ISSN: 0970-938X (Print) | 0976-1683 (Electronic)

An International Journal of Medical Sciences

- Biomedical Research (2013) Volume 24, Issue 1

^{1}Department of Mechanical Engineering, Iran University of Science and Technology, Narmak, Tehran, Iran

^{2}Department of Civil Engineering, Aalborg University, Sohngårdsholmsvej 57, 9000 Aalborg, Aalborg, Denmark

^{3}Department of Mechanical engineering, Babol University of Technology, Babol, Iran

^{4}Department of Civil Engineering, Sharif University of Technology, Tehran, Iran

- *Corresponding Author:
- N. Sedaghatizadeh

Department of Mechanical Engineering

Iran University of Science and Technology

Narmak, Tehran

Iran

Steady blood flow through a circular artery with rigid walls is studied by COSSERAT Continuum Mechanical Approach. To obtain the additional viscosities coefficients, feed forward multi-layer perceptron (MLP) type of artificial neural networks (ANN) and the results obtained in previous empirical works is used. The governing filed equations are derived and solution to the Hagen-Poiseuilli flow of a COSSERAT fluid in the artery is obtained analytically by Homotopy Perturbation Method (HPM) and numerically using finite difference method. Comparison of analytical results with numerical ones showed excellent agreement. . In addition microrotation and the velocity profile along the radius are obtained by using both numerical and analytical approaches.

Blood flow; artery; Cosserat continuum; viscoelastic response; non-Newtonian

Over 20 million people die every year due to blood related
diseases [1]. Since the blood flow is an important
subject in biomedical and medicine sciences, the behavior
of blood in the artery is one of the most important problems
in biomechanical engineering. Thus many studies
have been performed analytically, numerically and experimentally
to determine the blood behavior and its different
properties [1-5]. The phenomena which are associated
with flow of blood through arteries such as pulsatility
of flow, non-Newtonian behaviour of blood as a fluid and
flexibility of the arteries wall are very complicated; therefore,
theoretical study of blood flow is a very difficult
problem to attack. The nonlinear behavior of blood was
unknown until the second half of the last century [6].
From the rheological point of view, blood is a water based
solution which is the combination of organic and inorganic
substances and a variety of suspended cells mainly
red cells which strongly affects the dynamic of blood as
flood so that blood can be characterized as a non-
Newtonian fluid [7, 8].**Fig. 1** shows the component of a
blood sample [7]. Detailed information can be found in
previous literatures [6-8].

Experimental studies revealed that the blood viscosity decreased with an increase in shear rate and that blood has a small yield stress. Some constitutive models have been proposed for blood as a non-Newtonian fluid by several researchers. Casson proposed a model which is applied successfully for analysis of blood flow by Merrill and others later [8]. Biviscosity model is another model, which express that blood behaves as non-Newtonian fluid in small shear rates and Newtonian fluid in large shear rates. Nakamura and Sawada used this model successfully [9,10]. In the case of blood flow because of considerable mechanical properties which is emanate from its microstructure, the theory of simple materials cannot solve the problem, so a more sensitive continuum theory such as nonlocal, micropolar, multipolar, and gradient theories [5,11,12] with higher kinematic status will be applied more successfully. The idea of a material body endowed with both translational and rotational degrees of freedom stems from the seminal work of the Cosserat brothers (Cosserat and Cosserat) [13, 14]. In this continuum, the effect of couples and forces are considered independently from each other which later named micropolar continuum by Eringen [11]. Fluids of this continuum medium can support the couple stress, the body couple and nonsymmetric stress tensor and possess a rotation field, which is independent of the velocity field. The rotation field is no longer equal to the half of the curl of the velocity vector field. Because of the assumption of infinitesimal rotations, we can treat the rotation field as a vector field. The theory, thus, has two independent kinematic variables; the velocity vector V and the spin or microrotation vector ω .

Because of complex rheological behavior of blood flow and the fact that blood contains a variety of cells with a spinning movement that affects the blood flow velocity, we should applied Cosserat continuum model. This model can consider the effects of rotational movement and has the ability to describe the complex fluid flows such as non-Newtonian and turbulent fluid flow [11, 12, 15, and 16]. Recently

**Sedaghatizadeh** et. al [17] conducted a semi empirical
semi numerical study on blood flow through an artery in
Cosserat continuum, and revealed that the blood flow exhibits
a parabolic trend. They also calculated the non-
Newtonian factors exist in governing equations using
PSO algorithm and find out that these factors are not constant
during a cycle of a heart [17].

**Equation of motion in Cosserat continuum**

In Cosserat continuum both velocity and rotation vector field are considered at any material point. In order to develop the relation between current state of orthonormal directions attached to each material point and its initial state we have the following.

