Research Article  Open Access
Pulsatile Flow of TwoFluid Nonlinear Models for Blood Flow through Catheterized Arteries: A Comparative Study
Abstract
The pulsatile flow of blood through catheterized arteries is analyzed by treating the blood as a twofluid model with the suspension of all the erythrocytes in the core region as a nonNewtonian fluid and the plasma in the peripheral layer as a Newtonian fluid. The nonNewtonian fluid in the core region of the artery is represented by (i) Casson fluid and (ii) HerschelBulkley fluid. The expressions for the flow quantities obtained by Sankar (2008) for the twofluid Casson model and Sankar and Lee (2008) for the twofluid HerschelBulkley model are used to get the data for comparison. It is noted that the plugflow velocity, velocity distribution, and flow rate of the twofluid HB model are considerably higher than those of the twofluid Casson model for a given set of values of the parameters. Further, it is found that the wall shear stress and longitudinal impedance are significantly lower for the twofluid HB model than those of the twofluid Casson model.
1. Introduction
Catheters are of extensive use in modern medicine. In routine clinical studies, the measurement of arterial blood pressure/pressure gradient and flow velocity/flow rate are achieved by the use of an appropriate catheter tool device in the desired part of the arterial network [1]. Catheters are also used in diagnostic techniques (such as Xray angiography, intravascular ultrasound, and coronary balloon angioplasty) as well as in the treatment (balloon angioplasty) of various arterial diseases [2]. Catheters are even used to clear the short occlusions from the walls of the stenosed artery [3]. By reducing the obstruction through balloon angioplasty, the mean translesional pressure drop is reduced [4] and the coronary blood flow as well as the coronary flow reserve is increased [5]. The insertion of a catheter in an artery will alter the flow field, modify the pressure distribution, and hence increases the flow resistance [6]. Thus, the pressure or pressure gradient recorded by a transducer attached to the catheter will differ from that of an uncatheterized artery, and hence it is essential to know the catheterinduced error [7]. Even, a very small angioplasty guidewire leads to a sizable increase in flow resistance [8]. For smaller infusion catheter, the increase in flow resistance is less, although still appreciable. Hence, it is meaningful to study the increase in flow resistance due to catheterization.
Several theoretical and experimental investigations are performed to study the dynamics of blood flow through catheterized arteries [9–14]. MacDonald [15] discussed the blood flow characteristics in catheterized arteries using conformal transformation and finite difference method. Sarkar and Jayaraman [1] obtained the correction to flow ratepressure drop relationship in coronary angioplasty with steady steaming effect. Dash et al. [6] analyzed the effect of catheterization on various flow characteristics in a curved artery using perturbation method. Dash and Daripa [16] have studied the blood flow characteristics in an eccentric catheterized artery using a fast algorithm. In all the above investigations, Newtonian fluid represents blood. But it is well known that blood, being suspension of cells, behaves like a nonNewtonian fluid at low shear rate and during its flow through narrow blood vessels of diameter 0.02–0.1 mm [17–19]. Dash et al. [3] studied the steady and pulsatile flow of Casson fluid for blood flow through catheterized arteries and estimated the increase in frictional resistance using the perturbation analysis. Sankar and Hemalatha [7, 20] have studied the steady and pulsatile flow HerschelBulkley fluid for blood flow through catheterized arteries using perturbation method and estimated the increase in longitudinal impedance to flow.
Srivastava and Saxena [19] and Misra and Pandey [21] have mentioned that for blood flowing through narrow blood vessels there is a peripheral layer of plasma and a core region of suspension of all the erythrocytes. Hence, for a more realistic description of blood flow, it is appropriate to treat the blood as a twofluid model consisting of a core region containing all the erythrocytes as a nonNewtonian fluid and the plasma in the peripheral layer as a Newtonian fluid [19–21]. Sankar and Lee [22] and Sankar [23] have analyzed the pulsatile flow of twophase fluid models for blood flow through catheterized narrow arteries at low shear rates, by treating the fluid in the core region as HerschelBulkley (HB) model and (ii) Casson model, respectively. In both of the twofluid models, Newtonian fluid represents the fluid in the peripheral layer.
