*Magnesium in Synthesis of Porous and Biofunctionalized Metallic Materials DOI: http://dx.doi.org/10.5772/intechopen.102083*

230 GPa for Co-Cr alloy [15]. From the metallic biomaterials, the near β-Ti alloys have reported some of the lowest elastic modulus values ranging from 80 to 110 GPa [14]. Additionally, porous β-Ti alloys have obtained elastic modulus ranging from 43 to 75 GPa [4, 37, 73, 74], and it is expected that the addition of Mg will decrease even more such values [75]. An elastic modulus near to that of human bone improves the performance of multiple biomaterials used for dental and orthopedic implants. Considering that the microstructural effects of Mg additions are important to define the mechanical behavior of biomaterials, it is important to study the microstructure.

Due to the above, the effect of low (3 mass%) Mg additions into a Ti-34Nb-6Sn alloy was studied [28]. The selection of Nb and Sn was done for its β-stabilizing effect on Ti alloys [9, 76]. The β-phase has shown better mechanical compatibility with the human bone in comparison to the α-phase. This is due to lower elastic modulus values and a high strength-to-weight ratio [13, 14]. An elastic modulus of the implant that is close to that of the human bone decreases the mismatch of mechanical stress through the interface and avoids the damage of the organic tissue cells. Based on the above, the closer the elastic modulus between both, bone and implant material, the lower the probability of crack nucleation and failure of the implant [72]. Furthermore, a high strength-to-weight ratio allows reducing the thickness of biomedical metallic implants. Besides, the selected contents of Nb and Sn also contribute to obtaining elastic modulus of the Ti alloy of ~60 GPa [9, 74, 77]. The addition of 6 wt.% of Sn into the Ti-Nb alloy showed a good combination of corrosion resistance, strength, hardness, and lower elastic modulus [4]. On the other hand, considering the low elastic modulus of Mg (from 39 to 46 GPa) [78], it was added to reduce the elastic modulus of the Ti alloy. Besides, Mg is a natural component of human bones and is a required element for the metabolism process, i.e., Mg exhibits great biocompatibility, non-toxicity, and can stimulate hard tissue recovery [2]. The reported biodegradability of Mg2+ is one more of its valuable advantages [79]. The possibilities of its use in dental and orthopedic implants can be highly beneficial from controlling its degradation rate and ensuring its mechanical integrity during desire clinical periods. Moreover, the corrosion products of Mg, trigger the osteoconductivity of the bone [75]. For the last, the β-phase Ti alloys have shown a superior electrochemical performance that provides better resistance in corrosive environments as oral or body fluids [13, 32]. This is due to the surface TiO2 passive film that inherently protects these alloys [10, 32].

From the above-mentioned selection of components, four Ti alloys were prepared by powder metallurgy method. A typical four-stage route of milling – mixing - compaction – sintering was used. For this, measured amounts of titanium hydride (TiH), niobium hydride (NbH), and atomized Sn were used to obtain a mass ratio of Ti, Nb, and Sn of 60:34:6. The powders were mixed in a planetary ball mill and grounded at 200 rpm for 40 min. Posteriorly, the mixed powders were dried under a vacuum. Details of processing parameters can be found in previous work [28]. Half of the mixed powder was saved with the abovementioned chemical composition, while the other half was mixed with Mg powder in a 3 mass%. All the dried powders, whether with or without Mg, were compacted at 200 MPa for 15 s. To evaluate the microstructural and mechanical effect of Mg addition, two different sintering temperatures were used, 900°C and 1100°C. The sintering was carried out in a high vacuum resistive furnace for 2 h. A scheme of the elaboration process is shown in **Figure 1**.

Finally, two Ti-34Nb-6Sn (TNS) and two Ti-34Nb-6Sn/Mg (TNS/M) alloys were produced. The identification of the obtained alloys indicates the sintering temperature as postfix: TNS900, TNS1100, TNS/M900, and TNS/M1100.