(1)

Where and are the Kronecker delta and permutation
tensor, respectively. The associated Cosserat deformation
tensor ε_{ij} and torsion-curvature tensor k_{ij} are
written as

(2)

(3)

where the comma denotes partial differentiation. It is obvious that in the absence of the rotation vector ω , the classical continuum mechanics is recovered

It is assumed that the transfer of interaction between two
particles of the continuum through a surface element
n_{i} ds occurs by means of both a traction vector t_{i} ds and
a moment vector m_{i} ds . Surface forces and couples are
represented by the generally nonsymmetric force-stress
and couple stress tensor t_{ij} and m_{ij} , respectively. The
balance of linear momentum and angular momentum require
that the following equations must be satisfied

(4)

(5)

Where *ρ* *j,f _{i}* and

(6)

(7)

Where π is the thermodynamic pressure. The linear constitutive
equation for nonsymmetric stress tensor (i.e.
Cauchy's stress tensor), contains an additional viscosity
coefficient k_{v}. The value of k_{v} shows the influence of the
microrotation field on the stress tensor. The linear constitutive
equation for couple stress also contains three additional
viscosity coefficients α_{v} , β_{v} and γ_{v}.

At this stage the above equations should be combined to obtain the governing field equations. The field equations for micropolar fluids in the vectorial form are given by conservation of mass (i.e. continuity equation)

(8)

Balance of liner momentum

(9)

And balance of angular momentum

(10)

**Problem Definition **

**Fig. 2** shows a part of the femoral artery of a dog where
the measurements were made previously at USC School
of Medicine.
The pressure drop through the artery is measured using
two small branch arteries. To simplify the geometry the
vessel (**Fig. 2.a**) assumed and kept to be relatively straight
with mild taper (**Fig. 2. b**). The flow through a vessel is
determined by pressure gradient, which is assumed to be
constant in most of the problems with practical importance
and the vessel is considered as a pipe with radius
of R . The O_{z} axis overlaps the centerline of pipe. In this
case the velocity components and the microrotation velocities
become

(11)

(12)

From continuity when ρ = const [6], we have By neglecting the body forces and body couples, the equations of motion (i.e., Eqs. (9) and (10)) are reduced to the following form in cylindrical coordinate system.

(13)

(14)

The above two coupled partial differential equations should satisfy the following boundary and initial conditions

At (15)

At (16)

where is a dissipative boundary condition [11, 18] and 0 ≤n ≤1. This factor is small for a laminar flow and is increases as flow become turbulent.

**Analysis of the Homotopy Perturbation Method**

Homotopy Perturbation Method (HPM) [19-29] is based on the Homotopy which is an important part of the topology [30-31].

(17)

with the following boundary condition:

(18)

where A is a general differential operator, B a boundary operator, f (r) a known analytical function and Γ is the boundary of the domain Ω. A can be divided into two parts which are L and N, where L is linear and N is nonlinear. Eq. (17) can therefore be rewritten as follows:

(19)

Homotopy perturbation structure is shown as follows:

(20)

where,

(21)

In Eq. (21),p∈[0,1]is an embedding parameter and
μ_{0} is the first approximation that satisfies the boundary
condition. We can assume that the solution of Eq. (21)
can be written as a power series in p, as following:

(22)

and the best approximation for solution is:

(23)

In order to solve Eq. (17) using HPM, we construct the following homotopy; we need an initially condition which is Eq. (18). After constructing the Homotopy Perturbation of Eqs. (13) and (14) and rearranging based on powers of p-terms, we can obtain:

(24)

(25)

(26)

So V_{z} and w_{q} can be obtained as follows with four terms of approximation as

(27)

and

(28)

In the above solution, the viscosity coefficients of COSSERAT medium ( k_{v} , α_{v} , β_{v},γ_{v}) are unknown. Thus the
feed forward multi-layer perceptron (MLP) type of artificial neural networks (ANN) can be employed to calculate these
coefficients on the basis of experimental data. In this study, the flow field and the results of experiments done in Fachhochschule
Frankfurt by Silber [30] on steady blood flow through a dog artery with the diameter of 12 mm, is considered
to determine the viscosity coefficients.

In this study Homotopy Perturbation and Finite Difference
Methods are implemented to study the steady blood
flow through a circular artery. Results are compared in **Figs. 6** and **7** for both normalized microrotation and velocity,
respectively. As it is obvious analytic outcomes are
in good agreement with numerical ones.

Numerical solutions of governing coupled Eqs. (13) and
(14) with boundary conditions (15) and (16) are obtained
by finite difference FORTRAN code. Successive under
relaxation (SUR) method is implemented to solve linear
algebraic equations and the value of 0.1 is taken for under-relaxation parameter. The governing equations are
discretized by applying second-order accurate central difference
schemes. Grid independency was verified on different
node numbers from 51 to 201 and finally the number
of grid points is taken as 101 along the radius of artery.
The convergence criterion (maximum relative error
in the values of the dependent variables between two successive
iterations) was set at10^{-10}.