It is noticed that blood obeys Casson's equation only for moderate shear rate and the HerschelBulkley equation represents fairly closely what is occurring in blood [7, 18]. Chaturani and Ponnalagar Samy [24] have mentioned that for tube diameter 0.095 mm blood behaves like HB fluid rather than powerlaw and Bingham fluids. Iida [25] reports “That velocity profile in the arterioles having diameter less than 0.1 mm are generally explained fairly by the two models. However, velocity profiles in the arterioles whose diameters are less than 0.065 mm does not conform to the Casson model but can still be explained by HB fluid model. Moreover, HB fluid model can be reduced to powerlaw fluid model when the yield stress is zero and Bingham fluid model when its powerlaw index n takes the value 1, so that the twofluid powerlaw and Bingham models can be studied from the twofluid HB model itself as its particular cases. Thus, the twofluid HB model has more suitability than the twofluid Casson model in the studies of blood flow through narrow arteries. Hence, in this paper, the expressions for the flow quantities obtained by Sankar and Lee [22] for the twofluid HB model and the expressions for the flow quantities obtained by Sankar [23] for the twofluid Casson model are used to compare these fluid models and bring out the advantage of using the twofluid HB model over twofluid Casson model for blood flow in catheterized arteries. In this study, the governing equations and the boundary conditions of both of the twofluid models and the expressions obtained for the various flow quantities of these models by Sankar and Lee [22] and Sankar [23] are mentioned in brief and are used to perform a comparative study. The layout of the paper is as follows.
The formulation and method of solution of (i) twophase Casson fluid model and (ii) twophase HerschelBulkley (HB) fluid model are briefly given in Section 2. The variations of the flow quantities of these twofluid models on the yield stress, catheter radius ratio and pulsatility of the flow are analyzed in Section 3. The increase in the longitudinal impedance to flow due to catheterization for different types of catheters which are used in clinics, is also computed for both of the twophase fluid models and is discussed in Section 3. The results are summarized in the concluding Section 4.
2. Mathematical Formulation
Consider an axially symmetric, laminar, pulsatile, and fully developed unidirectional flow of blood (assumed to be incompressible) in the axial direction in an artery in which a catheter is introduced coaxially, where the artery is modeled as a rigidwalled circular tube of radius . The catheter radius is taken to be . Blood is represented by a twofluid model with the suspension of all of the erythrocytes in the core region as a nonNewtonian fluid and the plasma in the peripheral region as a Newtonian fluid. The nonNewtonian fluid in the core region is represented by (i) Casson fluid model and (ii) HerschelBulkley fluid model. We have used the cylindrical polar coordinates , where and denote the radial and axial coordinates and is the azimuthal angle. The flow geometries of the twofluid model for blood flow through catheterized artery are shown in Figure 1.
