Home | About | Journals | Submit | Contact Us | Français |

**|**J Biol Phys**|**v.37(1); 2011 January**|**PMC3006462

Formats

Article sections

- Abstract
- Introduction
- Physical model, governing equations and separation of waves
- Transmission and reflection coefficients and conditions for reflectionless arterial prosthesis
- Conclusion
- References

Authors

Related links

J Biol Phys. 2011 January; 37(1): 51–60.

Published online 2010 August 17. doi: 10.1007/s10867-010-9196-2

PMCID: PMC3006462

Laboratory of Modelling and Simulation in Engineering and Biological Physics, P.O. Box 812, Yaoundé, Cameroon

Guy Richard Kol, Email: kolguy_r/at/yahoo.fr.

Received 2010 April 17; Accepted 2010 July 15.

Copyright © Springer Science+Business Media B.V. 2010

We propose a new technique to characterize a reflectionless arterial prosthesis. The corresponding transmission and reflection coefficients are determined from the geometric and the elastic properties of the arterial wall, and the interaction between the latter and the prosthesis are studied accordingly.

Elastic tubes are widely used nowadays in fluid flow, and their importance no longer needs to be demonstrated. From the biological field up to the industrial domain, different types of elastic tubes are encountered and while many studies have been done, there still exist unresolved questions.

The first studies on elastic tubes presented a linear approach of the fluid flow therein with differences arising only in the consideration of the tube wall contribution [1, 2]. The nonlinear studies that came after brought out the importance of the fluid convective term and particularly took into account the dynamical properties of the tube wall where the nonlinear stress–strain relationship is considered [3–6]. Most of the works that followed contributed to improve the modeling, in a more realistic way, of elastic tubes [7–9] and took into account viscosity, tapering aspect of the tubes [5, 10], pathological cases of vessels and prostheses [11].

Some studies on blood waves focused on the study of the behaviour of waves in the regions of discontinuities [12, 13]. In general, any discontinuity zone present in an artery gives rise to the phenomena of wave transmission and reflection. These discontinuities are in general due to the change in the elastic properties of the arterial wall, diseases, bifurcations and prostheses.

The objective of this paper is to determine the expression of the transmission and reflection coefficients in the prosthesis using a nonlinear model of blood flow. This theory is based on the separation of pressure waves into their forward and backward running components proximal and distal to the prosthesis. We then use the same approach to investigate the characteristics of a reflectionless prosthesis.

In the study, we consider that the fluid in the vessel is Newtonian, viscous, homogeneous and incompressible. Furthermore, we take into account the variation of the radius and elasticity of the wall. The rest of the paper is therefore organized as follows: in Section 2, the physical and mathematical problems are formulated. In Section 3, the nonlinear model of the flow is used to determine the transmission and reflection coefficients of the prosthesis, and the conditions for a reflectionless prosthesis are derived accordingly. The last section is devoted to concluding remarks.

The artery is assumed to be conical, with a nonlinear stress–strain relation. The geometry of the artery with the presence of a prosthesis is illustrated in Fig. 1. The index *i* (*i*=1,2 or 3) represents each portion, where *i*=2 stands for the prosthesis and *i*=1 and *i*=3 represent the leftmost and rightmost portions of the vessel, respectively. For the present discrete case, we apply the averaging procedure, used in the literature, to the Navier–Stokes equations [5] and we derive the following one-dimensional model

1

2

where *z* is the axial coordinate, *t* is time, *P*_{i} is the pressure inside the vessel, *W*_{i} is the axial component of the fluid velocity, *R*_{i} is the radius of the vessel, *ρ* is the fluid density and *ν*_{i} is kinematic viscosity.

Geometry of the artery with prosthesis, where a stented vessel is shown in a cylindrical coordinate system (*r*, radial coordinate; *z*, longitudinal coordinate)

For the wall dynamics, we use the second law of Newton on a portion of the vessel wall and we obtain the relation [14]:

3

where is the pressure outside the vessel, *δR*_{i} denotes the vessel deformation, *ρ*_{0i} is the wall density, and *H*_{i} and *h*_{i} are the effective inertial thickness and the thickness of the wall, respectively. *γ*_{i} is the viscoelasticity coefficient due to the absorption effects of the material. *β*_{i} is the shear modulus, and _{i} is another coefficient of viscoelasticity. *σ*_{ti} is the approximated exponential function for the stress–strain relationship given in [15, 16]. This approximated relation is widely known as [3, 4, 6, 7]

4

where is the stationary radius of the vessel, *E*_{i} is the Young modulus and *a*_{i} represents the nonlinear coefficient of elasticity in the portion *i* .

