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Schiavone et al. Modelling of metallic and polymeric stents

levels of non-physiologic stresses provoked a more over the scaffolded vessel segments. [34]
aggressive pathobiological response of the vessel
[35]
walls, leading to a higher degree of neointimal formation. Residual stresses in stents were studied by Möller et al.
In-vitro studies have also proved that mechanical using X-ray diffraction method. Their work confirmed
stresses regulated the proliferation and migration of that even for as-produced stents, a significant amount
vascular cells, and the synthesis and reorganization of microstresses can be developed in the stent during
[32]
of extracellular matrix. Stent-induced vessel stresses crimping (this is also the case found in our work).
[33]
are closely linked with the level of artery injury, also Subsequent stent expansion caused an increase of
promoting the development of restenosis. From this stress due to tension. According to their study, the level
study, it is clear that the stresses in the artery appear to of stresses introduced by crimping and expansion can
be largely affected by the stent materials and designs. considerably affect the fatigue life of the stent. Based
PLLA has lower modulus, yield strength and strain on our results, crimping and expansion processes were
hardening, which soothed the stent-artery interaction found to induce comparable levels of stresses in the
and led to stress reduction in vessel layers as shown stent struts. Also, residual stresses developed during
in Figure 10. This is clinically beneficial. Bioresorbable crimping affected the expansion behaviour, though
polymeric stents are also more compliant than metallic only slightly, of stent in the deployment step. It is also
stents, diminishing associated vascular responses thought that residual stresses in the stent contributed to
the flexibility of the device, thus imposing less stresses
on the plaque during further deformation as confirmed
by our simulation results, although only marginally.
The targeted vessel diameter (i.e. 3 mm) could not
be achieved by stent expansion only. This is the
case for both polymer and metallic stents. Firstly,
it was due to the saturated expansion of the artery
layers, developed at a later stage of vessel stretching.
This happened when the stiffness of vessel layers,
especially the intima layer, increased steeply upon
large stretch [Figure 4]. Secondly, vessel layers were
assumed to deform purely elastically which imposed
a large recovery force on the expanded stent after
balloon deflation. Generally, the expansion of polymer
stent was slower than that of metallic stent. Higher
recoiling was also observed for polymeric stent due to
the weaker mechanical properties. This indicates that
there is a challenge to use polymer stents to achieve
desired lumen diameter, especially for patients with
stiffer artery and heavily calcified plaques. However, it
is recognised that polymer stent induced significantly
lower stresses in the artery than metallic stents, which
could reduce the occurrence of arterial injuries and in-
stent restenosis.
Figure 13: (A) Diameter change against pressure; and (B) recoiling
and dogboning effects obtained from simulations with and without
considering crimping-caused residual stresses on Xience stent In clinical practice, most stents are post-dilated with

Figure 14: (A) von Mises stress on the Xience stent; and (B) maximum principal stress on the artery-plaque system for simulation with (left)
and without (right) considering residual stresses
Vessel Plus ¦ Volume 1 ¦ March 31, 2017 19