Properties of the Vasculature

Velocity of the Bloodstream

Velocity, as relates to fluid movement, is the distance that a particle of fluid travels with regard to time, and it is expressed in units of distance per unit time (e.g., centimeters per second). Flow, in contrast, is the rate of displacement of a volume of fluid, and it is expressed in units of volume per unit time (e.g., cubic centimeters per second). In a rigid tube, velocity (v) and flow (Q) are related to one another by the cross-sectional area (A) of the tube:

The interrelationships among velocity, flow, and area are shown in Fig. 17.1 . Because conservation of mass requires that the fluid flowing through a rigid tube be constant, the velocity of the fluid varies inversely with the cross-sectional area. Thus fluid flow velocity is greatest in the section of the tube with the smallest cross-sectional area and slowest in the section of the tube with the greatest cross-sectional area.

As shown in Fig. 15.3 , velocity decreases progressively as blood traverses the arterial system. In the capillaries, velocity decreases to a minimal value. As the blood then passes centrally through the venous system toward the heart, velocity progressively increases again. The relative velocities in the various components of the circulatory system are related only to the respective cross-sectional areas.

Relationship Between Velocity and Pressure

The total energy in a hydraulic system consists of three components: pressure, gravity, and velocity. The velocity of blood flow can have an important effect on the pressure within the tube. Consider the effect of velocity on pressure in a tube with different cross-sectional areas ( Fig. 17.2 ). In this system, the total energy remains constant. The total pressure within the tube equals the lateral (static) pressure plus the dynamic pressure. The gravitational component can be neglected because the tube is horizontal. The total pressures in segments A, B, and C are equal, provided that the energy loss from viscosity is negligible (i.e., this fluid is an “ideal fluid”). The effect of velocity on the dynamic component (P _dyn ) can be estimated as follows:

where ρ is the density of the fluid (grams per cubic centimeters) and v is velocity (centimeters per second). Assume that the fluid has a density of 1 g/cm ³ . In section A in Fig. 17.2 , the lateral pressure is 100 mm Hg; note that 1 mm Hg equals 1330 dynes/cm ² . According to Eq. 17.2 , P _dyn = 5000 dynes/cm ² , or 3.8 mm Hg. In the narrow section B of the tube, where the velocity is twice as high, P _dyn = 20,000 dynes/cm ² , or 15 mm Hg. Thus the lateral pressure in section B is 15 mm Hg lower than the total pressure, whereas the lateral pressures in sections A and C are only 3.8 mm Hg lower. In most arterial locations, the dynamic component is a negligible fraction of the total pressure. However, at sites of an arterial constriction or obstruction, the high flow velocity is associated with large kinetic energy, and the dynamic pressure component may therefore increase significantly. Hence, the pressure would be reduced, and perfusion of distal segments would be correspondingly decreased. This example helps explain how pressure changes in a vessel that is narrowed by atherosclerosis or spasm of the blood vessel wall: that is, in narrowed sections of a tube, the dynamic component increases significantly because the flow velocity is associated with large kinetic energy.

Relationship Between Pressure and Flow

The most fundamental law that governs the flow of fluids through cylindrical tubes was derived empirically by the French physiologist Jean Léonard Marie Poiseuille in the 1840s. He was interested primarily in the physical determinants of blood flow, but he replaced blood with simpler liquids in his measurements of flow through glass capillary tubes. His work was so precise and important that his observations have been designated Poiseuille’s law.

Poiseuille’s Law

Poiseuille’s law applies to the steady (i.e., nonpulsatile) laminar flow of Newtonian fluids through rigid cylindrical tubes. A Newtonian fluid is one whose viscosity remains constant, and laminar flow is the type of motion in which the fluid moves as a series of individual layers, with each layer moving at a velocity different from that of its neighboring layers ( Fig. 17.3 A ). In the case of laminar flow through a tube, the fluid consists of a series of infinitesimally thin concentric tubes sliding past one another, of which the central tube has the highest velocity. The velocities of the concentric laminae decrease parabolically toward the vessel wall. Despite the differences between the vascular system (i.e., flow is pulsatile, the vessels are not rigid cylinders, and blood is not a Newtonian fluid), Poiseuille’s law does provide valuable insight into the determinants of blood flow through the vascular system. In certain unusual situations, however, flow can become turbulent (see Fig. 17.3 B ) rather than laminar. Under these conditions, vortices (swirls) are present, and the distribution of flow velocities is chaotic. This condition is described in more detail later in this chapter.

Poiseuille’s law describes the laminar flow of fluids through cylindrical tubes in terms of pressure, the dimensions of the tube, and the viscosity of liquid:

$\text{Q}=\pi \left({\text{P}}_{\text{i}}-{\text{P}}_{\text{o}}\right){\text{r}}^4/8\eta \text{l}$

where

• Q = flow

• P _i − P _o = pressure gradient from the inlet (i) of the tube to the outlet (o)

• r = radius of the tube

• η = viscosity of the fluid

• l = length of the tube

As is clear from the equation, flow through the tube increases as the pressure gradient is increased, and it decreases as either the viscosity of the fluid or the length of the tube increases. The radius of the tube is a critical factor in determining flow because it is raised to the fourth power.