The vessel radius *R*_{i}(*z*,*t*) in the portion *i* is described as

5

As vessel segments are tapered elastic tubes with increasing rigidity away from the heart, they represent general cases of thin-walled elastic tubes with variable Young modulus. Many authors [4, 7, 8] showed that the stationary radius and the variation of elasticity along the wall will obey the relations

6

with *f* (*z*)=−*m*_{i}(*z*−*z*_{0i}), *g*(*z*)=*n*_{i}(*z*−*z*_{0i}). Here, *f* (*z*) denotes the decrease of the tube radius and *g*(*z*) the increase of the tube wall rigidity along the tube originating from the heart. *R*_{0i} is the radius at the entrance of the portion *i* and *z*_{0i} the abscissa at the entrance of the vessel segment, in the same portion. The positive coefficients *m*_{i} and *n*_{i} characterize the rate of decrease and increase of the radius and wall rigidity in the portion *i*, respectively.

As we consider weak wall displacement, we introduce the approximation:

7

where and are the normalized cross sectional area of the tube and the cross sectional area at the entrance of the *i*th portions, respectively. *R*_{0}, *w*_{0} and *E*_{0} are, respectively, the reference values of radius, velocity and Young modulus at the entrance of the vessel. We define the viscosity coefficients and as:

8

and the dimensionless quantities

9

and we also include the following coefficients

10

We assume that the first and last portions (*i*=1 and *i*=3) are made up of the same material properties.

Following the mathematical procedure of [3, 17], we set

14

15

16

where *ϵ* is a small parameter measuring the weakness of dispersion and nonlinearity. Also weak are the parameters modelling the taper effect of the wall, the compliance variation and the blood viscosity which become

17

Next, the coordinate system is changed to slowly varying coordinates

18

where *c*_{i} represents the wave group velocity and is assumed to be the velocity of the linear wave as it enters the medium. *ξ*_{i} is an independent variable representing the forward wave whereas *η*_{i} stands for the backward wave.

The approximation of the flow motion at the first order, *ϵ*^{1}, gives:

19

20

21

In the present work, the study will be limited to this first order of perturbation. The second order, useful to obtain the analytic expression of the wave, is not used here. So, we consider the waves to be separated explicitly as forward and backward elements (where *f* and *b* are, respectively, used as superscripts) [18, 19].

22

23

24

In this part, we propose a new technique for the estimation of the transmission, *T*, and the reflection, *B*, coefficients in the prosthesis proportional to the local transmission, *T*_{i}, and reflection, *B*_{i}, coefficients at the interface of the prosthesis.

For the computation of the characteristic coefficients, we use the impedance of the system given by the ratio of the pressure and flow at the interfaces [16, 17]

28

The local transmission, *T*_{i}, and reflection, *B*_{i}, coefficients are defined as

29

and are expressed in terms of the flows. Of course, in the literature, this is not usual, but from the conservation law the relation *T*_{i}+*B*_{i}=1 is not verified when the pressures are mainly used. This also brings out the difference between the method used in this paper and the one used by Stergiopulos et al. [18].

If we consider at the interfaces the conservation of mass and energy, it will appear that, at the first interface,

30

while at the second interface

31

Inserting Eqs. 25–27 into Eqs. 30 and 31 helps to obtain the local transmission coefficient *T*_{1} (resp. *T*_{2}) and the local reflection coefficient *B*_{1} (resp. *B*_{2}), respectively, at the first and second interfaces

32

33

with *L*_{v} as the length of the left portion of the vessel and *L*_{p} as the length of the prosthesis.

