## 4 Discussion
Chewing forces are lower in conventional complete dentures and vary from 30 to 50 N [46], but are much higher in subjects with their own teeth or fixed partial dentures [44]. The differences between the results of this study and other studies on biomechanical behavior [37; 38; 47; 48; 49; 50; 51; 52; 53; 54] may be due to different mini-implant materials (Ti-Zr vs. Ti-Al6-V4), different retention mechanisms (Optiloc(r) Retentive System vs. \"O\"-ring), different implant designs and neck widths [55], different dimensions of implants [56; 57], small differences in strain measurements and different residual ridge forms. We bonded strain gauges as close as possible to the implant-bone interface due to the assumption that stress and strains would be highest where two materials with different stiffnesses meet. The cyanoacrylate glue was recommended by the strain gauge manufacturer and was also used in other studies utilizing SGs. Some experimental results indicate that SGs used with the same adhesive gave consistent deformation values, while different glues (cyanoacrylate, epoxy) led to different values under the same conditions [58]. Therefore, it is important to use the same glue as in other similar studies to be able to compare the results. Moreover, the adhesive thickness should be the lowest possible and should be uniformly applied. A linear relationship between adhesive thickness and strain error has been found [58]. For this reason, we used acetate foil to firmly press strain gauges against the model during the glue setting to have the thinnest possible glue layer. One of the limitations of this study is that we could not control the glue thickness completely. The new Ti-Zr (Roxolid(r)) Mini-Implant system utilizes binary titanium alloy with approximately 15% Zr, which is an \(\alpha\) alloy with a closely packed hexagonal crystallographic structure and represents a completely solid solution [35; 36; 59], contrary to Ti-6 Al-4V trinary alloy, which consists of \(\alpha\) and \(\beta\) phases (closely packed hexagonal and body-centered cubic crystals) [59; 60]. A solid solution of a Ti-Zr alloy allows better and faster cell adhesion and faster and better osseeintegration. There is no degradation of potentially cytotoxic vanadium. However, both alloys have similar mechanical properties [59; 60]. An elastic modulus of the Roxolid(r) alloy is reported to be 96.12 \(\pm\) 2.82 GPa, significantly lower than that of titanium [60], while according to [PERSON] et al. [35; 36] it varies between 102 and 104.7 GPa, with a Poisson coefficient of 0.33. The Roxolid(r) alloy has similar elastic characteristics to the Ti-6 Al-4V alloy (110-115 GPa) but with a higher tensile strength [35; 36]. The cortical bone modulus of elasticity is only 15 GPa [35; 36]. The hardness value of Ti by nanoindentation was 2.38 \(\pm\) 0.13 GPa and that of Ti-Zr Roxolid(r) alloy was 3.19 \(\pm\) 0.09 GPa [60]. [PERSON] et al. [61] reported the yield strength of the Roxolid(r) alloy to be 799 \(\pm\) 26 MPa, and the ultimate tensile strength to be 968 \(\pm\) 2.6, higher than that of grade four Ti, but similar to that of the Ti-6 Al-4V alloy (970 MPa). The data on the mechanical properties of these alloys show a certain amount of variability, as reported in different studies [62]. All microstrains reported in this study represented the peak (maximum) microstrains during the two-second OD loading interval after the desired loading force had been achieved. The maximum strain values during the 2 s loading period were chosen because the applied force varied by a very small amount during the 2 s. Additionally, repeated strains can interfere with bone reparatory mechanisms even in the range of 1500-3000, which corresponds to mild overloads [33]. However, none of the maximum microstrains registered in the present study exceeded a value of 3000, when bone reparatory mechanisms can be jeopardized.