**5. Mechanical behavior of Ti-Mg alloys**

It is well known that the microstructural features have a strong effect on the mechanical behavior of metallic components. In this section, the mechanical performance of the Ti-34Nb-6Sn and Ti-34Nb-6Sn/Mg alloys will be explained based on their microstructural features described in Section 4.

As it was explained before, the Mg additions into dental or orthopedic biomaterials should assist the adequate biocompatibility and a good mechanical strengthelastic modulus relationship. Lower elastic modulus is expected with Mg additions. Moreover, lower elastic moduli are expected from higher porosity percentages, this is, for lower sintering temperatures. Simultaneously, higher hardness values can be expected from the more compacted samples, i.e., the less porous. To evaluate these hypotheses, hardness and elastic modulus measurements were carried out.

From previous works [28–30], **Table 3** presents the Vickers hardness as a function of sintering temperature for the TNS and TNS/M samples. A general tendency to increase hardness as a function of sintering temperature was observed. This is congruent with the lower total porosity percentage in the samples sintered at higher temperatures. As it was explained in Section 2, the temperature increment fabrication atomic diffusional processes that result in bonding improvement and density increment. To observe the effect of temperature on the total porosity and hardness of the sintered alloys, **Figure 4** compares those three values. A general decrement of porosity with the increment of temperature can be observed in the TNS (**Figure 4a**) and TNS/M (**Figure 4b**) samples. Simultaneously, the higher density of the samples resulted in higher hardness values for all the samples.

For comparison purposes, the hardness of elemental Ti ranges from 1.3 to 2.0 GPa [16], while the hardness in the TNS and TNS/M samples ranges from 0.9 to 4.0 GPa. Besides, the minimum recommended hardness value for metallic biomaterials is 1.2 GPa for avoiding high wear damage susceptibility during chewing and daily oral processes [87]. However, the hardness of the natural human teeth ranges from 2.2 to 3.9 GPa [88]. Hardness values near to the ones of natural teeth could assist in decreasing wearing between teeth and implant. From the sintered TNS and TNS/M alloys, the TNS700, TNS800, TNS/M700, and TNS/M800 have hardness values below 1.5 GPa, this is, below the minimum acceptable for avoiding wear damage and being within the range of the hardness of natural teeth. As result, those four alloys cannot serve as a feasible biomaterial for dental applications. However, the hardness of the human bone ranges from ~0.3 to ~0.75 GPa [89, 90]. This means that all the sintered samples are within the acceptable hardness range for being applied as orthopedic implants.

From **Figure 4b**, it is also possible to observe a slower decrement rate of total porosity with the sintering temperature in comparison with that for the samples free of Mg additions (**Figure 4a**). This could be related to the abovementioned effect of Mg as a spacer. The partial evaporation of Mg created additional pores compared to the created in the Mg-free samples (TNS). As result, the hardness increment with sintering temperature in Mg-added (TNS/M) alloys also showed a slower rate of increment in comparison with the TNS samples.

As it was discussed in previous Section 4, the elastic modulus plays a key role in the success rate of dental and orthopedic implants. Considering that the elastic modulus is a measure of the stiffness of the material, it determines the resistance to


### **Table 3.**

*Comparison of Vickers hardness (HV) between the TNS and TNS/M samples sintered at 700 °C, 800 °C, 900 °C, and 1100 °C.*

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

**Figure 4.** *Hardness and total porosity as a function of sintering temperature for the a) TNS and b) TNS/M samples.*

deform a material in the elastic range. For increasing the feasibility of the implant, the constituent alloy should have an elastic modulus near to that of the human bone. The elastic modulus of human bone ranges from 5 to 30 GPa [15, 91].

For evaluating the elastic modulus of the sintered alloys, **Figure 5** presents a comparison between the values obtained for the TNS and TNS/M samples sintered at different temperatures. Considering that the samples sintered at 700°C and 800°C were discarded as potential materials for dental implant applications, the samples TNS700, TNS800, TNS/M700, and TNS/M800 were not included in **Figure 5**. Lower elastic modulus can be observed in the samples with Mg addition compared to the TNS systems. This result was congruent with the lower hardness values of the TNS/M samples (**Figure 4**), which implies that these alloys are softer than the TNS for similar sintering conditions. As it was explained before, the lower hardness in the TNS/M samples resulted from the spacer-like behavior of the Mg powders. Besides, a tendency to increase the elastic modulus with the sintering temperature was observed. Being congruent with lower porosity percentages measured in the TNS1100 and TNS/M1100 samples.

Comparing the obtained elastic moduli in the studied samples with the ones reported for human bone, the TNS/M900 sample can be the most adequate for its use in dental or orthopedic implants. Besides, the TNS/M900 alloy joins an acceptable hardness for biomedical implants and has adequate porosity features for triggering the anchorage between organic tissue and implant material. This is, the TNS/M900 sample

### **Figure 5.**

*Elastic modulus as function of temperature for samples sintered at 900°C and 1100°C with (TNSM) and without (TNS) Mg addition [28].*

combined the best microstructural and mechanical properties to be a potential biomaterial. However, the performance of these alloys under in vivo environments should be described to determine the feasibility of the studied alloys for biomedical purposes.
