## 4 Discussion
Additionally, we observed that an increased head size allowed more ROM during flexion and internal rotation, which was shown to extend the range for orienting the acetabular cup [41]; this is in agreement with previous studies [21, 25, 30, 41, 43]. The increase in ROM owing to the higher head size can be explained with the higher head-to-neck ratio, which causes delayed contact between the femoral neck and the acetabular cup [72]. We observed a similar tendency for internal rotation as [PERSON] et al. [21], where increasing head diameter led to more ROM. It is also noteworthy that increasing the head size would decrease the contact stresses at the liner during subluxation as well as lower component damage and the risk of dislocation [21, 25, 41], which is supported by clinical experience [73, 74]. It has to be taken into account that increasing the head size not always increased ROM. The reached ROM is also dependent on the femoral neck configuration, which can lead to the unexpected event of reduced ROM despite larger head sizes for very short neck length, caused by loss of femoral offset [49]. On the other hand, we observed higher ROM for an increased head offset, which was also observed by other research groups [25, 41, 49]. Our findings confirm this well-known correlation as well as the clinical experience that increasing the offset reduces the risk for dislocation in THR [18, 41, 52, 53, 75]. Furthermore, this is also in good agreement with the study by [PERSON] et al. [52], who reported an improved flexion of more than 20\({}^{\circ}\) and an improved internal rotation of more than 13\({}^{\circ}\), which corresponds well with our data (19.8\({}^{\circ}\) for flexion and 11.9\({}^{\circ}\) for internal rotation). A shorter neck, caused by a reduced head offset, is known to result in earlier contact between the proximal stem built-up and the liner component. Therefore, increasing the head offset is an option for the surgeon to mitigate issues with insufficient ROM intraoperatively, but this also causes an increased leg length and change in soft-tissue condition [25, 49]. Additionally, the femoral offset can also change the strain/stress pattern in the surrounding bone. Despite the progress in methodology for computational ROM analysis, our approach has some limitations. Firstly, the geometry generation is time-consuming and unacceptable in the clinical setting in its current state; however, this can be mitigated by using more advanced techniques in medical imaging data or with statistical shape models, which turned out as feasible for femoral and acetabular bones [67, 68, 76, 77]. Additionally, due to the rather limited resolution of MRI, segmentation, and deviations because of surface triangulation, a difference between the bony geometry and the derived model is unavoidable, which can influence the ROM, especially when bone-to-bone impingement or implant-to-bone impingement occurs. However, current methods for shape modeling [67, 68, 76] can be used in future studies to also mitigate discretization, but this may be associated with a less accurate representation of the patient's bone geometry, leading to less reliable results in terms of bone-to-bone or bone-to-implant impingement. This and the more favorable time effectivity gained by more sophisticated approaches in obtaining patient individual bone geometry will make the proposed framework more attractive as a preoperative analysis tool. Secondly, due to the lack of soft-tissue structures in the presented numerical model, the real ROM is likely to be smaller [11, 36, 47]. However, this is a typical assumption for studies investigating the ROM [18, 21, 32, 33, 64, 78]. Due to the fact that physiological ROM was achieved during flexion and internal rotation movements, we can assume that this limitation only affects the results for external rotation, where the detected ROM exceeded the physiological ROM.