chapter 3 physical and flow properties … 3 physical and flow properties of blood david elad and...

3.1

CHAPTER 3

PHYSICAL AND FLOWPROPERTIES OF BLOOD

David Elad and Shmuel EinavTel Aviv University, Tel Aviv, Israel

3.1 PHYSIOLOGY OF THE CIRCULATORY 3.5 BLOOD FLOW IN THESYSTEM 3.1 MICROCIRCULATION 3.16

3.2 PHYSICAL PROPERTIES OF BLOOD 3.4 3.6 BLOOD FLOW IN THE HEART 3.183.3 BLOOD FLOW IN ARTERIES 3.5 3.7 ANALOG MODELS OF BLOOD FLOW3.4 BLOOD FLOW IN VEINS 3.14 3.21

REFERENCES 3.23

The circulatory transport system is responsible for oxygen and nutrient supply to all body tissues andremoval of waste products. The discovery of the circulation of blood in the human body is related toWilliam Harvey (1578–1657). The circulatory system consists of the heart—the pump that generatesthe pressure gradients needed to drive blood to all body tissues, the blood vessels—the deliveryroutes, and the blood—the transport medium for the delivered materials. The blood travels continu-ously through two separate loops; both originate and terminate at the heart. The pulmonary circula-tion carries blood between the heart and the lungs, whereas the systemic circulation carries bloodbetween the heart and all other organs and body tissues (Fig. 3.1). In both systems blood is trans-ported in the vascular bed because of a pressure gradient through the following subdivisions: arteries,arterioles, capillaries, venules, and veins. The cardiac cycle is composed of the diastole, during whichthe ventricles are filling with blood, and the systole, during which the ventricles are actively contract-ing and pumping blood out of the heart (Martini, 1995; Thibodeau and Patton, 1999).

The total blood volume is unevenly distributed. About 84 percent of the entire blood volume is inthe systemic circulation, with 64 percent in the veins, 13 percent in the arteries, and 7 percent in thearterioles and capillaries. The heart contains 7 percent of blood volume and the pulmonary vessels 9percent. At normal resting activities heart rate of an adult is about 75 beats/min with a stroke volumeof typically 70 mL/beat. The cardiac output, the amount of blood pumped each minute, is thus 5.25L/min. It declines with age. During intense exercise, heart rate may increase to 150 beats/min andstroke volume to 130 mL/beat, providing a cardiac output of about 20 L/min. Under normalconditions the distribution of blood flow to the various organs is brain, 14 percent; heart, 4 percent;kidneys, 22 percent; liver, 27 percent; inactive muscles, 15 percent; bones, 5 percent; skin, 6 percent;

3.1 PHYSIOLOGY OF THE CIRCULATORY SYSTEM

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3.2 MECHANICS OF THE HUMAN BODY

bronchi, 2 percent. The averaged blood velocity in the aorta (cross-sctional area of 2.5 cm2) is 33 cm/s, while in the capillaries (cross-sectional area of 2500 cm2) it is about 0.3 mm/s. The blood remainsin the capillaries 1 to 3 seconds (Guyton and Hall, 1996; Saladin, 2001).

At normal conditions, the pulsatile pumping of the heart is inducing an arterial pressure thatfluctuates between the systolic pressure of 120 mmHg and the diastolic pressure of 80 mmHg (Fig.3.2). The pressure in the systematic capillaries varies between 35 mmHg near the arterioles to 10mmHg near the venous end, with a functional average of about 17 mmHg. When blood terminatesthrough the venae cavae into the right atrium of the heart, its pressure is about 0 mmHg. When theheart ejects blood into the aorta, a pressure pulse is transmitted through the arterial system. Thetraveling velocity of the pressure pulse increases as the vessel’s compliance decreases; in the aorta itis 3 to 5 m/s, in the large arteries 7 to 10 m/s, and in small arteries 15 to 35 m/s. Figure 3.3 depictsan example of the variations in the velocity and pressure waves as the pulse wave travels towardperipheral arteries (Caro et al., 1978; Fung, 1984).

FIGURE 3.1 General organization of the circulatory system with averaged values of normal blood flow to major organs.



PHYSICAL AND FLOW PROPERTIES OF BLOOD

3.3

FIGURE 3.2 Variation of blood pressure in the circulatory system.

FIGURE 3.3 Pressure and flow waveforms in different arteries of the human arterial tree. [From Mills et al.(1970) by permission.]





Blood is a suspension of cellular elements—red blood cells (erythrocytes), white cells (leukocytes),and platelets—in an aqueous electrolyte solution, the plasma. Red blood cells (RBC) are shaped as abiconcave saucer with typical dimensions of 2 × 8 µm. Erythrocytes are slightly heavier than theplasma (1.10 g/cm3 against 1.03 g/cm3); thus they can be separated by centrifugation from theplasma. In normal blood they occupy about 45 percent of the total volume. Although larger thanerythrocytes, the white cells are less than 1/600th as numerous as the red cells. The platelet concen-tration is 1/20th of the red cell concentration, and their dimensions are smaller (2.5 µm in diameter).The most important variable is the hematocrit, which defines the volumetric fraction of the RBCs inthe blood. The plasma contains 90 percent of its mass in water and 7 percent in the principal proteinsalbumin, globulin, lipoprotein, and fibrinogen. Albumin and globulin are essential in maintaining cellviability. The lipoproteins carry lipids (fat) to the cells to provide much of the fuel of the body. Theosmotic balance controls the fluid exchange between blood and tissues. The mass density of bloodhas a constant value of 1.05 g/cm3 for all mammals and is only slightly greater than that of water atroom temperature (about 1 g/cm3).

