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Research Papers

Comparative Study of Viscoelastic Arterial Wall Models in Nonlinear One-Dimensional Finite Element Simulations of Blood Flow

[+] Author and Article Information
Rashmi Raghu

Department of Mechanical Engineering, James H. Clark Center, E3.1, 318 Campus Drive,  Stanford University, Stanford, CA 94305,raghu21@stanford.edu

Irene E. Vignon-Clementel

 INRIA Paris Rocquencourt, BP 105, 78153 Le Chesnay Cedex, Franceirene.vignon-clementel@inria.fr

C. Alberto Figueroa

Department of Bioengineering, James H. Clark Center, E382, 318 Campus Drive,  Stanford University, Stanford, CA 94305,cafa@stanford.edu

Charles A. Taylor

Department of Bioengineering, Department of Surgery, James H. Clark Center, E350B, 318 Campus Drive,  Stanford University, Stanford, CA 94305,taylorca@stanford.edu

J Biomech Eng 133(8), 081003 (Sep 06, 2011) (11 pages) doi:10.1115/1.4004532 History: Received March 10, 2011; Revised June 05, 2011; Posted July 06, 2011; Published September 06, 2011; Online September 06, 2011

It is well known that blood vessels exhibit viscoelastic properties, which are modeled in the literature with different mathematical forms and experimental bases. The wide range of existing viscoelastic wall models may produce significantly different blood flow, pressure, and vessel deformation solutions in cardiovascular simulations. In this paper, we present a novel comparative study of two different viscoelastic wall models in nonlinear one-dimensional (1D) simulations of blood flow. The viscoelastic models are from papers by Holenstein in 1980 (model V1) and Valdez-Jasso in 2009 (model V2). The static elastic or zero-frequency responses of both models are chosen to be identical. The nonlinear 1D blood flow equations incorporating wall viscoelasticity are solved using a space-time finite element method and the implementation is verified with the Method of Manufactured Solutions. Simulation results using models V1, V2 and the common static elastic model are compared in three application examples: (i) wave propagation study in an idealized vessel with reflection-free outflow boundary condition; (ii) carotid artery model with nonperiodic boundary conditions; and (iii) subject-specific abdominal aorta model under rest and simulated lower limb exercise conditions. In the wave propagation study the damping and wave speed were largest for model V2 and lowest for the elastic model. In the carotid and abdominal aorta studies the most significant differences between wall models were observed in the hysteresis (pressure-area) loops, which were larger for V2 than V1, indicating that V2 is a more dissipative model. The cross-sectional area oscillations over the cardiac cycle were smaller for the viscoelastic models compared to the elastic model. In the abdominal aorta study, differences between constitutive models were more pronounced under exercise conditions than at rest. Inlet pressure pulse for model V1 was larger than the pulse for V2 and the elastic model in the exercise case. In this paper, we have successfully implemented and verified two viscoelastic wall models in a nonlinear 1D finite element blood flow solver and analyzed differences between these models in various idealized and physiological simulations, including exercise. The computational model of blood flow presented here can be utilized in further studies of the cardiovascular system incorporating viscoelastic wall properties.

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Copyright © 2011 by American Society of Mechanical Engineers
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Figures

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Figure 1

Bifurcation model for MMS verification. Variations of reference area, static modulus, and time constant (τ1 ) are shown along each vessel segment for viscoelastic model V2.

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Figure 2

MMS verification results for the finite element implementation of model V2. The MMS flow rate at the inlet is prescribed as a boundary condition for the simulations.

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Figure 3

Wave propagation study showing flow and pressure waves at different locations in the vessel for viscoelastic models V1, V2, and the elastic wall model. The vessel is 80 cm long with reflection-free (characteristic impedance) boundary condition at the outlet.

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Figure 4

Carotid artery model with transient in-flow boundary condition and RCR boundary condition at the outlet. Flow (prescribed), pressure, and cross-sectional area are shown for the inlet of the model.

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Figure 5

Inlet pressure, cross-sectional area and hysteresis for the carotid artery model corresponding to the fifth and ninth cardiac cycles in Fig. 4

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Figure 6

Abdominal aorta geometry of a normal subject from Yeung [61]

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Figure 7

Flow rate and pressure under resting conditions at the inlet, iliac bifurcation, left internal iliac, and right external iliac artery outlets for the subject-specific abdominal aorta model

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Figure 8

Cross-sectional area and hysteresis under resting conditions at the inlet and right external iliac outlet of the subject-specific abdominal aorta model

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Figure 9

Flow rate and pressure under lower limb exercise conditions at the inlet, iliac bifurcation, left internal iliac, and right external iliac artery outlets for the subject-specific abdominal aorta model

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Figure 10

Cross-sectional area and hysteresis under lower limb exercise conditions at the inlet and right external iliac outlet of the subject-specific abdominal aorta model

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