**Figure 1.** *Representation of the methodology to elaborate the Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys [28].*

For microstructural analysis, the samples were subjected to conventional metallographic preparation until a mirror-like surface. Final polishing with oxide polishing suspension (OPS) solution and hydrogen peroxide (10:2) was applied. Rietveld refinement of X-ray diffraction measurements was carried out for quantifying the present phases and estimating the lattice parameters. A Bruker/D2Phaser with Cu-Kα radiation was used at 30 kV and 10 mA. The measured 2θ range was 20 and 90° with a step size of 0.02° every 10 s. The Rietveld refinement was carried out by MAUD software (version 2.94) [80]. The morphology and chemical distribution of phases were studied by field emission scanning electron microscopy (FESEM) (ZEISS-ULTRA 55) and an energy-dispersive X-ray spectroscopy detector (EDS) (Oxford Instruments Ltda.).

Considering that the mechanical properties play a determining role in the performance of biomaterials, the elastic modulus was estimated by impulse excitation technique (ATCP, Sonelastic®). Hardness measurements were obtained using a load of 147 N by the Rockwell method (BECLA), using a spherical steel indenter with a diameter of 0.16 cm. More details of the whole methodology can be found in previous works [28–31].

As result, the four alloys resulted in tri-phasic microstructures of α-Ti (under hexagonal compact (hcp), structure), β-Ti (under body centered cubic (bcc) structure), and segregation of Nb. The microstructures can be observed in **Figure 2**, where the light gray color corresponds to the matrix of β-phase, the dark gray to the α-phase, and the bright particles to the Nb segregation. As it was expected from the increment of temperature, bigger grain sizes can be observed in the samples sintered at 1100°C (**Figure 2c** and **d**) compared to those sintered at 900°C. Both, α and β-phases, are randomly distributed in the microstructure, however, the linear chemical composition through the microstructure is not homogeneous. This is due to lower contents of Nb and Sn, especially Nb, in the α-phase (**Figure 2e**). Due to the well-known β-stabilizer nature of both alloying elements [1, 3–10], those chemical gradients were expected. On the other hand, the Nb segregation occurrence was reduced with the sintering temperature (**Figure 2c** and **d**), which indicates a better Nb diffusion in the matrix when temperature increases. However, the Nb particles are continuously observed at both sintered temperatures. From the Ti-Nb diagram phase, Nb has low solubility in Ti, so the continuous presence of Nb segregates was expected [3].

The phases percentages estimated by Rietveld refinement from XRD measurements and the total porosity obtained by Archimedes method for the four alloys are presented in **Table 1**. The TNS900, TNS1100, and TNS/M1100 samples showed similar phases percentages. However, the TNS/M900 showed a reduced β to α transformation during sintering. Additionally, both TNS/M samples obtained higher porosity

*Magnesium in Synthesis of Porous and Biofunctionalized Metallic Materials DOI: http://dx.doi.org/10.5772/intechopen.102083*

### **Figure 2.**

*Microstructure of a) TNS900, b) TNS/M900, c) TNS1100, and d) TNS/M1100, as well as e) linear chemical gradients through α and β phases representative of the four studied samples. Adapted from [28].*


### **Table 1.**

*Phases percentages and total porosity of TNS and TNS/M samples sintered at 900°C and 1100°C.*

percentages in comparison with the Mg-free alloys. This was a clear suggestion about the reduction of diffusional processes when Mg is added into the Ti-Nb-Sn system. Thus, Mg addition has an apparent α-phase stabilization effect.

Furthermore, the increment of porosity with the Mg content could be related to the highest content of oxygen from the intrinsic passivation layer of Mg2+. When temperatures increase, the release of gas also increases, generating pores at the microstructure [81]. Additionally, the Mg powders acted as a spacer in the TNS/M samples. Comparing the low melting point of Mg (~650°C) with that of Ti (~1668°C), it is evident that a fraction of Mg is evaporated during sintering, while the Ti content remains constant. The partial evaporation of Mg assisted in the formation of the pores during sintering. It is well known that the porosity tends to decrease at higher sintering temperatures during powder metallurgy methods [82]. Thus, the reduced porosity for the samples sintered at 1100°C compared to the samples sintered at 900°C was an expected result. The porosity could influence the mechanical properties of the alloy, especially in the strength and elastic modulus. The mechanical properties will be discussed in the next Section 3.

