was evaluated as ρ ¼ W=ðV 2 −V 3 Þ:
ð2Þ
ð4Þ
Compression mechanical tests were conducted on the scaffolds (10 mm diameter, 25 mm height) at room temperature using the MTS testing machine with 10 kN load cell at a constant crosshead speed of 0.5 mm/min. The slope of the initial linear portion of the stress–strain curve was used to define the elastic modulus. The cross point of the two tangents on the stress–strain curve around the yield point was used to calculate the yield strength. The test disks with a gage diameter of 10 mm and a gage thickness of 2 mm were machined from the porous TiNi scaffolds to study the cell biological responses in vitro. The scaffold disks were sterilized in an autoclave prior to be seeded with the human fetal osteoblast cell line (hFOB) and then cultured in Dulbecco's Modified Eagle Medium (DMEM) with 10% Fetal Calf Serum (FCS) at 37 °C and 5% CO2-atmosphere for 5 days, 7 days, 9 days and 11 days. After seeding and fixed in Phosphate Buffered Solution (PBS) with 3% glutaraldehyde for 2 days at 4 °C, the scaffold disks were washed twice by PBS. The samples were then dehydrated using a graded ethanol series from 10 to 100%, with three times 10 min incubation at each step. Dehydration was completed by critical point drying. The scaffolds were sputtered and cell adherence was qualitatively analyzed by SEM.
Table 1 Porosity and compressive properties of the porous NiTi scaffolds. Total porosity (%) Open porosity (%) Fraction of open porosity (%) Compressive strength (MPa) Elastic modulus (GPa)
55.93 ± 0.89 54.19 ± 0.52 96.89 97.60 ± 1.06 3.99 ± 0.04
66.56 ± 0.76 66.23 ± 0.61 99.50 78.11 ± 1.59 3.67 ± 0.07
69.24 ± 0.76 68.92 ± 0.59 99.54 74.82 ± 1.44 3.41 ± 0.06
71.58 ± 0.89 71.33 ± 0.56 99.65 67.36 ± 1.57 2.86 ± 0.08
Fig. 6. Compressive strength and elastic modulus of TiNi scaffolds compared with that of cancellous bone and cortical bone.
79.33 ± 1.24 79.17 ± 0.75 99.79 26.37 ± 2.03 0.81 ± 0.09
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
113
Fig. 7. SEM micrographs illustrating bone cell morphology after 5 days (a), 7 days (b), 9 days (c) and 11 days (d) of culture on the porous TiNi scaffold.
3. Results and discussions The crystallographic composition of TiNi pre-alloyed powder and the porous TiNi alloy scaffolds are shown in Fig. 3. After sintering at 1200 °C, the porous TiNi alloy scaffolds were found to be in the fully austenitic state, being similar with the TiNi powder as-received, while oxidation was found to have been prevented, and secondary intermetallics, TiC, were inconspicuously identified, as revealed by the results of the XRD analysis (Fig. 3). In Fig. 3, the peaks of both diffraction patterns were matched well with those of standard TiNi in the powder diffraction file (Card No. 65-4572), but X-ray diffraction line broadening was obvious in the pattern of sintered TiNi scaffold. Macrographic morphology from the TiNi scaffolds after sintering is shown in Fig. 4(a). This photograph presents that continuous open macroporous TiNi scaffolds have been successfully produced, which duplicate basic structural characteristics of polyurethane sponge. As shown in Fig. 4(b), a three-dimensionally interconnected, permeable structure with pore sizes varying from 250 to 500 μm, and the microporous structure on the macroporous skeleton was observed by SEM. Those characteristics are quite similar to human cancellous bone (Fig. 4(c)). Flatly et al. [19] reported that the optimum pore size for
osteoconduction was 500 μm. Itin et al. [20] showed that a pore size of 100–500 μm is compatible for new bone ingrowth. The appropriate pore morphology of TiNi scaffolds, the optimum pore size and interconnected macropore with micropore structure, like the ones shown in Fig. 4(b) should be beneficial for cell adhesion, proliferation and differentiation, rapid vascularization or bone ingrowth and remodeling [8,21–24]. The connected pores in the scaffolds allow also biomolecules, degraded substances and biofluid to freely flow into and out of the scaffold [6]. Therefore, porous TiNi scaffolds with 3D interconnected pore morphology and similar macro- and microarchitecture to cancellous bone, can be fabricated by slurry immersing with polymer sponge and sintering method. The open porosity of TiNi scaffolds ranges from 65% to 72%, which share over 99.5% total porosity, as the liquid displacement method determined. The strength, permeability, and structural defects of a porous scaffold highly relate to its porosity [25]. The high porosity, especially open porosity, which would lead to high specific surface area, promotes cell adhesion to the scaffold and promotes bone tissue regeneration. As the scaffolds duplicate the morphology of polyurethane sponges, shown in Fig. 4, the polyurethane sponge type influences the porosity of TiNi scaffolds greatly. The TiNi slurry concentration in the sponge is an
