**7. Results and discussion**

#### **7.1. Physico-chemical properties of the Ti14Nb4Sn alloy**

The X-ray diffraction pattern of sintered Ti14Nb4Sn is shown in Figure 5. Alpha peaks were observed at 39.0° and 40.5°, which are indexed as the reflection planes (101) and (103), while β peaks were observed at 38.5°, which is indexed as (110). The titanium alloy consisted of both α and β phases. Weak niobium peaks were also detected, while tin was not detected. Elemental analysis using EDS was performed concurrently with the SEM examination to identify the chemical composition of the samples. The EDS analyses verified that the alloy composition corresponded to Ti14Nb4Sn.

The pore connectivity, which can be determined by percolation theory, is a crucial param‐ eter that determines successful bone in-growth. Connectivity between the pore provides sufficient area for physiological fluid to flow throughout the new tissue that enhances nu‐ trient transportation. The images in Figure 6 exhibit variation in pores connectivity. High porosity results in high pore interconnectivity, *i*.*e*., samples with 80% porosity exhibit high

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**Figure 6.** Morphology of porous Ti14Nb4Sn alloys with different porosity: (a) 55%, (b) 60%, (c) 70%, (d) 72%, (e)

The application of SiO2 as a bond layer between the substrate and the coating should im‐ prove coating adhesion to the substrate. One advantage of using silica is its influence on the bone mineralization process. Li *et al*. applied silica onto a titanium surface using a sol gel process and demonstrated its bioactivity [11]. Hong *et al.* [4] conducted an *in vitro* bioactivity test on bioactive ceramic glass with higher silicon content and revealed a superior minerali‐ zation capability. Thian *et al*. [46] succeeded in incorporating silicon into hydroxyapatite (Si-HA) using magnetron sputtering, and reported that higher silica content was beneficial for

Figure 7 shows the surface topography of the HA-silica coating on titanium alloy Ti14Nb4Sn. The 2 µm thick hydroxyapatite coating and 200 nm thick SiO2 film were deposit‐ ed onto the titanium alloy using RF magnetron sputtering and e-beam evaporation, respec‐ tively. The HA coating was homogenous, which is characteristic of thin films deposited by sputtering. However, some cracks on the surface were observed. Some morphological fea‐ tures of rough coatings with some cracks could be advantageous for bone implant applica‐

**7.2. Physico-chemical properties of sputtered hydroxyapatite coated titanium alloys**

biomedical applications due to its higher corrosion resistance.

tions since this morphology could act as an anchorage for tissue growth.

pore connectivity.

75%, and (f) 80%

**Figure 5.** XRD pattern of sintered Ti14Nb4Sn alloy

SEM images of fabricated porous titanium alloys showed a combination of both macropores and micropores on the surface, as shown in Figure 6 (a)-(f). The micropore size ranged from 0.5 to 10 µm, while the macropore size ranged from 50 to 700 µm. Samples with greater po‐ rosity exhibited more interconnected features, more accessible inner surfaces, and interpene‐ trated macropores. It is believed that the optimal pore size to ensure vascularization and bone in-growth is 50-400 µm [13]. Compared to other studies, fabrication of Ti10Nb10Zr al‐ loy resulted in pore sizes ranging from 300 to 800 µm since the size of the space-holder par‐ ticles was set to be 500-800 µm [13].

Usually there are two types of pores when using the space-holder method to fabricate titani‐ um alloys: (i) macro-pores determined by the size of the space holder particles, and (ii) mi‐ cro-pores determined by the dimension of the titanium powder particles. The micropores can be designed to allow the scaffold to be impregnated with functional coatings or thera‐ peutic agents.

Porosity enhances the interlocking processes for the stability and immobility of the new im‐ plant, often referred to as stabilization and fixation of the implant. The porosity is influenced by several factors, namely the particle size of the metallic powder and the sintering pressure [13]. The porosity of the samples ranges from 55 to 80%. The optimum porosity of the im‐ plant for bone in-growth is in the range of 50-90%. It has been noted that the porosity level of an implant should be selected to provide the optimum mechanical behaviour, since po‐ rosity has a dominant and adverse influence on the strength of a porous material.