The macroscopic rheologic properties of blood are determined by its constituents. At a normalphysiological hematocrit of 45 percent, the viscosity of blood is µ = 4 × 10-2 dyne · s/cm2 (or poise),which is roughly 4 times that of water. Plasma alone (zero hematocrit) has a viscosity of µ = 1.1 × 10-2 to1.6 × 10-2 poise, depending upon the concentration of plasma proteins. After a heavy meal, when theconcentration of lipoproteins is high, the plasma viscosity is quite elevated (Whitmore, 1968). Inlarge arteries, the shear stress (τ) exerted on blood elements is linear with the rate of shear, and bloodbehaves as a newtonian fluid, for which,

(3.1)

where u is blood velocity and r is the radial coordinate perpendicular to the vessel wall.In the smaller arteries, the shear stress acting on blood elements is not linear with shear rate, and

the blood exhibits a nonnewtonian behavior. Different relationships have been proposed for thenonnewtonian characteristics of blood, for example, the power-law fluid,

(3.2)

where K is a constant coefficient. Another model, the Casson fluid (Casson, 1959), was proposed bymany investigators as a useful empirical model for blood (Cokelet, 1980; Charm and Kurland,1965),

(3.3)

where τy is the fluid yield stress.

3.2.2 Blood Rheology

3.2.1 Constituents of Blood

3.2 PHYSICAL PROPERTIES OF BLOOD




PHYSICAL AND FLOW PROPERTIES OF BLOOD 3.5

The aorta and arteries have a low resistance to blood flow compared with the arterioles and capillar-ies. When the ventricle contracts, a volume of blood is rapidly ejected into the arterial vessels. Sincethe outflow to the arteriole is relatively slow because of their high resistance to flow, the arteries areinflated to accommodate the extra blood volume. During diastole, the elastic recoil of the arteriesforces the blood out into the arterioles. Thus, the elastic properties of the arteries help to convert thepulsatile flow of blood from the heart into a more continuous flow through the rest of the circulation.Hemodynamics is a term used to describe the mechanisms that affect the dynamics of bloodcirculation.

An accurate model of blood flow in the arteries would include the following realistic features

1. The flow is pulsatile, with a time history containing major frequency components up to the eighthharmonic of the heart period.

2. The arteries are elastic and tapered tubes.3. The geometry of the arteries is complex and includes tapered, curved, and branching tubes.4. In small arteries, the viscosity depends upon vessel radius and shear rate.

Such a complex model has never been accomplished. But each of the features above has been“isolated,” and qualitative if not quantitative models have been derived. As is so often the case in theengineering analysis of a complex system, the model derived is a function of the major phenomenaone wishes to illustrate.

The general time-dependent governing equations of fluid flow in a straight cylindrical tube aregiven by the continuity and the Navier-Stokes equations in cylindrical coordinates,

(3.4)

(3.5)

(3.6)

Here, u and v are the axial and radial components of the fluid velocity, r and z are the radial and axialcoordinates, and ρ and µ are the fluid density and viscosity, respectively. Eqs. (3.5) and (3.6) are themomentum balance equations in the z and r directions.

The simplest model of steady laminar flow in a uniform circular cylinder is known as the Hagen-Poiseuille flow. For axisymmetric flow in a circular tube of internal radius R0 and length l, theboundary conditions are

(3.7)

3.3.1 Introduction

3.3.2 Steady Flow

3.3 BLOOD FLOW IN ARTERIES





For a uniform pressure gradient (∆P) along a tube we get the parabolic Poiseuille solution

(3.8)

The maximal velocity umax = (R0)2 ∆P/4µl is obtained at r = 0.The Poiseuille equation indicates that the pressure gradient ∆P required to produce a volumetric

flow Q = uA increases in proportion to Q. Accordingly, the vascular resistance R will be defined as

(3.9)

If the flow is measured in cm3/s and P in dyn/cm2, the units of R are dyn · s/cm5. If pressure ismeasured in mmHg and flow in cm3/s, resistance is expressed in “peripheral resistance units,” or PRU.

The arteries are composed of elastin and collagen fibers and smooth muscles in a complexcircumferential organization with a variable helix. Accordingly, the arteries are compliant vessels, andtheir wall stiffness increases with deformation, as in all other connective tissues. Because of theirability to expand as transmural pressure increases, blood vessels may function to store blood volumeunder pressure. In this sense, they function as capacitance elements, similar to storage tanks. Thelinear relationship between the volume V and the pressure defines the capacitance of the storageelement, or the vascular capacitance:

(3.10)

Note that the capacitance (or compliance) decreases with increasing pressure, and also decreases withage. Veins have a much larger capacitance than arteries and, in fact, are often referred to as capaci-tance or storage vessels.