2.1. TwoFluid Casson Model
2.1.1. Governing Equations and Boundary Conditions
For unidirectional flow of blood (incompressible fluid) in the axial direction, the equation of continuity reduces to
Since the flow is axisymmetric and one dimensional in the axial direction, the radial component of the velocity and the circumferential component of the velocity do not exit (become zero) and only the axial component of the velocity exists, and hence, the equation of continuity (2.1a) reduces to
The axial component of the momentum equation for the onedimensional flow of blood in the axial direction is given as
Using (2.1b) and (the radial component of the velocity and circumferential component of the velocity become zero for onedimensional flow of blood in the axial direction) in the momentum equation (2.2a), we get
The momentum equation (2.2b) is rewritten for the fluid flow in the core region and peripheral region of a twofluid model, respectively, as
where denotes the pressure; denote the density of the Casson fluid and Newtonian fluid, respectively; denote the shear stress of the Casson fluid and Newtonian fluid, respectively; denote the fluid's velocity in the core region and peripheral region, respectively; denotes the time; is the radius of the core region of the artery. The simplified form of the constitutive equations of the fluid in motion in the core region (Casson fluid) and peripheral layer (Newtonian fluid) are given by
where are the viscosities of the Casson fluid and Newtonian fluid, respectively; is the yield stress; 1 and 2 are the yield planes bounding the plugflow region. Equations (2.3a), (2.3b), and (2.6) are equipped with the following boundary conditions:
2.1.2. Nondimensionalization
Let be the absolute magnitude of the typical pressure gradient. Let us introduce the following nondimensional variables:
where are the Womersley numbers of the Casson fluid and Newtonian fluid, respectively, and is the nondimensional yield stress. The pressure gradient can be written as
where is the nondimensional pressure gradient along the axis, which is taken as a periodic function of time for pulsatile flow. Using (2.8) and (2.9), the momentum equations (2.3a)) and (2.3b), and the constitutive equations (2.4)–(2.6) are simplified, respectively, to
The boundary conditions (in the nondimensional form) are
2.1.3. Perturbation Method
Since it is not possible to find an exact solution of the nonlinear coupled implicit system of partial differential equations (2.10)–(2.15), the perturbation method is used to solve the system of nonlinear partial differential equations. When we nondimensionalize (2.3a)), the Womersley number occurs naturally, and hence; it is appropriate to expand the unknown in powers of as below:
Similarly, one can expand the other unknowns in the perturbation series as in (2.17). Hereafter, for our convenience, we have used “” instead of “”. Using the perturbation series expansions in the momentum equations (2.10) and (2.11) and then equating the constant terms and the firstorder terms, we get
Using the perturbation series expansions ofin the constitutive equations (2.12)–(2.15) and then equating the constant terms and the firstorder terms, we obtain
When ,
When ,
When ,
When ,
The boundary conditions (2.16) become
Equations (2.18)–(2.23) are solved explicitly with the help of the boundary conditions (2.24). For detailed derivation of the solution to variables and from (2.18)–(2.24), one can refer Sankar [23]. The detailed derivation for the flow rate, wall shear stress and longitudinal impedance are given by Sankar [23] and one can go through this reference to know about the details of obtaining these flow quantities.
2.2. TwoFluid HerschelBulkley Model
2.2.1. Governing Equations and Boundary Conditions
Following the derivation of (2.3a)) and (2.3b), the basic momentum equations in this case simplify to
where denotes the pressure; denote the density of the HerschelBulkley (HB) fluid and Newtonian fluid, respectively; denote the shear stress of the HB fluid and Newtonian fluid, respectively; denote the fluid's velocity in the core region and peripheral region, respectively; denotes the time; is the radius of the core region of the artery. The simplified form of the constitutive equations of the fluid in motion in the core region (HB fluid) and peripheral layer (Newtonian fluid) are given by
where are the viscosities of the HB fluid and Newtonian fluid; is the yield stress; and are the yield planes bounding the plugflow region. Equations (2.25), (2.26) and (2.27)–(2.30) can be solved with the help of the following boundary conditions:
2.2.2. Nondimensionalization
Let be the absolute magnitude of the typical pressure gradient. Let us introduce the following nondimensional variables:
where is the typical viscosity coefficient having the dimension as that of the Newtonian fluid's viscosity, are the Womersley numbers of the HB fluid and Newtonian fluid, respectively and is the nondimensional yield stress. The pressure gradient can be written as where is the nondimensional pressure gradient along the z axis. Using (2.32) and (2.33), the momentum equations (2.25) and (2.26) and the constitutive equations (2.27)–(2.30) are simplified, respectively, to
The boundary conditions (in the nondimensional form) are
2.2.3. Perturbation Method
As it is not possible to find an analytic solution of the nonlinear coupled implicit system of partial differential equations (2.34)–(2.39), a perturbation method is used to solve the system of partial differential equations. When we nondimensionalize (2.25) and (2.26), the Womersley numbers occur naturally and hence it is appropriate to expand the unknowns in powers of and . Let us expand the velocity in the perturbation series as below: Similarly, one can expand in the perturbation series as in (2.41). Hereafter, for convenience, we have used “” instead of “”. Using the perturbation series expansion of in the momentum equations (2.34) and (2.35) and the constitutive equations (2.36)–(2.39) and then equating the constant terms and the firstorder terms, one can obtain
When ,
When
When
When
The boundary conditions (2.40) become
Equations (2.42)–(2.46) can be solved explicitly with the help of the boundary conditions (2.47). Detailed derivation of the solution to the unknowns variables and from (2.42)–(2.47) are given by Sankar and Lee [22] and one can go through this reference for the detailed solution. The detailed derivation for the flow rate, wall shear stress, and longitudinal impedance are given by Sankar and Lee [22] and one can go through this reference to know about the details of obtaining these flow quantities.