The pore connectivity, which can be determined by percolation theory, is a crucial param‐ eter that determines successful bone in-growth. Connectivity between the pore provides sufficient area for physiological fluid to flow throughout the new tissue that enhances nu‐ trient transportation. The images in Figure 6 exhibit variation in pores connectivity. High porosity results in high pore interconnectivity, *i*.*e*., samples with 80% porosity exhibit high pore connectivity.

identify the chemical composition of the samples. The EDS analyses verified that the alloy

SEM images of fabricated porous titanium alloys showed a combination of both macropores and micropores on the surface, as shown in Figure 6 (a)-(f). The micropore size ranged from 0.5 to 10 µm, while the macropore size ranged from 50 to 700 µm. Samples with greater po‐ rosity exhibited more interconnected features, more accessible inner surfaces, and interpene‐ trated macropores. It is believed that the optimal pore size to ensure vascularization and bone in-growth is 50-400 µm [13]. Compared to other studies, fabrication of Ti10Nb10Zr al‐ loy resulted in pore sizes ranging from 300 to 800 µm since the size of the space-holder par‐

Usually there are two types of pores when using the space-holder method to fabricate titani‐ um alloys: (i) macro-pores determined by the size of the space holder particles, and (ii) mi‐ cro-pores determined by the dimension of the titanium powder particles. The micropores can be designed to allow the scaffold to be impregnated with functional coatings or thera‐

Porosity enhances the interlocking processes for the stability and immobility of the new im‐ plant, often referred to as stabilization and fixation of the implant. The porosity is influenced by several factors, namely the particle size of the metallic powder and the sintering pressure [13]. The porosity of the samples ranges from 55 to 80%. The optimum porosity of the im‐ plant for bone in-growth is in the range of 50-90%. It has been noted that the porosity level of an implant should be selected to provide the optimum mechanical behaviour, since po‐

rosity has a dominant and adverse influence on the strength of a porous material.

composition corresponded to Ti14Nb4Sn.

36 Titanium Alloys - Advances in Properties Control

**Figure 5.** XRD pattern of sintered Ti14Nb4Sn alloy

ticles was set to be 500-800 µm [13].

peutic agents.

**Figure 6.** Morphology of porous Ti14Nb4Sn alloys with different porosity: (a) 55%, (b) 60%, (c) 70%, (d) 72%, (e) 75%, and (f) 80%

#### **7.2. Physico-chemical properties of sputtered hydroxyapatite coated titanium alloys**

The application of SiO2 as a bond layer between the substrate and the coating should im‐ prove coating adhesion to the substrate. One advantage of using silica is its influence on the bone mineralization process. Li *et al*. applied silica onto a titanium surface using a sol gel process and demonstrated its bioactivity [11]. Hong *et al.* [4] conducted an *in vitro* bioactivity test on bioactive ceramic glass with higher silicon content and revealed a superior minerali‐ zation capability. Thian *et al*. [46] succeeded in incorporating silicon into hydroxyapatite (Si-HA) using magnetron sputtering, and reported that higher silica content was beneficial for biomedical applications due to its higher corrosion resistance.

Figure 7 shows the surface topography of the HA-silica coating on titanium alloy Ti14Nb4Sn. The 2 µm thick hydroxyapatite coating and 200 nm thick SiO2 film were deposit‐ ed onto the titanium alloy using RF magnetron sputtering and e-beam evaporation, respec‐ tively. The HA coating was homogenous, which is characteristic of thin films deposited by sputtering. However, some cracks on the surface were observed. Some morphological fea‐ tures of rough coatings with some cracks could be advantageous for bone implant applica‐ tions since this morphology could act as an anchorage for tissue growth.