To obtain the additional viscosities coefficients, feed forward multi-layer perceptron (MLP) type of artificial neural networks (ANN) and the result obtained in [30] is used in this work.

5.1. Modeling using MLP-type neural networks

An example of a MLP type of neural network with one
input node, a single hidden layer with two neuron and one
output neuron is shown in **Fig.3**.
An additional input called bias with constant value of 1 is
added to the previous input node which works as a shift
operator. Each input node is related to each neuron in
hidden layer by a connecting weight. The sum of the
products of the weights and the inputs is calculated by
each neuron in hidden layer and then treated by an activation
function. The obtained result is then multiplied by the
associated weight C_{3} and again the previous procedure
will be repeated in the output neuron. In the present study
hyperbolic tangent and linear functions are used as the
activation functions in the hidden and output layers respectively.

The final output of the current network is calculated as

(29)

where,

(30)

Once the number of layers and the number of neurons in
each layer, have been selected, the network's weights
must be adjusted to minimize the prediction error made
by the network. This is the general role of the *training
algorithms*. In this investigation Back-propagation (BP)
method is applied to train the ANN which is the most
widely used learning process in neural networks today.

**Back-Propagation algorithm**

Back-propagation was firstly proposed by Werbos [33]
which is based on searching an error surface (error as a
function of ANN weights) using the gradient descent algorithm
for points with minimum error.
Consider a network with one hidden layer and one output
neuron as shown in **Fig 4.** When a set of input data (input vector) are propagated
through the network, for the current set of weights there is
an output *Est* . The training of perceptron is a supervised
learning algorithm where weights are adjusted to minimize
the absolute error between the estimated output Est
of network and the desired output *Des* whenever the estimated
output does not match the desired output. If the
network error ( *NE* ) is defined as

*Network Error = Est - Des = NE* (31)

The training algorithm should adjust the weights to minimize
NE^{2} . For this purpose an artificial neuron with its
basic elements is considered as shown in **Fig .5**.

The neuron consists of three basic components; weights, a
summing junction and an activation function. The outputs
of n neurons NO_{1},..., NO_{n} lead in neuron *N* as the inputs.
If neuron *N* is in the hidden layer then this is the
input vector of the network. These outputs are multiplied
by the associated weights W_{1N} ,..., W_{nN}.The summing
junction adds together all these products to provide the
input I_{N} for activation function of neuron *N* . Then
I_{N} passes through the activation function *AF ( )* and
gives the final output of neuron *N* , which is NO_{N} . To
commence the calculations, consider neuron *M* and
weight W_{MN} which connects the two neurons. The equation
for weight update is as follows:

(32)

where LR is the learning rate parameter and is error gradient with reference to the
weight W_{MN}. The chain rule gives

(33)

since the rest of the inputs to neuron N are independent
of the weight W_{MN} we have

(34)

Eqs. (32), (33) and (34) give

(35)

For the case N is an output neuron we have:

(36)

Substituting Eq. (36) into Eq. (35) gives

(37)

For hyperbolic tangent and linear activation functions *AF'(I _{N} )* and the final form of weight update rule can be written
as follows:

(38)

(39)

When N is a hidden layer neuron

(40)

where the subscript *on* represents the output neuron. In Eq. (40) we have

(41)

substituting Eq. (41) into Eq. (40) gives

(42)

In Eq. (42),is now written as a function ofwhich was calculated in Eq. (36). Hence the weight update rule for a hidden layer neuron takes the following form

(43)

The most common parameters for evaluation of a neural network’s performance are minimum total squared errors (or RMS error) and minimum total absolute error (or MAE error). MAE and RMS errors are defined as

(44)

(45)

where *n* is the number of training data. The number of hidden layers and the neurons in each of them should be determined
in a way to minimum the above errors.

The velocity profile along the radius is shown in **Fig. 6**. The diagram shows that the velocity has a parabolic trend and
change from maximum value at center to zero at the wall. Microrotation is illustrated in **Fig. 7** and it shows that, it varies
linearly along the radius.

The maximum values of *V* and ω (V_{max}, ω_{max}) can be obtained analytically by substituting r = 0 and r = R in Eqs.(27, 28) respectively as follows:

(29)

(30)

In this study a flow simulation respect to time and dimensionless radius based on Cosserat model for a complex fluid like blood has been performed. ANN is applied to obtain the non-Newtonian coefficients. Using these values, the velocity and rotation profiles are calculated over radius of the artery are obtained. The obtained results show that, the maximum values of velocity occur in the inlet core; while, the maximum values of rotation happen near and on the boundary during systoles and diastoles.

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