Cath Lab Digest

Introduction
Early stent-graft designs were ill-prepared to face the hemodynamic forces generated by aortic blood flow and pressure. Long-term results were marred by high rates of migration,¹ component separation,² and structural failure.³ Current designs have proven to be more stable.² Though more rigorous pre-clinical testing has helped to assess the likelihood of each known failure mode before devices are implanted in patients, there is still no consensus as to the test conditions, because the hemodynamic environment of the endovascular stent-graft remains poorly understood.

Previously published data on fluid dynamics in aortic stent-grafts have used steady state flow conditions and hypothetical geometries to model the hemodyanamic load.^4,5 These studies helped to identify key determinants of displacement force, such as inlet diameter, blood pressure and limb angulation. Despite the simplicity of the models, clinical studies have shown that the same factors correlate with rates of stent-graft migration.⁵

Nevertheless, steady state modeling of the summated forces within simple constructs cannot provide an accurate representation of clinical conditions. The repetitive fluctuation of flow, pressure and force during the cardiac cycle is far more destabilizing. Cyclical stress and strain degrade not only stent-graft attachment, but also stent-graft structure. Persistent flexing of the metal stents generates stress that can cause stent fracture, and the movement of the metal stents along the graft fabric can cause holes in the stent fabric.³ Our group has used cine-flouroscopy to observe both pulsatile displacement of the graft and pulsatile diameter change of stent-grafts.⁶ We also used computational methods to model the effects of pulsatile pressure and flow at various points throughout the cardiac cycle, and various points on the surfaces of realistic geometries (derived from CT data in real stent-grafts) under realistic conditions of external (sac) pressure.⁶ In addition, we used cine-fluoroscopy to make direct observations of the cyclical strain (pulsatile micro-movement) exhibited by stent-grafts at various stages after clinical implantation. Both types of data are helpful in designing better stent-grafts and pre-clinical tests.

Computational Fluid Dynamics (CFD)
Methods. The fluid flow equations (Navier-Stokes) can be solved by analytic mathematical means only with simplified hypothetical models. It is only possible to determine internal hemodynamic effects, such as wall shear stress and local pressure, using numerical methods. Computational Fluid Dynamics (CFD) is a method for generating the numerical solutions to these equations at discrete points over a three-dimensional surface. The three-dimensional computational grid can be generated using either hypothetical stent-graft models or 3-dimensional imaging of real stent-grafts in clinical use. In either case, the validity of the CFD simulation depends on the accuracy of the chosen boundary conditions, including the viscosity and shear properties, the inlet and outlet flow velocities, the inlet and outlet blood pressures, and the pressure on the outside of the graft.

Previously published simulations of the hemodynamic forces through the stent-graft have used both analytical and numerical techniques, but none of these simulations modeled internal fluid motion. Most only calculated the hemodynamic force due to steady state flow, usually a mean peak inlet flow velocity. Furthermore, all these simulations used simplified models of the stent-graft geometry, incorporating only inlet and outlet diameters.^4,5,7,8 The models were generally composed of bifurcated tubes with geometrical stent-graft characteristics with proximal and branch components lying in a single plane. Limb angulation was limited to curvature in the same plane as the rest of the graft.^4,7 The sole exception was an analysis of a single patient-specific geometry.⁸

We have used both hypothetical (Figure 1) and patient-specific (Figure 2) geometries, for the 3-dimensional computational grids in our CFD simulations.⁶ In the hypothetical geometries, the different elements of graft design were varied to isolate factors such as diameter and angulation, and assess their effects on the displacing forces within various parts of the stent-graft. Our patient-specific models were derived from the arterial phase of high-resolution contrast-enhanced CT scans. Each 3-dimensional arrangement of contrast-enhanced voxels within a CT representation of luminal anatomy was converted to a surface mesh of interconnected triangles using an automated algorithm. A series of CT scans were selected to cover a range of possible stent-graft orientations. In addition, the different regions of the graft, the proximal trunk, the bifurcation and the iliac branches, were modeled separately to identify their contribution to the forces on the stent-graft.

In any CFD analysis of pressure-generated forces, the pressure on the outside of the stent-graft (sac pressure) is as important as the pressure on the inside of the stent-graft (systemic arterial blood pressure.) The transmural pressure gradient is what matters, not the absolute pressure. Most previous CFD models have assumed this outside pressure to be zero. In the absence of patient-specific data on sac pressure, published data were used on the mean pressure in each of four clinical circumstances, relating to endoleak status and aneurysm diameter change.⁹