Analysis of Prolapse in Cardiovascular Stents: A Constitutive Equation for Vascular Tissue and Finite-Element Modelling

Abstract

The effectiveness of a cardiovascular stent depends on many factors, such as its ability to sustain the compression applied by the vessel wall, minimal longitudinal contraction when it is expanded, and its ability to flex when navigating tortuous blood vessels. The long-term reaction of the tissue to the stent is also device dependant; in particular some designs provoke in-stent restenosis (i.e., regrowth of the occlusion around the stent). The mechanism of restenosis is thought to involve injury or damage to the vessel wall due to the high stresses generated around the stent when it expands. Because of this, the deflection of the tissue between the struts of the stent (called prolapse or "draping") has been used as a measure of the potential of a stent to cause restenosis. In this paper, uniaxial and biaxial experiments on human femoral artery and porcine aortic vascular tissue are used to develop a hyperelastic constitive model of vascular tissue suitable for implementation in finite-element analysis. To analyze prolapse, four stent designs (BeStent 2, Medtronic AVE; NIROYAL, Boston Scientific; VELOCITY, Cordis; TETRA, Guidant) were expanded in vitro to determine their repeating-unit dimensions. This geometric data was used to generate a finite element model of the vascular tissue supported within a repeating-unit of the stent. Under a pressure of 450 mm Hg (representing the radial compression of the vessel wall), maximum radial deflection of 0.253 mm, 0.279 mm, 0.348 mm and 0.48 mm were calculated for each of the four stents. Stresses in the vascular wall were highest for the VELOCITY stent. The method is proposed as a way to compare stents relative to their potential for restenosis and as a basis for a biomechanical design of a stent repeating-unit that would minimize restenosis.

Effects of material, coating, design and plaque composition on stent deployment inside a stenotic artery—Finite element simulation

“…The difference between Mooney-Rivlin and Ogden models is just the type of formulation for strain energy potential (Prendergast et al, 2003;Gastaldi et al, 2010; Zahedmanesh and Lally, 2009; Karimi et al, 2014). We have compared the simulations using the Ogden model and the Mooney-Rivlin model, and the two models gave very similar results of stent expansion behaviour.…”

Differential impact of local stiffening and narrowing on hemodynamics in repaired aortic coarctation: an FSI study

“…At the boundaries, only radial displacement is allowed. The material behavior of the aortic tissue is described using a polynomial hyperelastic model (hyperelastic constants: C10 = 18.9 kPa, C01 = 2.75 kPa, C20 = 400 kPa, C11 = 847.2 kPa [34]) with the value of C20 obtained in an iterative way, such that the deformations of the descending aorta in the FSI simulation corresponded to the deformations measured with MRI (9%).…”

Fluid–structure interaction simulation of pulse propagation in arteries: Numerical pitfalls and hemodynamic impact of a local stiffening

“…In the second case, a physiological pressure pulse (period T of 1 s, amplitude of approximately 60 mmHg) is used as an inlet boundary condition. As the deformations become more pronounced in this case, a hyperelastic material model is used (a reduced form of the generalized Mooney-Rivlin) with experimentally obtained coefficients for the human aorta (C 10 = 18.9, C 01 = 2.75, C 20 = 590.4, C 11 = 857.18 and C 30 = 0 kPa) [24]. The behavior of the local stiffening is modeled using a linear elastic model (Young's modulus 1500 kPa).…”