The titanium alloys are likely to be oxidized during the annealing process. Therefore, TiO2 appeared in the CaO.SiO2.TiO2 phase and CaTiO3 phase. The CaTiO3 peak was detected at 47.8° with an orientation of (800). The four peaks at 38.8°, 39.7°, 40.5° and 48.9° correspond‐ ing to calcium phosphate (Ca2P2O7) are indexed as the reflection planes (222), (223), (301) and (320), respectively. However this phase might have higher solubility compared to HA. It is possible that during the sputtering process not all components of the HA target were sputtered and transferred onto the substrate. The titanium peak was present at 53.5° and in‐ dexed as (102). The results indicated that HA coatings using magnetron sputtering could

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39

This chapter describes the importance of developing a bioactive titanium alloy scaffold for bone tissue engineering applications. Ti14Nb4Sn alloy was designed and then fabricated us‐ ing powder metallurgy method. The porosity ranged from 55 to 80% with pore sizes of

Powder metallurgy that employed the space-holder sintering method was successful in fabricating samples for biomedical implant studies. The method produced porous struc‐ tures that (i) enable better fixation, (ii) lower elastic modulus to match the properties of natural bone, and (iii) construct morphologies that mimic the features of natural bone

To further enhance the biocompatibility of titanium alloys, 2 µm thick hydroxyapatite and 200 nm thick SiO2 coatings were deposited onto Ti alloys using e-beam evaporation and RF magnetron sputtering. SEM images showed that the microstructure of the hydroxyapatite coating is homogenous, with some cracks appearing on its surface. XRD results confirmed that the coatings consisted of an HA phase with some CaO.SiO2.TiO2, CaTiO3 and phases. Silica was also present in the XRD spectrum, which corresponds to the CaO.SiO2.TiO2 phase. It was demonstrated that the e-beam evaporation and magnetron sputtering methods are suitable for depositing silica and hydroxyapatite coatings. The hydroxyapatite-silica config‐ uration may be useful for biomedical implants, as it provides better adhesion strength for rapid osseointegration acceleration. Further study will focus on the biological response of

CW acknowledges the financial support from the Australian Research Council (ARC)

produce the crystalline apatite phase.

**8. Conclusions**

100-600 µm.

structures.

these coatings.

**Acknowledgements**

through the ARC Discovery Project DP110101974.

**Figure 7.** Morphology of HA-SiO2 coated Ti14Nb4Sn alloy

The XRD pattern of the HA-SiO2 coated titanium alloys is shown in Figure 8. The identified phases were hydroxyapatite, CaO.SiO2.TiO2, calcium pyrophosphate, CaTiO3 and titanium. After annealing, the crystalline phase of HA was present at 2*θ* = 30° which matches the (107) plane. A peak corresponding to CaO.SiO2.TiO2 phase was also observed at 43.5° and in‐ dexed as (223). In addition, the peak confirms the presence of the silica phase. The phase CaO, *i.e.,* in CaO.SiO2.TiO2 observed in the XRD pattern could be related to the partial de‐ composition of hydroxyapatite during the deposition process.

**Figure 8.** XRD patterns of HA-SiO2 coatings on Ti14Nb4Sn alloy

The titanium alloys are likely to be oxidized during the annealing process. Therefore, TiO2 appeared in the CaO.SiO2.TiO2 phase and CaTiO3 phase. The CaTiO3 peak was detected at 47.8° with an orientation of (800). The four peaks at 38.8°, 39.7°, 40.5° and 48.9° correspond‐ ing to calcium phosphate (Ca2P2O7) are indexed as the reflection planes (222), (223), (301) and (320), respectively. However this phase might have higher solubility compared to HA. It is possible that during the sputtering process not all components of the HA target were sputtered and transferred onto the substrate. The titanium peak was present at 53.5° and in‐ dexed as (102). The results indicated that HA coatings using magnetron sputtering could produce the crystalline apatite phase.