In this work, we suppose that the distal forward wave, , is the sum of two waves: the part of the proximal forward wave that became transmitted through the prosthesis and the part of the distal backward wave that became reflected by the prosthesis. And we obtain a similar equation for the proximal backward running wave, . These equations are given by

34

where *T* and *B* are the global transmission and reflection coefficients due to the overall prosthesis. Using Eq. 34, we obtain the expressions of the transmission and reflection coefficients, as

35

From Eq. 37, we get the conditions for a reflectionless arterial prostheses. The reflectionless arterial prosthesis is obtained when the reflection coefficient is equal to 0 and the transmission coefficient equal to 1. This gives us the following condition (with *B*_{1}≠0):

38

with

39

and

40

Using this condition (see relation 38), the optimal thickness of the prosthesis is derived as

41

The manufacturing of reflectionless arterial prostheses being one of the primary objectives of research in this field, our result gives a new mathematical condition for the reflectionless arterial prosthesis. This condition takes into account the geometric and elastic properties of the vessel wall.

For numerical calculation, we have used the characteristic parameters for the femoral artery of a dog [6]: *L*_{v}=4.5 cm, dyn/cm^{2}, *n*_{1}=0.067 cm^{−1}, *m*_{1}=0.080 cm^{−1} and the thickness of the vessel is 0.018 cm. We consider the prosthesis parameters to be *L*_{p}=4.0 cm, *n*_{2}=0.069 cm^{−1} and *m*_{2}=0.089 cm^{−1}.

Figure 2 shows that the transmission coefficient increases to the value 1 on the positive side of the ordinate and the reflection coefficient increases to the value 0 on the negative side, as the ratio of the product of the Young modulus and the wall thickness is increased and for *R*_{01}/*R*_{02}=0.86. This behaviour is comparable with the effects of the severity of an extended stenosis as obtained by Stergiopulos et al. [18]. The negative values obtained in the case of the reflection coefficient are due to the fact that the reflected waves velocity is negative (see Eq. 26).

The dependence of the transmission and reflection coefficients on the ratio of the product of Young modulus and wall thickness of the stented vessel for *R*_{01}/*R*_{02}=0.86

In Fig. 3, the ratio of thickness (Eq. 41) is plotted as the ratio of the radius varies. When the ratio of the radii increases, the thickness ratio decreases. This figure illustrates a reflectionless arterial prosthesis. For example, in the case of a Palmaz stent with Young modulus dyn/cm^{2} [20], inserted in the femoral artery, the ratio of the thickness of the stent to that of the vessel is . Using this condition, the value of the transmission coefficient is approximately equal to 1 while the reflection coefficient is equal to 0, when the ratio of the radius is *R*_{01}/*R*_{02}=0.94. Taking into account the thickness of the femoral artery (*h*_{01}=0.018 cm), the thickness of the stent is obtained as *h*_{02}= 0.173 cm.

Dependence of the ratio of thickness on that of the radius in the case of the reflectionless prosthesis

In Fig. 4, we show that the variation of the length of the left portion of the vessel brings about modifications in the characteristics of the perfect prosthesis. When the length of the vessel increases, the characteristics of the reflectionless prosthesis decrease. As a whole, these results present the interactions between the vessel wall and the prosthesis. Such a behaviour would also be observed for the variation of the elasticity of the vessel. These results are further confirmed by the case of Fig. 5, where the variation of the prosthesis length is taken into account. So, when the length of the prosthesis increases, the characteristics of the reflectionless prosthesis increase. In fact, if there is a perturbation at the entrance of the prothesis, that prothesis, due to its characteristics will overcome the perturbation for a better blood flow. That is why increasing the length of the prothesis deeply modifies the condition for a reflectionless prothesis as depicted in Fig. 5.

In this study, the concept of wave separation has been used to derive the reflection and transmission coefficients of blood waves created by a prosthesis inserted in an artery. This has led to the derivation of conditions for a reflectionless prosthesis. These conditions are useful for the design of a reflectionless arterial prosthesis. In our methodology, we take explicitly into account the interactions between the vessel wall and the prosthesis at the entrance and the exit of the prosthesis.

1. Fung YC. New York. New York: Springer; 1984. Biodynamics: Circulation.

2. Pedley TJ. Perspectives in Fluid Mechanics. Cambridge: Cambridge University Press; 2000.

3. Yomosa S. Solitary waves in large blood vessels J. Phys. Soc. Jpn. 1987. 56506–520.5208931161987JPSJ...56..506Y. doi: 10.1143/JPSJ.56.506. [Cross Ref]

4. Noubissie S, Woafo P. Dynamics of solitary blood waves in arteries with prostheses. Phys. Rev. E. 2003;67:419111–419118. doi: 10.1103/PhysRevE.67.041911. [PubMed] [Cross Ref]