Another simple and useful expression is the arterial compliance per unit length, Cu, that can bederived when the tube cross-sectional area A is related to the internal pressure A = A(P, z). For a thin-wall elastic tube (with internal radius R0 and wall thickness h), which is made of a hookean material(with Young modulus E), one can obtain the following useful relation,

(3.11)

Arterial pulse propagation varies along the circulatory system as a result of the complex geometryand nonuniform structure of the arteries. In order to learn the basic facts of arterial pulse character-istics, we assumed an idealized case of an infinitely long circular elastic tube that contains a homog-enous, incompressible, and nonviscous fluid (Fig. 3.4). In order to analyze the velocity of propaga-tion of the arterial pulse, we assume a local perturbation, for example, in the tube cross-sectional area,that propagates along the tube at a constant velocity c.

The one-dimensional equations for conservation of mass and momentum for this idealized caseare, respectively (Pedley, 1980; Fung, 1984),

3.3.3 Wave Propagation in Arteries





(3.12)

(3.13)

where A(z) = tube cross-sectional area u(z) = uniform axial velocity of blood P(z) = pressure in the tube ρ = fluid viscosity z = axial coordinate t = time

The elasticity of the tube wall can be prescribed by relationship between the local tube pressure andthe cross-sectional area, P(A).

We further assume that the perturbation is small, while the wave length is very large compared withthe tube radius. Thus, the nonlinear inertia variables are negligible and the linearized conservationequation of mass and momentum become, respectively,

(3.14)

(3.15)

Next, we differentiate Eq. (3.14) with respect to t and Eq. (3.15) with respect to z, and upon addingthe results we obtain the following wave equation:

(3.16)

FIGURE 3.4 Cross section of artery showing the change of volume andvolume flow.





for which the wave speed c is given by

(3.17)

This suggests that blood pressure disturbances propagate in a wavelike manner from the heart towardthe periphery of the circulation with a wave speed c. For a thin-wall elastic tube (with internal radiusR0 and wall thickness h), which is made of a hookean material (with Young modulus E) and subjectedto a small increase of internal pressure, the wave speed c can be expressed as

(3.18)

This equation was obtained by Thomas Young in 1808, and is known as the Moens-Kortweg wavespeed. The Moens-Kortweg wave speed varies not only with axial distance but also with pressure. Thedominant pressure-dependent term in Eq. (3.18) is E, the modulus of elasticity; it increases withincreasing transmural pressure as the stiff collagen fibers bear more of the tension of the artery wall.The high-pressure portion of the pressure wave therefore travels at a higher velocity than the low-pressure portions of the wave, leading to a steepening of the pressure front as it travels from the hearttoward the peripheral circulation (Fig. 3.5). Wave speed also varies with age because of the decreasein the elasticity of arteries.

The arteries are not infinitely long, and it is possible for the wave to reflect from the distal end andtravel back up the artery to add to new waves emanating from the heart. The sum of all suchpropagated and reflected waves yields the pressure at each point along the arterial tree. Branching isclearly an important contributor to the measured pressures in the major arteries; there is a partialreflection each time the total cross section of the vessel changes abruptly.

Blood flow in the large arteries is driven by the heart, and accordingly it is a pulsating flow. Thesimplest model for pulsatile flow was developed by Womersley (1955a) for a fully developed oscil-latory flow of an incompressible fluid in a rigid, straight circular cylinder. The problem is defined fora sinusoidal pressure gradient composed from sinuses and cosinuses,

(3.19)

where the oscillatory frequency is ω/2π. Insertion of Eq. (3.19) into Eq. (3.5) yields

(3.20)

The solution is obtained by separation of variables as follows:

(3.21)

3.3.4 Pulsatile Flow

FIGURE 3.5 Steepening of a pressure pulse with distance along an artery.





Insertion of Eq. (3.21) into (3.20) yields the Bessel equation,

(3.22)

The solution for Eq. (3.22) is

(3.23)

where J0 is a Bessel function of order zero of the first kind, υ = µ/ρ is the kinematic viscosity, and ais a dimensionless parameter known as the Womersley number and given by

(3.24)

When α is large, the velocity profile becomes blunt (Fig. 3.6).

Pulsatile flow in an elastic vessel is very complex, since the tube is able to undergo local deforma-tions in both longitudinal and circumferential directions. The unsteady component of the pulsatileflow is assumed to be induced by propagation of small waves in a pressurized elastic tube. Themathematical approach is based on the classical model for the fluid-structure interaction problem,which describes the dynamic equilibrium between the fluid and the tube thin wall (Womersley,1955b; Atabek and Lew, 1966). The dynamic equilibrium is expressed by the hydrodynamic equa-tions (Navier-Stokes) for the incompressible fluid flow and the equations of motion for the wall of anelastic tube, which are coupled together by the boundary conditions at the fluid-wall interface. Themotion of the liquid is described in a fixed laboratory coordinate system ( , θ, ), and the dynamic

FIGURE 3.6 Theoretical velocity profiles of an oscillating flow resulting from a sinusoidal pressuregradient (cos ω) in a pipe, α is the Womersley number. Profiles are plotted for intervals of ∆ωt = 15°. Forωt > 180°, the velocity profiles are of the same form but opposite in sign. [From Nichols and O’Rourke(1998) by permission.]





equilibrium of a tube element in its deformed state is expressed in a lagrangian (material) coordinatesystem ( , , θ), which is attached to the surface of the tube (Fig. 3.7).

