Research Papers

Effects of Cyclic Motion on Coronary Blood Flow

[+] Author and Article Information
Wei Yin

e-mail: wei.yin@stonybrook.edu
School of Mechanical and
Aerospace Engineering,
Oklahoma State University,
Stillwater, OK 74078

1Corresponding author. Bioengineering Building, Room 109, Department of Biomedical Engineering, Stony Brook University, Stony Brook, NY 11803.

Contributed by the Bioengineering Division of ASME for publication in the JOURNAL OF BIOMECHANICAL ENGINEERING. Manuscript received June 11, 2013; final manuscript received August 26, 2013; accepted manuscript posted September 6, 2013; published online October 4, 2013. Assoc. Editor: Dalin Tang.

J Biomech Eng 135(12), 121002 (Oct 04, 2013) (8 pages) Paper No: BIO-13-1262; doi: 10.1115/1.4025335 History: Received June 11, 2013; Revised August 26, 2013; Accepted September 06, 2013

The goal of this study was to establish a computational fluid dynamics model to investigate the effect of cyclic motion (i.e., bending and stretching) on coronary blood flow. The three-dimensional (3D) geometry of a 50-mm section of the left anterior descending artery (normal or with a 60% stenosis) was constructed based on anatomical studies. To describe the bending motion of the blood vessel wall, arbitrary Lagrangian–Eularian methods were used. To simulate artery bending and blood pressure change induced stretching, the arterial wall was modeled as an anisotropic nonlinear elastic solid using the five-parameter Mooney–Rivlin hyperelastic model. Employing a laminar model, the flow field was solved using the continuity equations and Navier–Stokes equations. Blood was modeled as an incompressible Newtonian fluid. A fluid–structure interaction approach was used to couple the fluid domain and the solid domain iteratively, allowing force and total mesh displacement to be transferred between the two domains. The results demonstrated that even though the bending motion of the coronary artery could significantly affect blood cell trajectory, it had little effect on flow parameters, i.e., blood flow velocity, blood shear stress, and wall shear stress. The shape of the stenosis (asymmetric or symmetric) hardly affected flow parameters either. However, wall normal stresses (axial, circumferential, and radial stress) can be greatly affected by the blood vessel wall motion. The axial wall stress was significantly higher than the circumferential and radial stresses, as well as wall shear stress. Therefore, investigation on effects of wall stress on blood vessel wall cellular functions may help us better understand the mechanism of mechanical stress induced cardiovascular disease.

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Fig. 1

The beginning section of the left anterior descending artery. (a) The 3D geometry constructed in ANSYS DesignModeler; the inserted figure on the top left corner demonstrates the relative location of the region of interest in the left coronary artery (circled). (b) The meshed volume of the solid domain, i.e., the arterial wall. (c) The meshed volume of the fluid domain, i.e., blood.

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Fig. 2

The stenosed (60%) left anterior descending artery. Vessel diameter was reduced asymmetrically or symmetrically. The center of the stenosis throat was located 8 mm downstream of the inlet.

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Fig. 3

In CFX, y location of each node on LAD wall can be determined by the current radius of curvature R(t), the distance between the inlet and outlet (2 l), and x coordinate of the node. R(t) and center of curvature both change with time.

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Fig. 4

The radius of curvature as a function of time

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Fig. 5

Fully developed inlet flow velocity profiles during one cardiac cycle. The maximum flow occurred at t = 0.85 s during diastole; and the minimum flow occurred at t = 0.69 s during systole.

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Fig. 6

Velocity distribution in the normal LAD during one cardiac cycle at t = 0.15, 0.30, 0.75, and 0.85 s. The velocity varied between 0 and 18.8 cm/s, with the maximum velocity occurring at the end of diastole, around the curvature. The black arrows indicate the direction of flow.

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Fig. 7

Velocity vector distribution around an (a) asymmetric and (b) symmetric stenosis in the left anterior descending artery, at t = 0.85 s during diastole. The maximum velocity around the asymmetric stenosis was 31.2 cm/s and that around the symmetric stenosis was 28.5 cm/s. Recirculation zones induced by asymmetric stenosis were generally larger than those induced by symmetric stenosis. The red arrows indicate the flow direction.

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Fig. 8

Wall shear stress contours at t = 0.30 and 0.85 s, in the normal and stenosed LAD artery. Small figures next to each contour depict wall shear stress distribution 8 mm downstream of the bifurcation (marked by red double arrows). For the normal artery, the maximum wall shear stress during one cardiac cycle was 1.2 Pa; with the presence of a 60% stenosis, the maximum wall shear stress increased to 3.31 Pa for the asymmetric case, and 3.02 Pa for the symmetric case. The maximum shear stress occurred at 0.85 s. White arrows indicate flow direction.

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Fig. 9

Longitudinal section view of wall stress distribution at t = 0.85 s in normal and stenosed arteries: (a) axial stress; (b) circumferential stress; and (c) radial stress

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Fig. 10

The trajectories of three randomly chosen particles during one cardiac cycle. The red arrows demonstrate the direction of the arterial motion, and the white arrows demonstrate the flow direction.

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Fig. 11

The trajectories of three randomly chosen particles in the stenosed left coronary arteries at the end of the cardiac cycle (t = 0.9 s)

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Fig. 12

Particle shear stress history along their trajectories, in normal and stenosed left coronary arteries during one cardiac cycle



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