3. Results and Discussion
The objective of the present study is to compare the twofluid HB model and twofluid Casson model. The typical value of the powerlaw index n of the HB fluid for blood flow models is generally taken as 0.95 [26]. Though the yield stress of blood at a haematocrit of 40 is dyne/cm^{2} [27], the range to 0.1 is more suitable when a catheter is inserted into the blood vessels [3]. Just to pronounce the variations in the flow quantities, we have taken the range of yield stress as 0 to 0.25 in this study. The range 0–0.6 is used for the catheter radius ratio k [3].
Since the flow is pulsatile and any periodic function can be represented by a Fourier series, it is appropriate to choose the pressure gradient as , where is the amplitude parameter and is taken as less than 1. In the present study, we use the range 0.2–0.5 for the amplitude parameter to discuss its influence [22]. The ratio or between the Womersley numbers of the Newtonian fluid and HB fluid or Casson fluid is called Womersley number ratio. It is noted that in this ratio, the numerator corresponds to the Womersley number of the Newtonian fluid and the denominator corresponds to the Womersley number of a nonNewtonian fluid. It is well known that the Womersley number of the Newtonian fluid would be higher than that of the nonNewtonian fluid. Thus, we have chosen the Womersley number ratio as less than 1 and particularly as 0.5. Although the Womersley number of the HB fluid also ranges from 0 to 1 [22], the value 0.5 is used in this study. Given the values of and , the value of can be obtained from . Similarly, the Womersley number of the Casson fluid also ranges from 0 to 1 [23]; the value 0.5 is used in the present study.
3.1. Yield Plane Locations
The location of a point where the shear stress is equal to the yield stress is called a yield point and the locus of such points is called yield surface or yield plane. In the case of a tube flow, there is only one yield plane, whereas, for annular flow, there are two yield planes and and these two yield planes form the boundary of the plugflow region. The width of the plug core region is denoted by and is defined as , where is the yield stress in the nondimensional form which ranges from 0–0.25, is the nondimensional pressure gradient which is taken as for pulsatile flow of blood, is the amplitude of the flow whose range is taken as 0.2–0.5, and is the time parameter, Knowing the values of , and , one can compute the value of . For pulsatile flow, the yield plane locations change not only during the course of motion, but also, with respect to the other parameters.
Figure 2 illustrates the width of the plugflow region of the different fluid models for blood flow through catheterized arteries. The variation of the yield plane locations in a time cycle for different fluid models and is depicted in Figure 2. It is clear that for all the fluid models, the width of the plugflow region decreases as the time parameter t increases from 0° to 90° and then it increases as the time variable t increases from 90° to 270° and then again it decreases as the time t increases further from 270° to 360°. The width of the plugflow region is minimum at 90° and maximum at 270°. It is also observed that at any instant of time, the width of the plugflow region of the twofluid models are marginally lower than those of the singlefluid models. There is not much difference between the widths of the twofluid HB model and Casson models and the similar behavior is noticed for the singlefluid HB and Casson models. It is of importance to note that the plot of the singlefluid HB model is in good agreement with Figure of Sankar and Hemalatha [7] and the plot of the singlefluid Casson model is in good agreement with Figure of Dash et al. [3].