The first-order approximations for the axial (u1) and radial (ν1) components of the fluid velocity, andthe pressure (P1) as a function of time (t) and space (r, z), are given by

(3.25)

(3.26)

(3.27)

The dimensionless parameters m, x, k, τθ, and F10 are related to the material properties and defined as

(3.28)

FIGURE 3.7 Mechanics of the arterial wall: (a) axisymmetricwall deformation; (b) element of the tube wall under biaxialloading. The T’s are longitudinal and circumferential internalstresses.





where c = wave speedω = 2πHR/60 = angular frequency HR = heart rate A

1= input pressure amplitude

J0 and J

1= Bessel functions of order 0 and 1 of the first kind

ρF and ρ

T= blood and wall densities

R0

= undisturbed radius of the tube

Excellent recent summaries on pulsatile blood flow may be found in Nichols and O’Rourke (1998)and Zamir (2000).

Turbulence has been shown to exist in large arteries of a living system. It is especially pronouncedwhen the flow rate increases in exercise conditions (Yamaguchi and Parker, 1983). Turbulence ischaracterized by the appearance of random fluctuations in the flow. The transition to turbulence is avery complex procedure, which schematically can be described by a hierarchy of motions: growth oftwo-dimensional infinitesimal disturbances to final amplitudes, three-dimensionality of the flow, anda maze of complex nonlinear interactions among the large-amplitude, three-dimensional modes re-sulting in a final, usually stochastically steady but instantaneously random, motion called turbulentflow (Akhavan et al., 1991; Einav and Sokolov, 1993).

In a turbulent flow field, all the dynamic properties (e.g., velocity, pressure, vorticity) are randomfunctions of position and time. One thus looks at the statistical aspects of the flow characteristics (e.g.,mean velocity, rms turbulent intensity). These quantities are meaningful if the flow is stochasticallyrandom (i.e., its statistics are independent of time) (Nerem and Rumberger, 1976). The time averageof any random quantity is given by

(3.29)

One can thus decompose the instantaneous variables u and v as follows:

(3.30)

(3.31)

(3.32)

We assume that u′ is a velocity fluctuation in the x direction only and ν′ in the y direction only. Theoverbar denotes time average, so that by definition, the averages of u′, ν′, and ρ′ fluctuations are zero(stochastically random), and the partial derivatives in time of the mean quantities , , are zeros(the Reynolds turbulence decomposition approach, according to which velocities and pressures canbe decomposed to time-dependent and time-independent components).

By replacing u with etc. in the Navier-Stokes equation and taking time average, it can beshown that for the turbulent case the two-dimensional Navier-Stokes equation in cartesian coordinatesbecomes

(3.33)

3.3.5 Turbulence





and for the y direction one obtains

(3.34)

Integration yields

(3.35)

As ν′ must vanish near the wall, the values of P + ρgz will be larger near the wall. That implies that,in a turbulent boundary layer, the pressure does not change hydrostatically (is not height or depthdependent), as is the case of laminar flow. Equation (3.35) implies that the pressure is not a functionof y, and thus,

(3.36)

Since P is independent of y, integration in y yields

(3.37)

where C1 = τ0 is the shear stress near the wall.We see that in addition to the convective, pressure, and viscous terms, we have an additional term,

which is the gradient of the nonlinear term ρu′ν′, which represents the average transverse transport oflongitudinal momentum due to the turbulent fluctuations. It appears as a pseudo-stress along with theviscous stress µ∂U/∂y, and is called the Reynolds stress. This term is usually large in most turbulentshear flows (Lieber and Giddens, 1988).

The arteries and veins are generally not straight uniform tubes but have some curved structure,especially the aorta, which has a complex three-dimensional curved geometry with multiplanar cur-vature. To understand the effect of curvature on blood flow, we will discuss the simple case of steadylaminar flow in an in-plane curved tube (Fig. 3.8). When a steady fluid flow enters a curved pipe inthe horizontal plane, all of its elements are subjected to a centripetal acceleration normal to theiroriginal directions and directed toward the bend center. This force is supplied by a pressure gradientin the plane of the bend, which is more or less uniform across the cross section. Hence, all the fluidelements experience approximately the same sideways acceleration, and the faster-moving elementswith the greater inertia will thus change their direction less rapidly than the slower-moving ones. Thenet result is that the faster-moving elements that originally occupy the core fluid near the center of thetube are swept toward the outside of the bend along the diametrical plane, and their place is taken byan inward circumferential motion of the slower moving fluid located near the walls. Consequently,the overall flow field is composed of an outward-skewed axial component on which is superimposeda secondary flow circulation of two counter-rotating vortices.

