## 4 Discussion
However, in the experimental study [15], vertical occlusal loads were applied, and failure occurred in the lower fixation screws at a load range of 1678-2163 N. This experiment showed that the system's load-bearing capacity was considerably higher than the numerical predictions made in study [15]. The experiment applied occlusal forces to rigidly supported, significantly shortened mandibular stumps, altering bending conditions by changing the bone length and the lever arm of the occlusal force relative to joint support. Both the FEM model and the experiment in [15] involved fixation at both temporomandibular joints, which differs significantly from real-world conditions. Despite this, the stress analysis results indicated that bone failure would occur before implant failure. Furthermore, the experimental validation in study [15] confirmed that failure tends to occur in the bone screw regions. In our study, bone tissue stresses also approached the critical threshold of 140 MPa but did not exceed the cortical bone strength of the mandible [13]. However, it can be assumed that patients do not fully regain muscular function, and occlusal forces sufficient for food mastication range between 250 and 300 N. In clinical practice, overload-induced bone resorption occurs around bone screws, which aligns with stress values when occlusal forces in the model are reduced by half. However, the problem addressed in this study was the risk assessment of implant fracture, which occurs in clinical practice [10; 14; 15; 16; 17; 18]. An example is the study by [PERSON] et al. [15], where 40 out of 187 implanted lattice implants experienced clinical fractures. The stress in the designed implant in one of the lattice bars was below 80 MPa, significantly lower than the permissible values for titanium obtained via the selective laser melting (SLM) method. The authors attempted to justify fractures by stress concentration in a single finite element, which, however, suggests an FEM artifact due to finite element mesh. The failure was explained by the lower actual cross-sectional strength compared to the model due to the presence of pores inherent to the SLM technology. Our results indicate that in preclinical numerical verification of mandibular implant load capacity, occlusal forces cannot act solely in a vertical direction, and temporomandibular joint support should not be symmetrical [14; 16; 17; 19]. Our objective is not to engage in extensive discussions with every study in this domain, as their number is vast, but rather to highlight the implications from studies [3; 12]. In summary, it should be emphasized that in simulation models for the study of mandibular prosthesis load-bearing capacity, the most important criteria are the method of model support and muscle activity. In our research, the assumption of perfect fusion had a secondary impact on the result. There is general agreement that the decisive factor in load-bearing studies is the maximum muscular effort that generates the maximum bite force. However, our studies, in comparison with others, indicate that this is not the only criterion. According to the fundamentals of mechanics, it is known that internal forces (bending moments, torsion, and shear) depend on the compliance of the supports. Various approaches to mandibular support during the load-bearing capacity studies of mandibular prostheses can be distinguished. The first approach involves neglecting the active forces of the muscles, and instead of generating reactions in the joint, the condylar head is directly loaded with an active force until the implant fractures, with the mandible rigidly fixed in an arbitrarily chosen cross-section. This approach is suitable for comparing load-bearing capacities between solutions but overestimates the load because the mandible is not a cantilever support rigidly fixed on one side. The second approach involves obtaining bending through active forces on the mandibular frame supported on the tooth and in the joints.