3.2. PlugFlow Velocity
Figure 3 depicts the simultaneous effects of the nonNewtonian nature of the fluid and the catheter on plugflow velocity of different twofluid models for blood flow through catheterized arteries. The variation of the plugflow velocity with catheter radius ratio k for different twofluid models with , and is shown in Figure 3. The plugflow velocity for different twofluid models decreases nonlinearly with the increase of the catheter radius ratio k. The plugflow velocity decreases rapidly as the catheter radius ratio k increases from 0.1 to 0.3 and then it decreases gradually as the catheter radius ratio increases further from 0.3 to 0.6. It is found that for a given value of the catheter radius ratio k, the plugflow velocity is maximum for the twofluid powerlaw model and minimum for the twofluid Casson model. It is also clear that the plugflow velocity for the twofluid HB model is higher than that of the twofluid Casson model. It is also observed that the plugflow velocity decreases slightly with the increase of the powerlaw index n.
3.3. Velocity Distribution
The velocity distribution for different twofluid models with , and is shown in Figure 4. One can easily observe the flattened velocity profiles for the twofluid models, which have fluids with yield stress in the core region, and the usual parabolic velocity profile for the twofluid power model, which has no yield stress. The twofluid powerlaw model has the velocity with highest magnitude and the twofluid Casson model has the velocity with lowest magnitude. The velocity for the twofluid HB model is considerably lower than that of the twofluid powerlaw model and significantly higher than that of the twofluid Casson model.
3.4. Flow Rate
Figure 5 depicts the transient changes in the flow rate of different fluid models for blood flow through catheterized arteries. The variation of the flow rate in a time cycle for different fluid models with , and is plotted in Figure 5. It is found that for all the fluid models, the flow rate increases as the parameter t increases from 0° to 90° and then it decreases as the time t increases from 90° to 270° and then it increases as the time t increases from 270° to 360°.The flow rate is maximum at 90° and minimum at 270°. At any instant of time t, the flow rate of the twofluid HB model is significantly higher than that of the twofluid Casson model. A similar behavior is observed for the singlefluid HB and Casson models, but the difference between the flow rates of these models is high when the time parameter t lies between 0° and 180° and marginal when the time variable t lies between 180° and 360°. It is also noticed that for a given set of values of the parameters, the flow rate of the twofluid HB model is significantly higher than that of the singlefluid HB model and the flow rate of the twofluid Casson model is marginally higher than that of the singlefluid HB model. It is of interest to note that the plot of the singlefluid HB model is in good agreement with Figure of Sankar and Hemalatha [7] and the plot of the singlefluid Casson model is in good agreement with Figure of Dash et al. [3].
Figure 6 shows the influence of the nonNewtonian effects on the flow rate of the different twofluid models. The variation of the flow rate with yield stress for different twofluid models with , °, is depicted in Figure 6. It is observed that the flow rate decreases linearly with the increase of the yield stress for the twofluid HB model and it decreases very slowly with the increase of the yield stress for the twofluid Casson model. It is also noted that for a given set of values of the parameters, the flow rate decreases marginally with the increase of the powerlaw index n. One can note that the flow rate of the two fluid HB model is significantly higher than that of the twofluid Casson model when all of the other parameters were kept as constant.