The analytical solution for a fully developed, steady viscous flow in a curved tube of circularcross section was developed by Dean in 1927, who expressed the ratio of centrifugal inertial forcesto the viscous forces (analogous to the definition of Reynolds number Re) by the dimensionless Deannumber,

(3.38)

3.3.6 Flow in Curved Tubes





FIGURE 3.8 Schematic description of the skewed axial velocity profile andthe secondary motions developed in a laminar flow in a curved tube.

FIGURE 3.9 Qualitative illustration of laminar flow downstream of a bifurcation with a possible region offlow separation, secondary flow, and skewed axial profile.





where r is the tube radius and Rcurve is the radius of curvature. As De increases, the maximal axialvelocity is more skewed toward the outer wall. Dean’s analytic solutions are limited to small ratios ofradius to radius of curvature for which De < 96. However, numerical solutions extended the range upto 5000. Blood flow in the aortic arch is complex and topics such as entry flow from the aortic valve,pulsatile flow, and their influence on wall shear stress have been the subject of numerous experimentaland numerical studies (Pedley, 1980; Berger et al., 1983; Chandran, 2001).

The arterial system is a complex asymmetric multigeneration system of branching and bifurcatingtubes that distribute blood to all organs and tissues. A simplified arterial bifurcation may be repre-sented by two curved tubes attached to a straight mother tube. Accordingly, the pattern of blood flowdownstream of the flow divider (i.e., bifurcating region) is in general similar to flow in curved tubes(Fig. 3.9). Typically, a boundary layer is generated on the inside wall downstream from the flowdivider, with the maximum axial velocity just outside the boundary layer. As in flow in curved tubes,the maximal axial velocity is skewed toward the outer curvature, which is the inner wall of thebifurcation.

Comprehensive experimental and computational studies were conducted to explore the pattern ofblood flow in a branching vessel, energy losses, and the level of wall shear stress in the branch region(Ku and Giddens, 1987; Pinchak and Ostrach, 1976; Liepsch et al., 1989; Liepsch, 1993; Pedley,1995; Perktold and Rappitsch, 1995). Of special interest are the carotid bifurcation and the lowerextremity bypass graft-to-artery anastomosis whose blockage may induce stroke and walkinginability, respectively. A recent review of computational studies of blood flow through bifurcatinggeometries that may aid the design of carotid endartectomy for stroke prevention and graft-to-arteryconfiguration may be found in Kleinstreuer et al., 2001.

The veins are thin-walled tubular structures that may “collapse” (i.e., the cross-sectional area does notmaintain its circular shape and becomes less than in the unstressed geometry) when subjected tonegative transmural pressures P (internal minus external pressures). Experimental studies (Moreno etal., 1970) demonstrated that the structural performance of veins is similar to that of thin-walled elastictubes (Fig. 3.10). Three regions may be identified in a vein subjected to a transmural pressure: WhenP < 0, the tube is inflated, its cross section increases and maintains a circular shape; when P > 0, thetube cross section collapses first to an ellipse shape; and at a certain negative transmural pressure, acontact is obtained between opposite walls, thereby generating two lumens. Structural analysis of thestability of thin elastic rings and their postbuckling shape (Flaherty et al., 1972), as well as experi-mental studies (Thiriet et al., 2001) revealed the different complex modes of collapsed cross sections.In order to facilitate at least a one-dimensional fluid flow analysis, it is useful to represent themechanical characteristics of the vein wall by a “tube law” relationship that locally correlates betweenthe transmural pressure and the vein cross-sectional area.

Venous flow is a complex interaction between the compliant structures (veins and surroundingtissues) and the flow of blood. Since venous blood pressure is low, transmural pressure can becomenegative, thereby resulting in blood flow through a partially collapsed tube. Early studies with a thin-walled elastic tube revealed the relevant experimental evidence (Conrad, 1969). The steady flow rate(Q) through a given length of a uniform collapsible tube depends on two pressure differences

3.3.7 Flow in Bifurcating and Branching Systems

3.4.1 Vein Compliance

3.4.2 Flow in Collapsible Tubes

3.4 BLOOD FLOW IN VEINS





selected from among the pressures immediatelyupstream (P1), immediately downstream (P2),and external (Pe) to the collapsible segment(Fig. 3.11). Thus, the pressure-flow relation-ships in collapsible tubes are more complexthan those of rigid tubes, where Q is related to afixed pressure gradient, and may attain differ-ent shapes, depending on which of the pres-sures (e.g., P1, P2, Pe) are held fixed and whichare varied. In addition, one should also con-sider the facts that real veins may be neitheruniform nor straight, and that the external pres-sure is not necessarily uniform along the tube.

The one-dimensional theory for steady in-compressible fluid flow in collapsible tubes(when P - Pe < 0) was outlined by Shapiro(1977) in a format analogous to that for gasdynamics. The governing equations for thefluid are that for conservation of mass,

(3.39)

and that for conservation of momentum,

(3.40)

where u = velocity P = pressure in the flowing fluid ρ = mass density of the fluid A = tube cross-sectional area t = time z = longitudinal distance

FIGURE 3.10 Relationship between transmural pressure, P - Pe,and normalized cross-sectional area, (A - A0)/A0, of a long seg-ment of inferior vena cava of a dog. The solid line is a com-puter solution for a latex tube. [From Moreno et al. (1970) bypermission.]