114
J. Li et al. / Materials Science and Engineering C 34 (2014) 110–114
important influencing factor as well. With the increase of TiNi content, the open porosity decreased, which would lead to higher mechanical strength and less favorable biological environment, so the appropriate TiNi content must be found for the specific application. The compressive strength and elastic modulus of TiNi scaffolds were studied using the compression test. The compressive stress–strain curve of TiNi scaffolds is shown in Fig. 5. This curve is similar to the typical porous materials with open pores showing the three regimes of behavior [26,27], as linear elasticity, non-linear elasticity and densification, and plastic collapse and densification. While plateau is relatively short and densification has not occurred in the third regime of TiNi's curve, the scaffolds are collapsed when the moment exerted on the pore walls exceeds the fully plastic moment. These results mentioned above can attribute to the excessive porosity, and the existence of TiC, a brittle phase, together with the lack of excellent plasticity of TiNi, which is common in intermetallics. The mechanical properties of porous metallic materials are mainly determined by porosity and the morphology, size and distribution of the pores [23,28]. The relationship between porosity and compressive properties is shown in Table 1. The compressive strength and elastic modulus of TiNi scaffolds decrease with increase of the porosity. The comparison of elastic modulus and compressive strength of the TiNi scaffolds with those of cancellous bone and cortical bone is shown in Fig. 6. The compressive strength and compressive elastic modulus of the TiNi scaffolds are 73.52 MPa and 3.36 GPa, respectively, which are located between the cortical bone (130–150 MPa, 4.4–28.8 GPa) and cancellous bone (10–50 MPa, 0.01–3.0 GPa) [29], and met the mechanical properties of requirement of trabecular bone well. This fact is very attractive for its potential to minimize the stress shielding effect and to optimize short term and long term fixation. Fig. 7 shows the cell growth on the surface of the porous TiNi scaffolds after 5, 7, 9, 11 days of incubation period. Within 5 days, bone cells successfully anchored onto the porous surface, and proliferated well (Fig. 7(a)). After 7 days of culturing, the osteoblasts had migrated into the pores and were bridged to the substrate in addition to neighboring cells (Fig. 7(b)). There was a large quantity of pseudopodia or microvilli, as the arrow in Fig. 7(b) showed. Cells appeared to be more elongated and connected with each other then grew in a monolayer and had a flat, well-spread-out morphology, as shown in Fig. 7(c). Fig. 7(d) shows that cells spread to almost the entire surface, including inner surface, and developed into a multilayer, which indicates that the inner surface of pores also possesses good osteoconductivity, and bone cells can fully link and grow on the inner surface of the pores. The scaffold's structural parameters lead to direct osteogenesis, which play crucial roles not only in the cellular adhesion properties, but also in cell viability, ingrowth, distribution and the formation of an extracellular matrix [8,30]. Porous TiNi scaffold does not inhibit osteoblast proliferation and even accelerates cell adhesion and growth. 3D structure and rough surface created by slurry immersing with polymer sponge and sintering method can promote cell adhesion, trophic circulation and substance metabolism. However, other methods, such as conventional sintering, HIP and SHS, have limitations in formation of the three-dimensional structure.