3.5. Wall Shear Stress
The variation of the wall shear stress in a time cycle for different fluid models with , and is plotted in Figure 7. It is observed that the wall shear stress increases when the time parameter t increases from 0° to 90° and then it decreases as the time t increases from 90° to 270° and then it increases when the time t increases from 270° to 360°. The wall shear stress is maximum at 90° and minimum at 270°. It is found that for a given set of values of the parameters, the wall shear stress of the twofluid models are marginally lower than those of the singlefluid models. Also, it is noticed that the wall shear stress of the twofluid HB model is marginally lower than that of the twofluid Casson model. It is of interest to note that the plot of the singlefluid HB model is in good agreement with Figure of Sankar and Hemalatha [7] and the plot of the singlefluid Casson model is in good agreement with Figure of Dash et al. [3].
Figure 8 shows the effects of catheterization on wall shear stress of the different fluid models for blood flow through catheterized arteries. The variation of wall shear stress with catheter radius ratio for different fluid models with , , and is sketched in Figure 8. It is seen that the wall shear stress decreases almost linearly with the increase of the catheter radius ratio for all the fluid models. For a given set of values of the parameters, the wall shear stress of the twofluid models is slightly higher than that of the singlefluid models. There is not much of the difference between the wall shear stress of twofluid HB and Casson models.
3.6. Resistance to Flow
Figure 9 shows the variation of the longitudinal impedance to flow with yield stress for twofluid HB and Casson models with and °. It is observed that for the twofluid HB model, the longitudinal impedance to flow increases very slowly with the increase of the yield stress, but, for the twofluid Casson model, the impedance increases linearly when the yield stress increases from 0 to 0.15 and then it increases slowly when the yield stress increases from 0.15 to 0.25. The longitudinal impedance to flow of the twofluid HB model is significantly lower than that of the twofluid Casson model.
The increase in the longitudinal impedance due to the catheterization is defined as the ratio between the longitudinal impedance of a fluid model in a catheterized artery for a given set of values of the parameters and the longitudinal impedance of the same fluid in the uncatheterized artery for the same set of values of the parameters [23]. The estimates of the increase in the longitudinal impedance with effects on catheterization for different values of the yield stress for the twofluid HB and Casson models with , and ° are given in Table 1. For the range 0.1–0.5 of the catheter radius ratio, the range of increase in longitudinal impedance of the twofluid HB model are 1.34–4.26, 1.35–4.53, 1.37–4.86 and 1.39–5.26 when the yield stress values are 0.1, 0.15, 0.2 and 0.25, respectively. For the twofluid Casson model, the estimates of the increase in the longitudinal impedance increase are 1.42–7.47, 1.44–7.68, 1.45–7.75 and 1.46–7.86 when the yield stress values are 0.1, 0.15, 0.2 and 0.25, respectively. It is important to note that the estimates of the increase in the longitudinal impedance considerably much smaller for the twofluid HB model than those of the twofluid Casson model.

Catheters play an important role in the clinical investigations, since they are used to measure different types of flow quantities. Some types of catheters used in clinics, their sizes and their usage are mentioned in Table 2 [20], where is the diameter of the catheter and is the diameter of the artery. As an application of the present study to the medical field, the different types of the catheters (with sizes), which are used in the medical field [23], and the corresponding range of estimates of the increase in the longitudinal impedance for the twofluid HB and Casson models with , and and are computed in Table 3. It is observed that the range of estimates of the increase in the longitudinal impedance to flow for the twofluid HB model are significantly very lower than those of the twofluid Casson model. Hence, it is strongly felt that the twofluid HB model will have more applicability than the twofluid Casson model in the clinical use.


3.7. Usefulness of the Present Study
Figure 4 could be useful to the physicians in predicting the postcatheterization velocity profiles and thus, they can predict the effect of introducing the catheter on the velocity profiles and flow rate of the blood in the artery. Figure 8 might be useful to clinicians to predict and analyze the wall shear stress after inserting the catheter into the artery coaxially. Tables 1 and 3 might be used by clinicians to obtain the rough estimates of increase in longitudinal impedance due to the insertion of the catheter into the artery and the influence of nonNewtonian behavior of blood on impedance to flow. Since, catheters are used widely clinically; these estimates might be useful to physicians to decide their future course of action. Furthermore, as catheters are used clear the short occlusions or stenosis in the arterial wall, the present study could also be useful in estimating the increase in the longitudinal impedance and wall shear stress, since, the insertion of a catheter into the artery alters the flow field, modifies the pressure distribution and hence increases the flow resistance. Thus, there is considerable usefulness of the present study in the physiological context.