FIGURE 3.11 Sketch of a typical experimental system for investigation of liquid flow in acollapsible tube. See text for notation.





The governing equation for the tube deformation may be given by the tube law, which is also anequation of state that relates the transmural pressure to the local cross-sectional area,

(3.41)

where A0 is the unstressed circular cross section and Kp is the wall stiffness coefficient. Solution ofthese governing equations for given boundary conditions provides the one-dimensional flow patternof the coupled fluid-structure problem of fluid flow through a collapsible elastic tube.

Shapiro (1977) defined the speed index, S = u/c, similar to the Mach number in gas dynamics, anddemonstrated different cases of subcritical (S < 1) and supercritical (S > 1) flows. It has been shownexperimentally in simple experiments with compliant tubes that gradual reduction of the downstreampressure progressively increases the flow rate until a maximal value is reached (Holt, 1969; Conrad,1969). The one-dimensional theory demonstrates that for a given tube (specific geometry and wallproperties) and boundary conditions, the maximal steady flow that can be conveyed in a collapsibletube is attained for S = 1 (e.g., when u = c) at some position along the tube (Dawson & Elliott, 1977;Shapiro, 1977; Elad et al., 1989). In this case, the flow is said to be “choked” and further reductionin downstream pressure does not affect the flow upstream of the flow-limiting site. Much of itscomplexity, however, is still unresolved either experimentally or theoretically (Kamm et al., 1982;Kamm and Pedley, 1989; Elad et al., 1992).

The concept of a closed circuit for the circulation was established by Harvey (1578–1657). Theexperiments of Hagen (1839) and Poiseuille (1840) were performed in an attempt to elucidate theflow resistance of the human microcirculation. During the past century, major strides have been madein understanding the detailed fluid mechanics of the microcirculation and in depicting a concretepicture of the flow in capillaries and other small vessels.

We include in the term “microcirculation” those vessels with lumens (internal diameters) that are somemodest multiple—say 1 to 10—of the major diameter of the unstressed RBC. This definition includesprimarily the arterioles, the capillaries, and the postcapillary venules. The capillaries are of particularinterest because they are generally from 6 to 10 µm in diameter, i.e., about the same size as the RBC.In the larger vessels, RBC may tumble and interact with one another and move from streamline tostreamline as they course down the vessel. In contrast, in the microcirculation the RBC must travel insingle file through true capillaries (Berman and Fuhro, 1969; Berman et al., 1982). Clearly, anyattempt to adequately describe the behavior of capillary flow must recognize the particulate nature ofthe blood.

The tortuosity and intermittency of capillary flow argue strongly that the case for an analytic descrip-tion is lost from the outset. To disprove this, we must return to the Navier-Stokes equations for amoment and compare the various acceleration and force terms, which apply in the microcirculation.The momentum equation, which is Newton’s second law for a fluid, can be written

(3.42)

3.5 BLOOD FLOW IN THE MICROCIRCULATION

3.5.1 The Microvascular Bed

3.5.2 Capillary Blood Flow





Since we are analyzing capillaries in which the RBC are considered solid bodies traveling in a tubeand surrounded by a waterlike fluid (plasma), a good representation of the viscous shear forces actingin the fluid phase is the newtonian flow,

(3.43)

We now examine the four terms in the momentum equation from the vantage point of an observersitting on the erythrocytes. It is an observable fact that most frequently the fluid in the capillarymoves at least 10 to 20 vessel diameters before flow ceases, so that a characteristic time for theunsteady term (A) is, say, 10 D/U. The distance over which the velocity varies by U is, typically, D.(In the gap between the RBC and the wall, this distance is, of course, smaller, but the sense of ourargument is not changed.)

Dividing both sides of Eq. (3.42) by ρ, we have the following order-of-magnitude comparisonsbetween the terms:

The term UD/ν is the well-known Reynolds number. Typical values for human capillaries are U ≈ 500µm/s, D ≈ 7 µm, ν ≈ 1.5 × 10-2 cm2/s, so that the Reynolds number is about 2 × 10-3. Clearly, theunsteady (A) and convective acceleration (B) terms are negligible compared to the viscous forces (Le-Cong and Zweifach, 1979; Einav and Berman, 1988).

This result is most welcome, because it allows us to neglect the acceleration of the fluid as it passesaround and between the RBCs, and to establish a continuous balance between the local net pressureforce acting on an element of fluid and the viscous stresses acting on the same fluid element. Theequation to be solved is therefore

(3.45)

subject to the condition that the fluid velocity is zero at the RBC surface, which is our fixed frame ofreference, and U at the capillary wall. We must also place boundary conditions on both the pressureand velocity at the tube ends, and specify the actual shape and distribution of the RBC. This requiressome drastic simplifications if we wish to obtain quantitative results, so we assume that all the RBChave a uniform shape (sphere, disk, ellipse, pancake, etc.) and are spaced at regular intervals. Thenthe flow, and hence the pressure, will also be subject to the requirement of periodicity, and we canidealize the ends of the capillary as being substantially removed from the region being analyzed. Ifwe specify the relative velocity U between the capillary and the RBC, the total pressure drop acrossthe capillary can be computed.

