## 4 Discussion
### Model Validation  
Results of the current study were compared with available in vitro and in silico data from the literature. [PERSON] et al. described that the validation via diverse results increase the accuracy and reliability of the developed models' predictions [2]. Therefore, validation was performed in the current study utilising several published experimental and in silico results. The LBM and PSM yielded similar ROM results against pure bending moment of 7.5 Nm indicating that the material assignment of the bony structures slightly influences the spinal kinematics of the spine. Similar findings were described by [PERSON] et al., stating that for FE simulations, the material properties of the bony structures have a minor influence on the spinal kinematics; thus, vertebral bodies could be modelled as rigid bodies [79]. Moreover, [PERSON] et al. presented, calibrated, and validated a hybrid finite element-multibody model with rigid vertebral bodies, indicating that the hybrid simulation model is powerful and efficient for providing valid mechanical responses. Their study confirmed that the soft tissues such as the spinal ligaments and the IVD have primarily influence the kinematic properties of the spine [29]. Thus, nonlinear modelling of such structures is crucial, and therefore in the current study, the spinal ligaments and components of the intervertebral discs were modelled with nonlinear materials. The two models were in a good agreement with the published in silico and in vitro studies in L1-L5 ROM of flexion-extension, lateral bending, and axial rotation. However, the nonlinear load-displacement curve (Figure 3b) demonstrates that in flexion, both the patient-specific and literature-based models are slightly outside the in vitro range, while in extension, the LBM is within, the PSM is slightly outside the range. The same trend is reported by [PERSON] et al. [2], that majority of their assessed FE models overestimate the motion in extension and underestimate it in flexion, moreover, good agreement was found with the reported in silico results for lateral bending and axial rotation as well [2]. The FJF depends on several patient-specific factors, such as the underlying geometry of the individual cartilage surfaces, the gap size between the adjacent surfaces, or the thickness of the cartilage layer [80]. There is limited data available in the literature on the in vitro measurement of FJF, and these studies have some limitations and uncertainties [51]. Nevertheless, in pure bending, the results of the current study show good agreement with the available in vitro [58] and in silico [2] data, although a slight overestimation of the in vitro result appeared in extension (Figure 3c). The current study results indicate that the FJF in lateral bending is notably lower than in extension and axial rotation, which may originate from the patient-specific anatomy of the current study's FE model agreeing with the findings of [PERSON] et al. [80]. In flexion, no force was transmitted due to the distancing cartilage surfaces during the load [34]. The IDP predictions of the FE models in pure compressive follower load showed excellent agreement with the median in vitro results at 300 N and 1000 N (Figure 3d). Among the published FE models in the literature, the NP is often modelled as incompressible fluid-filled cavity [25; 49], which allows simulating the effect of the initial hydrostatic pressure in the IVD [81]. Nonetheless, as [PERSON] et al. pointed out, most NPs modelled with fluid-filled cavities neglect this effect since IDP values measured in the unloaded state are less significant [2]. The IVR values were obtained and evaluated for both pure and combined load cases (Figure 4a-h). [PERSON] et al. previously reported that 7.5 Nm pure bending moment is the most appropriate for simulating the loads on the lumbar section, including effects of muscle forces [48]. [PERSON] et al. performed a human cadaveric test [60], while [PERSON] et al.