4. Conclusions
The pulsatile flow of blood through catheterized arteries is analyzed, assuming blood as a (i) twofluid Casson model and (ii) twofluid HerschelBulkley model. This study brings out the advantages of using the twofluid HerschelBulkley (HB) model over the twofluid Casson model for pulsatile blood flow through catheterized arteries. The effects of the catheterization, nonNewtonian nature of blood and pulsatility of the flow on the yield plane locations, velocity, flow rate, wall shear stress and longitudinal impedance are analyzed for different twofluid models. It is found that the width of the plugflow velocity, velocity distribution and thee flow rate for the twofluid HB model are considerably higher than those of the twofluid Casson model for a given set of values of the parameters. Also, it is observed that the longitudinal impedance is significantly very low for the twofluid HB model than those of the twofluid Casson model. It is of interest to note that the difference between the estimates of the increase in the longitudinal impedance of the twofluid HB model and the twofluid Casson model is substantial and hence, one can expect a significant increase in the flow of the twofluid HB model. Thus, it is concluded that the twofluid HB model will have more applicability in analyzing the blood flow through catheterized arteries.
References
 A. Sarkar and G. Jayaraman, “Correction to flow ratepressure drop relation in coronary angioplasty: steady streaming effect,” Journal of Biomechanics, vol. 31, pp. 781–791, 1998. View at: Google Scholar
 G. Jayaraman and K. Tiwari, “Flow in a catheterized curved artery,” Medical and Biological Engineering and Computing, vol. 33, pp. 1–6, 1995. View at: Google Scholar
 R. K. Dash, G. Jayaraman, and K. N. Metha, “Estimation of increased flow resistance in a narrow catheterized artery—a theoretical model,” Journal of Biomechanics, vol. 29, pp. 917–930, 1996. View at: Google Scholar
 R. V. Anderson, G. S. Roubin, P. P. Leimgruber et al., “Measurement of transstenotic pressure gradient during percutaneous transluminal coronary angioplasty,” Circulation, vol. 73, pp. 1223–1230, 1986. View at: Google Scholar
 R. Wilson, M. R. Johnson, M. L. Marcus et al., “The effect of coronary angioplasty on coronary flow reserve,” Circulation, vol. 77, pp. 873–885, 1988. View at: Google Scholar
 R. K. Dash, G. Jayaraman, and K. N. Metha, “Flow in a catheterized curved artery with stenosis,” Journal of Biomechanics, vol. 32, pp. 49–61, 1999. View at: Google Scholar
 D. S. Sankar and K. Hemalatha, “Pulsatile flow of HerschelBulkley fluid through catheterized arteries—a mathematical model,” Applied Mathematical Modeling, vol. 31, pp. 1497–1517, 2007. View at: Google Scholar
 L. H. Back, E. Y. Kwack, and M. R. Back, “Flow ratepressure drop relation in coronary angioplasty: catheter obstruction effect,” Journal of Biomechanical Engineering, vol. 118, pp. 83–89, 1996. View at: Google Scholar
 H. Kanai, E. Y. Kwack, and M. R. Back, “Flow ratepressure drop relation in coronary angioplasty: catheter obstruction effect,” Journal of Biomechanical Engineering, vol. 118, pp. 83–89, 1996. View at: Google Scholar
 T. A. MacMahon, C. Clark, V. S. Murthy, and A. H. Shapiro, “Intraaortic balloon experiment in a lumpedelement hydraulic model of the circulation,” Journal of Biomechanics, vol. 4, pp. 335–350, 1971. View at: Google Scholar