For isolated and modestly spaced RBC, the fluid velocities in the vicinity of a red cell is schematicallyshown in Fig. 3.12. In the gap, the velocity varies from U to zero in a distance h, whereas in the“bolus” region between the RBC, the same variation is achieved over a distance of D/4. If h < D/4, asis often observed in vivo, then the viscous shear force is greatest in the gap region and tends to “pull”the RBC along in the direction of relative motion of the wall.

Counteracting this viscous force must be a net pressure, Pu – Pd, acting in a direction opposite tothe sense of the shear force. This balance of forces is the origin of the parachutelike shape shown inFig. 3.3 and frequently observed under a microscope.

3.5.3 Motion of a Single Cell

(3.44)





For h � D/4, we can approximate the net pressure,

(3.46)

where 2b is the axial extent of the region of the gap. Suppose we use Eq. (3.46) to estimate thepressure drop across a typical capillary. Taking h = 0.02D, b = 0.1D, U = 500 µm/s, D = 7 µm, andµ = 1.4 × 10-2 dyn · s/cm2, then Pu – Pd ≈ 40 dyn/cm2.

Under normal physiological conditions, systole and diastole occur in a definite coordination andconstitute the cardiac cycle. Each cycle is considered to start with the atrial systole. The contractionbegins a wave in that part of the right atrium where the orifices of the venae cavae are, and theninvolves both atria, which have a common musculature. With the cardiac rhythm of 75 contractionsper minute, an atrial (auricular) systole lasts 0.1 second. As it ends, the ventricle systole begins, theatria then being in a state of diastole, which lasts 0.7 second. The contraction of the two ventriclesoccurs simultaneously, and their systole persists for about 0.3 second. After that, ventricular diastolebegins and lasts about 0.5 second. One-tenth second before the end of the ventricular diastole, anew atrial systole occurs, and a new cycle of cardiac activity begins. The interconnection andsequence of the atrial and ventricular contractions depend upon where stimulation arises in the heartand how it spreads. Contraction of the ventricular myocardium ejects blood into the aorta andpulmonary arteries.

The heart valves are unidirectional valves, and in normal physiological conditions, blood flows inonly one direction in the heart cavities: from the atria into the ventricles, and from the ventricles intothe arterial system (Fig. 3.13). The ring-shaped muscle bundles of the atria, which surround theorifices, like a sphincter contract first during atrial systole, constricting these orifices so that bloodflows from the atria only in the directions of the ventricles, and does not return into the veins. As theventricles are relaxed during the atrial systole, and the pressure within them is lower than that in thecontracting atria, blood enters them from the atria.

The human heart contains four unidirectional valves that are anatomically grouped into two types: theatrioventricular valves and the semilunar valves. The tricuspid and mitral (bicuspid) valves belong to

3.6 BLOOD FLOW IN THE HEART

3.6.1 Flow in the Heart Ventricles

3.6.2 Flow through Heart Valves

FIGURE 3.12 Diagram of the fluid pressure and velocity near thered blood cell within a capillary.





the first type, whereas the pulmonic and aortic valves compose the second. Valves of the same typeare not only structurally similar, but also functionally alike. For this reason, conclusions derived fromstudies on aortic and mitral valves are generally applicable to pulmonic and tricuspid valves,respectively.

One-way passages of blood from the ventricles into the main arteries is due to the heart valves.Though relatively simple in structure, heart valves are in their healthy state remarkably efficient,opening smoothly with little resistance to flow and closing swiftly in response to a small pressuredifference with negligible regurgitation. Yet they are essentially passive devices that move in reactionto fluid-dynamical forces imposed upon them.

The motions of heart valves have drawn a considerable amount of attention from physicians andanatomists. One of the major reasons for this interest can be attributed to the frequent diagnosis ofvalvular incompetence in association with cardiopulmonary dysfunctionings. The first studyexploring the nature of valve functioning, by Henderson and Johnson (1912), was a series of simplein vitro experiments, which provided a fairly accurate description of the dynamics of valve closure.Then came the Bellhouse and Bellhouse experiments (1969) and Bellhouse and Talbot (1969)analytical solution, which showed that the forces responsible for valve closure, are directly related tothe stagnation pressure of the flow field behind the valve cusps.

Computational fluid dynamics models were developed over the years that include the effects ofleaflet motion and its interaction with the flowing blood (Bellhouse et al., 1973; Mazumdar, 1992).Several finite-element structural models for heart valves were also developed in which issues such asmaterial and geometric nonlinearities, leaflet structural dynamics, stent deformation, and leafletcoaptation for closed valve configurations were effectively dealt with (Bluestein and Einav, 1993;1994). More recently, fluid-structure interaction models, based on the immersed boundary technique,

FIGURE 3.13 Structure of the heart and course of blood flow throughthe heart chambers. [From Guyton and Hall (1996) by permission.]





were used for describing valvular function (Sauob et al., 1999). In these models, strong coupledfluid-structure dynamic simulations were obtained. These models allowed for the inclusion of bend-ing stresses, contact between adjacent leaflets when they coapted, and transient three-dimensionalblood flow through the valve.