4. Conclusions Porous TiNi scaffolds with suitable pore property and good biocompatibility were successfully fabricated by a novel approach combining slurry immersing with polymer sponge and sintering methods. The composition of TiNi scaffolds remains the same as that of the initial
powder as-received. A compressive strength of 73.52 MPa and modulus of 3.36 GPa for the scaffold agree well with the mechanical properties of requirement of trabecular bone. The TiNi scaffolds with a 3D interconnected macroporous structure have an open porosity and a pore size ranging from 65 to 72% and from 250 to 500 μm, respectively. The pore size and shape of the produced scaffold are controllable, as its macroporous structure is the replicate of the polymer sponge template. Furthermore, the relationships between porous structure of TiNi scaffolds and osteoblast ingrowth as well as formation of bone tissues are investigated. The in vitro results reveal that osteoblasts can adhere to and proliferate well on the entire surface of the 3D interconnected macroporous TiNi scaffold. Moreover, osteoblasts penetrate into the scaffold through the pores and interconnected channels, which can enhance biological fixation. These results suggest that the porous TiNi scaffolds fabricated by slurry immersing with polymer sponge and sintering method have broad potential for clinical applications. Acknowledgments The authors would like to acknowledge the Fundamental Research Funds for the Central Universities of Central South University (No. 2012zztsa012), the National Natural Science Foundation of China (No. 51274247) and the National High Technology Research and Development Program (No. 2012BAE06B00) for their financial support. References [1] T.L. Livingston, Bioactive Foam for Bone Tissue Engineering: An In Vivo Study, 1999. [2] J.D. Bobyn, G.J. Stackpool, S.A. Hacking, M. Tanzer, J.J. Krygier, J. Bone Joint Surg. Br. 81B (1999) 907–914. [3] C.A. Vacanti, L.J. Bonassar, Clin. Orthop. Relat. Res. (1999) S375–S381. [4] D.J. Hoh, B.L. Hoh, A.P. Amar, M.Y. Wang, Neurosurgery 64 (2009) 199–214. [5] A. Bansiddhi, T.D. Sargeant, S.I. Stupp, D.C. Dunand, Acta Biomater. 4 (2008) 773–782. [6] O. Prymak, D. Bogdanski, M. Koller, S.A. Esenwein, G. Muhr, F. Beckmann, T. Donath, M. Assad, M. Epple, Biomaterials 26 (2005) 5801–5807. [7] C. Greiner, S.M. Oppenheimer, D.C. Dunand, Acta Biomater. 1 (2005) 705–716. [8] X.M. Liu, S.L. Wu, K.W.K. Yeung, Y.L. Chan, T. Hu, Z.S. Xu, X.Y. Liu, J.C.Y. Chung, K.M.C. Cheung, P.K. Chu, Biomaterials 32 (2011) 330–338. [9] C.H. Turner, Osteoporos. Int. 13 (2002) 97–104. [10] S. Sahin, M.C. Cehreli, E. Yalcin, J. Dent. 30 (2002) 271–282. [11] M.C. Peters, D.J. Mooney, Mater. Sci. Forum 250 (1997) 43–52. [12] H.R. Ramay, M.Q. Zhang, Biomaterials 24 (2003) 3293–3302. [13] G. Tosun, N. Orhan, L. Ozler, Mater. Lett. 66 (2012) 138–140. [14] S.L. Wu, C.Y. Chung, X.M. Liu, P.K. Chu, J.P.Y. Ho, C.L. Chu, Y.L. Chan, K.W.K. Yeung, W.W. Lu, K.M.C. Cheung, K.D.K. Luk, Acta Mater. 55 (2007) 3437–3451. [15] G.I. Nakas, A.F. Dericioglu, S. Bor, J. Mech. Behav. Biomed. 4 (2011) 2017–2023. [16] S.A. Hosseini, S.K. Sadrnezhaad, A. Ekrami, Mater. Sci. Eng. C Mater. 29 (2009) 2203–2207. [17] M.H. Ismail, R. Goodall, H.A. Davies, I. Todd, Mater. Sci. Eng. C Mater. 32 (2012) 1480–1485. [18] M. Kohl, T. Habijan, M. Bram, H.P. Buchkremer, D. Stover, M. Koller, Adv. Eng. Mater. 11 (2009) 959–968. [19] T.J. Flatly, K.L. Lynch, M. Benson, Clin. Orthop. Relat. Res. 179 (1983) 246–252. [20] V.I. Itin, V.E. Gjunter, S.A. Shabalovskaya, R.L.C. Sachdeva, Mater. Charact. 32 (1994) 179–187. [21] G. Tripathi, B. Basu, Ceram. Int. 38 (2012) 341–349. [22] F. Likibi, M. Assad, C. Coillard, G. Chabot, C.H. Rivard, Ann. Chir. 130 (2005) 235–241. [23] V. Karageorgiou, D. Kaplan, Biomaterials 26 (2005) 5474–5491. [24] B.S. Chang, C.K. Lee, K.S. Hong, H.J. Youn, H.S. Ryu, S.S. Chung, K.W. Park, Biomaterials 21 (2000) 1291–1298. [25] P. Sepulveda, J.G.P. Binner, J. Eur. Ceram. Soc. 19 (1999) 2059–2066. [26] F. Linde, P. Norgaard, I. Hvid, A. Odgaard, K. Soballe, J. Biomech. 24 (1991) 803–809. [27] D.A. Shimko, V.F. Shimko, E.A. Sander, K.F. Dickson, E.A. Nauman, J. Biomed. Mater. Res. B 73B (2005) 315–324. [28] L.D. Zardiackas, D.E. Parsell, L.D. Dillon, D.W. Mitchell, L.A. Nunnery, R. Poggie, J. Biomed. Mater. Res. 58 (2001) 180–187. [29] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Prog. Mater. Sci. 54 (2009) 397–425. [30] H. Haugen, J. Will, A. Kohler, U. Hopfner, J. Aigner, E. Wintermantel, J. Eur. Ceram. Soc. 24 (2004) 661–668.