 K. V. Manjula and R. Devanathan, “A theoretical study of catheter probe in a stenosed artery,” in Biomechanics, K. B. Sahay and R. K. Saxena, Eds., pp. 240–246, Wiley Eastern, New Delhi, India, 1988. View at: Google Scholar
 L. H. Back, “Estimated mean flow resistance increase during coronary artery catheterization,” Journal of Biomechanics, vol. 27, pp. 169–175, 1994. View at: Google Scholar
 K. T. Karahalios, “Some possible effects of a catheter on the arterial wall,” Medical Physics, vol. 17, pp. 922–925, 1990. View at: Google Scholar
 G. Jayaraman and R. K. Dash, “Numerical study of flow in a constricted curved annulus: an application to flow in a catheterised artery,” Journal of Engineering Mathematics, vol. 40, no. 4, pp. 355–376, 2001. View at: Publisher Site  Google Scholar  Zentralblatt MATH  MathSciNet
 D. A. MacDonald, “Pulsatile flow in a catheterized artery,” Journal of Biomechanics, vol. 19, pp. 239–249, 1986. View at: Google Scholar
 R. K. Dash and P. Daripa, “Analytical and numerical studies of a singularly perturbed Boussinesq equation,” Applied Mathematics and Computation, vol. 126, no. 1, pp. 1–30, 2002. View at: Publisher Site  Google Scholar  Zentralblatt MATH  MathSciNet
 C. Tu and M. Deville, “Pulsatile flow of nonNewtonian fluids through arterial stenosis,” Journal of Biomechanics, vol. 29, pp. 899–908, 1996. View at: Google Scholar
 P. Chaturani and R. P. Samy, “Pulsatile flow of a Casson fluid through stenosed arteries with application to blood flow,” Biorheology, vol. 23, no. 5, pp. 499–511, 1986. View at: Google Scholar
 V. P. Srivastava and M. Saxena, “Twolayered model of Casson fluid flow through stenotic blood vessels: applications to the cardiovascular system,” Journal of Biomechanics, vol. 27, pp. 921–928, 1994. View at: Google Scholar
 D. S. Sankar and K. Hemalatha, “A nonNewtonian fluid flow model for blood flow through a catheterized artery—steady flow,” Applied Mathematical Modeling, vol. 31, pp. 1847–1864, 2007. View at: Google Scholar
 J. C. Misra and S. K. Pandey, “Peristaltic transport of blood in small vessels: study of a mathematical model,” Computers & Mathematics with Applications, vol. 43, no. 89, pp. 1183–1193, 2002. View at: Publisher Site  Google Scholar  Zentralblatt MATH  MathSciNet
 D. S. Sankar and U. Lee, “Twofluid nonlinear mathematical model for pulsatile blood flow through catheterized arteries,” Journal of Mechanical Science and Technology, vol. 23, pp. 1650–1669, 2009. View at: Google Scholar
 D. S. Sankar, “A twofluid model for pulsatile flow in catheterized blood vessels,” International Journal of NonLinear Mechanics, vol. 44, pp. 337–351, 2009. View at: Google Scholar
 P. Chaturani and R. P. Samy, “A study of nonNewtonian aspects of blood flow through stenosed arteries and its applications in arterial diseases,” Biorheology, vol. 22, pp. 521–531, 1985. View at: Google Scholar
 N. Iida, “Influence of plasma layer on steady blood flow in microvessels,” Japanese Journal of Applied Physics, vol. 17, pp. 203–214, 1978. View at: Google Scholar
 D. S. Sankar and U. Lee, “Twofluid nonlinear model for flow in catheterized blood vessels,” International Journal of NonLinear Mechanics, vol. 43, pp. 622–631, 2008. View at: Google Scholar
 E. W. Merrill, “Rheology of blood,” Physiological Reviews, vol. 49, pp. 863–888, 1969. View at: Google Scholar
Copyright
Copyright © 2010 D. S. Sankar and Usik Lee. This is an open access article distributed under the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.