ARUP; ISBN: 978-0-9562121-5-3 - CMBBE 2012 - Cardiff University

The non-linear behaviour of the artery was described using nine parameter Mooney-
Rivlin hyperelastic constitutive equation, which were taken from literature 5 . The plaque
was modelled with an elasto-plastic material: the yielding point was identified from
experimental data of tensile tests on plaque samples 6 .
Both the vessel model and the stent were meshed with eight-node Solid185 elements.
Appropriate boundary conditions were used to simulate stent crimping and expansion
into the 3D model of the stenotic vessel, after a procedure of percutaneous transluminal
angioplasty (PTA). After that, physiological loading condition of blood pressure and
axial compression of the artery were simulated.
The simulation of the PTA requires a definition of a balloon model: this was considered
as a cylindrical rigid body, placed into the vessel (Fig. 4a) and then expanded in radial
direction until a diameter ensuring the recovery of the original vessel lumen (Fig. 4b).
After this step, the balloon is deflated, allowing the elastic recoil of the plaque and the
vessel (Fig. 4c). In the same time of the plaque pre-dilatation, the crimping of the stent
was performed, using a cylindrical rigid body in displacement control, resembling the
sheath (Fig. 4b). After the pre-dilatation of the plaque, the stent was released into the
vessel (Fig. 4d). In order to simulate the physiological condition of the vessel, constant
diastolic pressure acting on the inner surface of the artery was considered during all
stages of the procedure.
Figure 4. Stenting procedure: initial configuration (a), end of stent crimping and expansion of the balloon
for angioplasty (b), after the release of the angioplasty balloon (c), after stent self-expansion (d).
Flexible–rigid contact pairs between the rigid body and stent outer surface and between
the balloon and the plaque inner surface, and flexible–flexible contact pairs between
stent outer surface and plaque inner surface were created. A friction coefficient (0.2)
was set to all contact pairs for simulation stability.
After the stent placement into the vessel, in vivo cycling loading conditions acting on
SFA were applied: blood pressure and axial compression. In the first case, additional
steps were added to the previous model in which systolic and diastolic pressure were
applied on the inner surface of vessel and plaque. During the cardiac cycle the devices
reaches a minimum value of diameter in diastole and a maximum value in systole: it is
subjected to repeated cycles of deformation (1.1x10 5 cycles systole-diastole/day) that
could determine its fatigue failure.
In the second case, after the stent was deployed into the vessel, it was subjected to a
cyclic loading of 5% axial compression that reproduces the deformations caused by hip
and knee joint movement which could determine the device fatigue failure. The stent
was subjected to cycles of compression through displacement applied on plaque ends:
the plaque-vessel was axially stretched by 5% before the stent deployment, then after
the device release it was carried back to its original length and finally stretched again.
As a result of the friction (0.2) defined for the contact between the outer stent surface