## Materials and Methods
### Numerical Discretization, Blood Rheology and Boundary Conditions  
All geometries were imported into the commercial package IcomCFD, Version 16 (ANSYS Inc., Canonsburg, PA, USA) for the generation of body-conforming volumetric grids. The computational meshes were composed of tetrahedral elements.  
IcomCFD creates computational grids utilizing user-specified parameters. In particular, the physical domain of each model was firstly reduced into a number of regions. Then, different mesh densities were used for each region: a finer mesh with smaller tetrahedral elements was used in the regions around the stent struts, due to their small dimensions. For ensuring a correct resolution of the velocity and hence a proper computation of the wall shear stress and its variables, the numerical mesh was refined near all walls. In addition, a grid independence study was carried out prior to performing the present simulations. Details can be found in a previous study [28] that was used as a basis for generating the numerical grids in the present work. The total number of mesh elements ranged from approximately 0.3 to \(20\times 10^{6}\) elements depending on the presence or not of the intravascular device and on the size of the stent. The final grids used in this study are summarized in the Table 3.  
Unsteady CFD simulations for the 12 geometrical configurations considered were carried out using the commercial software Ansys CFX, Version 16.0 (Ansys Inc., Canonsburg, PA, USA). Blood flow was modeled as an incompressible and non-Newtonian fluid by means of the Carreau-Yasuda model. This model, among others, is largely used in the literature for the simulation of hemodynamic disorders [28; 33]. The constitutive equation that represents this model is:  
\[\mu_{eff}=\mu_{\infty}+(\mu_{0}-\mu_{\infty})\cdot\left[1+\left(\lambda\dot{ \gamma}\right)^{a}\right]^{\frac{n-d}{a}} \tag{1}\]  
where \(\mu_{0}=0.056\) is the viscosity at zero shear rate expressed in \(\mathrm{Pa}\cdot\mathrm{s}\), \(\mu_{\infty}=0.00345\) is the viscosity for an infinite shear rate expressed in \(\mathrm{Pa}\cdot\mathrm{s}\), \(\lambda=3.313\) is the relaxation time expressed in \(s\), \(n=2\) is the power exponent and \(a=0.64\) the Yasuda exponent. The blood flow was set to 1060 \(\mathrm{kg/m^{3}}\). The vessel wall was assumed to be rigid with a no-slip boundary condition. As in the previous study [28], a time-dependent flow waveform found in the literature [39] was imposed at the inlet and outlet of the model using uniform velocity profiles. For that reason, the inflow and outflow regions were extended sufficiently to provide fully developed flow in the region near the device and to reduce the influence of the boundary conditions on the computational domain. Cardiac cycles of 1 s were discretized in time steps of 0.001 s [28]. Finally, three cardiac cycles were computed for each simulation in order to damp the initial transients. The flow regime has been considered as laminar as the inflow Reynolds number based on the peak flow at systole was 604.  
\begin{table}
\begin{tabular}{c c c c} \hline \hline
**Vessel** & **Diameter [mm]** & **Aspect Ratio [mm]** & **Sient** \\ \hline V-bent, C-bent, S-bent & 4.75 & 1.68 & No stent, STENT 1, 2, 3 \\ \hline \hline \end{tabular}
\end{table}
Table 2: Summary of the geometrical cases analyzed.  
\begin{table}
\begin{tabular}{c c c c c} \hline \hline  & **No Stent** & **STENT 1** & **STENT 2** & **STENT 3** \\ \hline
**3D elements** & 329,286 & 10,949,532 & 14,403,306 & 18,358,866 \\ \hline
**2D elements** & 94,830 & 1,143,257 & 1,040,456 & 1,440,800 \\ \hline \hline \end{tabular}
\end{table}
Table 3: Summary of the computational meshes used in this study.