## 3 Results  
The computational procedure presented in the Materials and Methods section is implemented in our FE solver PAK [37], supported by the pre- and post-processing user interface software CAD Solid and Field [38] ([[https://github.com/miljanmilos/CAD-Solid-Field](https://github.com/miljanmilos/CAD-Solid-Field)]([https://github.com/miljanmilos/CAD-Solid-Field](https://github.com/miljanmilos/CAD-Solid-Field)), accessed on 23 October 2023). The two virtual patients used in this study are generated based on the geometric data obtained from two patients with developed HCM and DCM. A simple parametric model of the left ventricle consists of a base part, aortic and mitral valves, and a connecting part (the junction between the base and the valves; see Figure 3a). Each geometric component of the model has a length, radius, and number of layers (division parameter). Two different left ventricle (LV) models are generated as presented in Figure 3a,b: (1) a model with a shorter base length component mimicking the HCM case (Figure 3a), and (2) a model with a longer base length component and thinner lateral wall, mimicking the DCM case (Figure 3b). A fluid domain (coloured blue in Figure 3a,b) is bordered by a solid wall (coloured black in Figure 3a,b) in the FE model. The solid wall is constrained at its top (\"Solid top\" in Figure 3a,b), and the inlet velocity is prescribed at the inlet mitral valve cross-section, while the outlet velocity is prescribed at the aortic valve cross-section. The functions for the prescribed inlet and outlet velocities are shown in Figure 3c. In the LV parametric model, the base radius is 25.31 mm, the base part length is 79.47 mm, the connection part length is 10.11 mm, the mitral valve cross-section radius is 12.1 mm, the mitral branch length is 18.1 mm, the aortic valve cross-section radius is 14.9 mm, the aortic branch length is 18.1 mm, and the solid wall thickness is 10 mm. The mitral and aortic branches have lengths that are larger in dimension in comparison to the thickness of the aortic and mitral valves to prevent possible velocity fluctuations in these zones. The time step used in the FE simulation is 20 \(\times\) 0.05 s.  
Figure 3d shows a one-layered solid representation of the fibres and their orientations. An algorithm is used to calculate the fibre direction automatically, according to [39]. The fibres are helical and their angles vary from 60\({}^{\circ}\) in the endocardium (yellow arrows in Figure 3d) to \(-60^{\circ}\) in the epicardium (red arrows in Figure 3d). A linear distribution of the angle along the thickness of the LV wall is assumed. In the first step of our fibre generation algorithm, the inner (endocardium) and outer (epicardium) surfaces of the LV wall are extracted. At each node of the model, we calculate the distances to the extracted surfaces. Based on the ratio of the distance to the epicardium and the total thickness of the wall, we vary the angle at which the fibre should be generated. Finally, the fibre muscle is generated by rotating the tangential vector to the wall surface by the calculated angle. Here, we present the results of the simulations with two LV geometries and accompanying scenarios for the LV models that closely resemble the HCM and DCM scenarios. Each case addresses the following three possibilities: (1) the effect of Ca\({}^{2+}\) concentration, (2) the effect of the Holzapfel scale factor (wall stiffness), and (3) the effect of the inlet and outlet velocities. Finally, we present the effects on ejection fraction of different wall stiffness and inlet/outlet velocity values. Figure 4a shows the pressure distribution inside the HCM case LV model from Figure 3a across the 1.0 s time cycle, while Figure 4b shows the distribution of velocity within the fluid. Figure 4: Pressure (**a**) and velocity (**b**) fields obtained using FE simulation of LV parametric model with HCM at three different time points. Figure 3: Left ventricle parametric model with structural mesh.