Biomedical Engineering Reference
In-Depth Information
Fig. 8.2  Computational model showing the ( a ) fluid and ( b ) solid domain meshes. Grid indepen-
dence analysis was performed at 3 different mesh refinement levels: coarse (  20,000 elements ),
medium (  500,000 elements ) and fine (  1,000,000 elements ). Only 2 % of the dissimilitude between
the fine and medium mesh is observed
in the solid solver to calculate the artery deformation. The coupling performance
is enhanced by the “artificial compressibility” method (Raback et al. 2001). The
blood flow is modelled as laminar where the peak Reynolds number, which oc-
curs at the stenosis, is approximately 1000. The time step size is set to 0.015s and
the results are obtained at the 4th cardiac cycle to avoid any start-up effects from
initial conditions.
Figure 8.2 shows the computational mesh of the stenosed model with approxi-
mately 1,000,000 cells made up of tetrahedral cells in the inner fluid domain and
prism cells at the near wall region along the wall boundaries
8.2.2
Simulation Details
In the solid domain, each end of the artery (CCA, ICA, ECA) are modelled as
fixed supports while a symmetrical condition is assumed at the plane of the bi-
furcation. In the fluid domain, a pressure inlet boundary condition is imposed
using a time-varying pressure waveform as shown in Fig. 8.3a referred from
Tada (2005), while outlet boundaries are governed by transient mass flow rate
profiles as given in Fig. 8.3b . For simplicity, the arterial wall is assumed, the
material elasticity properties are treated as isotropic based on the Hooke's Law
(Salzar et al. 1995; Thubrikar and Robicsek 1995), and the blood flow is assumed
to be Newtonian with a constant dynamic viscosity of 0.0035 Pa-s. An alterna-
tive approach to represent the material hyper-plastic elasticity property and the
non-Newtonian flow behaviour is the Mooney-Rivlin elastic model (Torii et al.
2009b) and Casson fluid model, which require more computational sources and
longer simulation time.
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