FIGURE 3.14 The arterial coronary circulation from anterior view. Arteries near theanterior surface are darker than those of the posterior surface seen through the heart.[From Thibodeau and Patton (1999) by permission.]

FIGURE 3.15 Blood flow through the coronary arteries: left (top) and right(bottom).





The supply of blood to the heart is provided by the coronary vessels (Fig. 3.14), and occurs mainlyduring the diastole, opposite to other blood vessels. Coronary flow is low and even reversing duringsystole, while venous outflow is high. In diastole, coronary flow is high while venous flow is low.Contraction of the myocardium during the period of ventricular tension compresses the small arterieslying within it to such an extent that blood flow in the coronaries is sharply reduced (Fig. 3.15).Some blood from the cardiac veins enters the coronary sinus, which empties into the right atrium(Kajiya et al., 1990; Hoffman et al., 1985). The sinus receives blood mainly from the veins of the leftventricle, around 75 to 90 percent. A large amount of blood flows from the myocardium of theinteratrial septum and from the right ventricle along numerous microvessels, and drains into the rightventricle (Krams et al., 1989). From 200 to 250 mL of blood flows through the coronaries of ahuman being per minute, which is about 4 to 6 percent of the minute volume of the heart. Duringphysical exertion, coronary flow may rise to 3 or 4 L per minute (Bellhouse et al., 1968).

The circulatory system is a complex system of branching compliant tubes, which adjusts itself accord-ing to complex controllers. Mechanical and electrical analog models that can be solved analyticallywere devised to study either single arteries or the whole circulatory subsystem of specific organs (e.g.,heart, kidney). Because of the complexity of the human system, there are a multitude of variables,which affect the functions, properties, and response of the circulatory system. Experimentally it isimpossible to include all the known variables in a single system. However, analog models can dealwith a group of variables at a time, and even make it possible to study the interaction between thevariables. For extreme circulatory and physiological situations, it is simply too dangerous to studythem experimentally, while the models are immune to studies.

The Windkessel (air-cell in German) was suggested by Otto Frank (1899) to represent the cardiovas-cular system (Fig. 3.16a). In this linear mechanical model, the aorta and large blood vessels are

3.6.3 Coronary Blood Flow

3.7 ANALOG MODELS OF BLOOD FLOW

3.7.1 Introduction

3.7.2. Mechanical Models (Windkessel)

FIGURE 3.16 (a) The Windkessel model of the aorta and peripheral circulation, (b) Electricalmodel of the Windkessel model.





represented by a linear compliant air-cell and the peripheral vessels are replaced by a rigid tube witha linear resistance (Dinnar, 1981; Fung, 1984). Accordingly,

(3.47)

where V, P, and C = volume, pressure, and compliance of the air-cell and are linearly related Qout = outflow of the air-cell R = linear resistance of the rigid tube Pvenous = 0

Conservation of mass requires that

(3.48)

where Qin is the inflow to the air-cell and represents the cardiac stroke. This governing equation canbe solved analytically for a set of given boundary conditions. The scheme of this analog is shown inFig. 3.16b.

A realistic model of blood flow in a compliant artery or vein requires a coupled solution of thegoverning equations for both the fluid and the elastic tube wall along with boundary conditions.These models are known as fluid-structure problems and can be solved by complex numericaltechniques. In order to investigate linear approximations of complex blood flow phenomena, electri-cal analogs, which result in analogous linear differential equations, were developed (Dinnar, 1981).In these models the vessel resistance to blood flow is represented by a resistor R, the tube complianceby a capacitor C, and blood inertia by an inductance L. Assuming that blood flow and pressure areanalogous to the electrical current and voltage, respectively, blood flow through a compliant vessel(e.g., artery or vein) may be represented by an electrical scheme (Fig. 3.17). Following Kirchhoff’slaw for currents, the governing differential equation becomes

(3.49)

For a system of vessels attached in series (or small units of a nonuniform vessel) a more complicatedscheme, composed of a series of compartments whose characteristics are represented by Ri, Ci, and Li

(Fig. 3.18), is used. For this case, Kirchhoff’s law applied to each loop will yield

3.7.3. Electrical Models

FIGURE 3.17 Electrical analog for blood flow in a compliant vessel.





(3.50)

(3.51)

(3.52)

The problem is reduced to a set of linear differential equations and can be solved for given boundaryconditions. The RCL characteristics of a given problem may be derived from the known physicalproperties of blood and blood vessels (Dinnar, 1981; Van der Twell, 1957). This compartmentalapproach allowed for computer simulations of complex arterial circuits with clinical applications(McMahon et al., 1971; Clark et al., 1980; Barnea et al., 1990; Olansen et al., 2000; Westerhof andStergiopulos, 2000; Ripplinger et al., 2001).

The authors are thankful to Dr. Uri Zaretsky for his technical assistance.

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ACKNOWLEDGMENT

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FIGURE 3.18 Lumped-parameter electrical model of a multicompartment arterial segment.





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chapter 3 physical and flow properties … 3 physical and flow properties of blood david elad and...

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