## 4 Discussion
His 3D model was an idealized anatomical model (with a sample size of one), while our models were CBCT- and anatomical-based (with a sample size of nine). The stress display areas for VM non-linear looked vaguely like ours (i.e., the cervical third stress in the tipping, and entire PDL stress for the intrusion), while the linear simulations totally disagreed. However, it must be emphasized that [PERSON] et al. 's [5] stress distribution areas are biomechanically unnatural, and not clinically met (while the present study is both mechanically and clinically feasible). The quantitative reports for the non-linear PDL are under 4.7 KPa of the MHP for 0.25 N and in exceedance for 2.4 N. His conclusion was that only hydrostatic stresses could be used in the FEA of PDLs, in total contradiction with the physical-mechanical engineering field knowledge (as above-mentioned), other FEA reports [1; 21; 22; 23; 24; 25], and the present study. [PERSON] et al. [14], in a model of the lower mandibular arch (incisor, canine, and premolar, with mesh size 1.2 mm, 10 nodes tetrahedral elements, 23,565 elements, 32,812 nodes, 0.35-0.5 N of tipping, VM, S1, hydrostatic stress), reported higher stress for the multiple teeth vs. the single tooth models (i.e., which is biomechanically debatable). The stress display areas (qualitative) were closer to the present study. However, the qualitative color-coded areas were entirely red-orange, signaling a quasi-uniform stress, which is biomechanically unnatural and clinically not true. Nevertheless, the applied 0.5 N of tipping was reported to produce a maximum amount of stress of 235.2-324 KPa, highly exceeding the maximum 12.8-16-26 KPa [5; 6; 7; 8] reported for MHP, signaling high ischemic and resorptive risks for a light force, contradicting both clinical knowledge [20] and the present results. Other two studies, namely [PERSON] et al. [15] (a single simplified model of the 2 nd upper premolar, intrusion) and [PERSON] et al. [16] (a single simplified model of the 1 st lower premolar, intrusion), reported qualitatively comparable but quantitatively higher results. [PERSON] et al. [18] reported the intrusion and extrusion being more stressful than the rotation, translation, and tipping in a VM study, totally contradicting the present study and biomechanical and clinical knowledge [20]. In different VM-based research, [PERSON] et al. [19] reported the tipping to be more stressful than intrusion, in agreement with the results reported herein. [PERSON] et al. [17] reported, for an intact PDL (specifically, an artificially created upper incisor), for intrusion (0.15 N), extrusion (0.3 N), and tipping (0.3 N), smaller amounts of stress than here and maximum stress areas in the apical third (which is biomechanically unnatural for such small forces), contradicting our results. The above-mentioned correlations with available studies [3; 4; 5; 6; 7; 8; 9; 10; 11; 12; 13; 14; 15; 16; 17; 18; 19] clearly show the main shortcomings of applying the FEA method for the study of dental tissue, and why it lacks the accuracy shown in the engineering field. Those reported above are due to significant differences between the analyzed 3D models (i.e., mostly idealized, with low anatomical accuracy due to a high global element size of 1.2 mm and fewer elements/nodes 1674/5205-23,563/32,812, up to 1.67 million, vs. a reduced global element size of 0.08-0.116 mm and a higher number of elements/nodes, specifically 5.06-6.05 million/0.97-1.07 million, in the present study). Most of these studies do not motivate their selection of their employed failure criteria (material type-based), instead employing, comparing, and correlating brittle-like or liquid criteria, which are not suited for the ductile-like nature of dental tissue. Thus, it is expected that the reported stress distribution (qualitatively) varies and displays unnatural or unrealistic clinical results, as proven in our earlier comparative studies [1; 21; 22; 23; 24; 25].