Posteriorly, the previous Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys were also reported through the same powder metallurgy methodology, except for sintering temperatures of 700°C and 800°C [29–31]. These samples will be identified as TNS700, TNS/M700, TNS800, and TNS/M800 for this book chapter. Compared to the previous TNS/M900, and TNS/M1100, the same three constitutive phases, α, β and Nb segregation, were observed at the TNS and TNS/M alloys sintered at 700°C and 800°C. For

comparison purposes, representative EDS measurements of TNS/M800 and TNS/M900 are shown in **Figure 3**. **Figure 3a** is representative of the distribution of the elements in the samples sintered at 700°C and 800°C, while **Figure 3b** represents the distribution of the elements in samples sintered at 900°C and 1100°C. Similar distributions of the alloying elements were observed in both cases, except for the Nb segregates. **Figure 3a** shows a greater presence of Nb segregates in comparison with that representative of sintering above 900°C (**Figure 3b**). This can be explained by the lower solubility of Nb in Ti below 1100 K (~827°C) [3]. The solubility, together with the effect of temperature, could also be related to the smaller grain size in the samples sintered below 900°C (**Figures 2** and **3**). While the β-phase matrix of the TNS1100 has an average grain size of 15 μm, that of the TNS700 has an average of 6 μm. This could be due to the lower solubility of Nb at lower sintering temperatures generated more segregated particles through the microstructure. Those particles could act as a pin for grain growth. The pin-like behavior of Nb segregates in a Ti matrix has been reported before [16].

From **Table 2**, the TNS700, TNS/M700, TNS800, and TNS/M800 samples presented higher porosity percentages in comparison with the samples sintered at 900°C and 1100°C (**Table 1**). Besides, the samples sintered at 800°C were more compacted than the samples sintered at 700°C. This was a confirmation about the higher the sintering temperature, the lower the porosity percentage. It is well-known that increasing the sintering temperature favors the diffusional processes and more compacted microstructures with higher relative densities, i.e., lower porosity percentages [82]. Besides, **Table 1** and **Table 2** showed an increment of porosity with the Mg additions for samples sintered at the same temperature. As it was explained before, the partial evaporation of Mg and its passivating oxide, contributed to the increment of porosity. Among the most studied spacers for powder metallurgy methods are carbamide, sodium chloride, ammonium hydrogen carbonate, and Mg [83]. From those, Mg has shown superior advantages over the organic spacers due to its good biocompatibility

### **Figure 3.**

*Comparison between alloying elements distribution of the Ti-34Nb-6Sn/Mg alloy representative of a) the sintered at 700°C and 800°C and b) the sintered at 900°C and 1100°C. adapted from [28, 30].*

*Magnesium in Synthesis of Porous and Biofunctionalized Metallic Materials DOI: http://dx.doi.org/10.5772/intechopen.102083*


### **Table 2.**

*Phases percentages and total porosity of TNS and TNS/M samples sintered at 700 °C and 800 °C.*

and good mechanical properties [79]. The increment of pore formation with the Mg additions could be beneficial for the implant by decreasing the elastic modulus. This topic will be covered in Section 3.

Besides, from **Figure 3**, a larger pore size for the samples sintered below 800°C is notable in comparison with that for the samples sintered above 900°C. Macropores of ~100 μm in average were acquired (**Figure 3a**) for the samples sintered below 800°C. This was contrasting with the pores in the range from 5 to 35 μm obtained in the samples sintered above 900°C. This could be related to lower atomic diffusion resulting from the lower sintering temperature that reduced the bonding ratio in the samples. Porosity improves the bonding between bone and implant material, encouraging the anchorage and growth of the organic tissue [15, 75]. Thus, porous materials facilitate tissue generation enabling body fluid transmission [84]. Considering that allowing successful osseointegration is one of the main requirements of dental and orthopedic implant materials, the porous structures are highly promising for those applications. It has also been reported that macro-pores are beneficial for multiple biological processes as cell attachment, ingrowth of osteoblasts, vascularization, and osteoconductivity [85]. However, an adequate vascularization requires pores with diameters larger than 100 μm, specifically in the range from 100 to 500 μm [15, 86]. From the above and the fact that increments of porosity percentage used to assist on the decrement of elastic modulus [75], the alloys sintered at temperatures below 800°C could be more appropriate for dental or orthopedic implant applications. However, other mechanical properties as strength and hardness are also crucial for the performance of metallic implants. The mechanical behavior will be described in the following Section 5. This will contribute to clarifying the current concerns about the role of porosity *in-vivo* environments.