5. Demiray H. Weakly nonlinear waves in a viscous fluid contained in a viscoelastic tube with variable cross-section Eur. J. Mech. 2005. 24337–347.3471069.740162164364. doi: 10.1016/j.euromechsol.2004.12.002. [Cross Ref]

6. Ntchantcho R, Noubissie S, Woafo P. Numerical simulation of solitary blood waves in an elastic tube subjected to a localised deformation Commun. Nonlinear Sci. Numer. Simulat. 2007. 121572–1583.15831118.3503523326472007CNSNS..12.1572N. doi: 10.1016/j.cnsns.2006.03.009. [Cross Ref]

7. Paquerot JF, Remoissenet M. Dynamics of nonlinear blood pressure waves in large arteries Phys. Lett. A 1994. 19477–82.821994PhLA..194...77P. doi: 10.1016/0375-9601(94)00729-9. [Cross Ref]

8. Noubissie S, Woafo P. Dynamics of solitary waves through taper-thin elastic tube with localised deformation Phys. Scr. 2004. 69249–256.2561067.761052004PhyS...69..249N. doi: 10.1238/Physica.Regular.069a00249. [Cross Ref]

9. Paquerot JF, Lambrakos SG. Monovariable representation of blood flow in large elastic arteries Phys. Rev. E 1994. 493432–3439.34391994PhRvE..49.3432P. doi: 10.1103/PhysRevE.49.3432. [PubMed] [Cross Ref]

10. Duan WS, Wang BR, Wei RJ. The decay of soliton in small blood arteries J. Phys. Soc. Jpn. 1994. 65945–947.9471996JPSJ...65..945D. doi: 10.1143/JPSJ.65.945. [Cross Ref]

11. Tortoriello A, Pedrizzetti G. Flow-tissue interaction with compliance mismatch in a model stented artery. J. Biomech. 2004;37:1–11. doi: 10.1016/S0021-9290(03)00259-8. [PubMed] [Cross Ref]

12. Khir AW, Parker KH. Measurements of wave speed and reflected waves in elastic tubes and bifurcations. J. Biomech. 2002;35:775–783. doi: 10.1016/S0021-9290(02)00025-8. [PubMed] [Cross Ref]

13. Khir AW, Parker KH. Wave intensity in the ascending aorta: effects of arterial occlusion. J. Biomech. 2005;38:647–655. doi: 10.1016/j.jbiomech.2004.05.039. [PubMed] [Cross Ref]

14. Formaggia L, Lamponi D, Quarteroni A. One-dimensional models for blood flow in arteries J. Eng. Math. 2003. 47251–276.2761070.760592038983. doi: 10.1023/B:ENGI.0000007980.01347.29. [Cross Ref]

15. Akio S, Masamitsu H, Yoshihisa U. Pressure pulse wave for blood flow in the aorta from the viewpoint of the nonlinear Toda lattice. Phys. Lett. A. 1996;221:395–399. doi: 10.1016/0375-9601(96)00563-4. [Cross Ref]

16. Xie J, Zhou J, Fung YC. Bending of blood vessel wall: stress–strain laws of the intima-media and adventitial layers. J. Biomech. Eng. 1995;117:136–145. doi: 10.1115/1.2792261. [PubMed] [Cross Ref]

17. Duan WS, Wang BR, Wei RJ. Reflection and transmission blood waves due to arterial branching Phys. Rev. E 1997. 551773–1778.17781997PhRvE..55.1773D. doi: 10.1103/PhysRevE.55.1773. [Cross Ref]

18. Stergiopulos N, Spiridon M, Pythoud F, Meister JJ. On the wave transmission and reflection properties of stenoses. J. Biomech. 1996;29:31–38. doi: 10.1016/0021-9290(95)00023-2. [PubMed] [Cross Ref]

19. Li JK, editor. Dynamics of the vascular system. Singapore: World Scientific; 2004.

Articles from Journal of Biological Physics are provided here courtesy of **Springer Science+Business Media B.V.**

PubMed Central Canada is a service of the Canadian Institutes of Health Research (CIHR) working in partnership with the National Research Council's Canada Institute for Scientific and Technical Information in cooperation with the National Center for Biotechnology Information at the U.S. National Library of Medicine(NCBI/NLM). It includes content provided to the PubMed Central International archive by participating publishers. |