Resistance to Flow

In electrical theory, Ohm’s law is that the resistance (R) equals the ratio of voltage drop (E) to current flow (I).

$\text{R}=\text{E}/\text{I}$

Similarly, in fluid mechanics, hydraulic resistance (R) may be defined as the ratio of the pressure drop (P _i − P _o ,) to flow (Q):

$\text{R}=\left({\text{P}}_{\text{i}}-{\text{P}}_{\text{o}}\right)/\text{Q}$

For the steady, laminar flow of a Newtonian fluid through a cylindrical tube, the physical components of hydraulic resistance may be appreciated by the rearranging of Poiseuille’s law to yield the hydraulic resistance equation:

$\text{R}=\left({\text{P}}_{\text{i}}-{\text{P}}_{\text{o}}\right)/\text{Q}=8\eta \text{l}/\pi {\text{r}}^4$

Thus when Poiseuille’s law applies, the resistance to flow depends on only the dimensions of the tube and the characteristics of the fluid.

The principal determinant of resistance to blood flow through any vessel is the caliber of the vessel because resistance varies inversely as the fourth power of the radius of the tube. In Fig. 17.4 , the resistance to flow through small blood vessels is measured, and the resistance per unit length of vessel (R/l) is plotted against the vessel diameter. As shown, resistance is highest in the capillaries (diameter of 7 µm), and it diminishes as the vessels increase in diameter on the arterial and venous sides of the capillaries. Values of R/l are virtually inversely proportional to the fourth power of the diameter (or radius) of the larger vessels on both sides of the capillaries.

Changes in vascular resistance occur when the caliber of vessels changes. The most important factor that leads to a change in vessel caliber is contraction of the circular smooth muscle cells in the vessel wall. Changes in internal pressure also alter the caliber of blood vessels and therefore alter the resistance to blood flow through these vessels. Blood vessels are elastic tubes. Hence, the greater the transmural pressure (i.e., the difference between internal and external pressure) across the wall of a vessel, the greater the caliber of the vessel and the less its hydraulic resistance.

It is apparent from Fig. 15.3 that the greatest drop in pressure occurs in the very small arteries and arterioles. However, capillaries, which have a mean diameter of approximately 7 µm, have the greatest resistance to blood flow. Nevertheless, of all the different varieties of blood vessels that lie in series with one another (as in Fig. 15.3), the arterioles, not the capillaries, have the greatest resistance. This seeming paradox is related to the relative numbers of parallel capillaries and parallel arterioles: There are far more capillaries than arterioles in the systemic circulation, and total resistance across the many capillaries arranged in parallel is much less than total resistance across the fewer arterioles arranged in parallel. In addition, arterioles have a thick coat of circularly arranged smooth muscle fibers that can vary the lumen radius. Even small changes in radius alter resistance greatly, as can be seen from the hydraulic resistance equation ( Eq. 17.6 ), wherein R varies inversely with r ⁴ .

Resistances in Series and in Parallel

In the cardiovascular system, the various types of vessels listed along the horizontal axis in Fig. 15.3 lie in series with one another. The individual members of each category of vessels are ordinarily arranged in parallel with one another (see Fig. 15.1). Thus capillaries are in most instances parallel elements throughout the body, except in the renal vasculature (in which the peritubular capillaries are in series with the glomerular capillaries) and the splanchnic vasculature (in which the intestinal and hepatic capillaries are aligned in series with each other). The total hydraulic resistance of components arranged in series or in parallel can be derived in the same manner as those for analogous combinations of electrical resistance.

Resistance of Vessels in Series

In the system depicted in Fig. 17.5 , three hydraulic resistances, R ₁ , R ₂ , and R ₃ , are arranged in series. The pressure drop across the entire system (i.e., the difference between inflow pressure [P _i ] and outflow pressure [P _o ]) consists of the sum of the pressure drops across each of the individual resistances (equation [a] in Fig. 17.5 ). In the steady state, the flow (Q) through any given cross-section must equal the flow through any other cross-section. When each component in equation (a) is divided by Q (equation [b] in Fig. 17.5 ), it is evident from the definition of resistance ( Eq. 17.5 ) that for resistances in series, the total resistance (R _t ) of the entire system equals the sum of the individual resistances; that is,

${\text{R}}_{\text{t}}={\text{R}}_{\text{l}}+{\text{R}}_2+{\text{R}}_3$

Resistance of Vessels in Parallel

For resistances in parallel, as illustrated in Fig. 17.6 , inflow and outflow pressure is the same for all tubes. In steady state, the total flow (Q _t ) through the system equals the sum of the flows through the individual parallel elements (equation [a] in Fig. 17.5 ). Because the pressure gradient (P _i − P _o ) is identical for all parallel elements, each term in equation (a) may be divided by that pressure gradient to yield equation (b). From the definition of resistance, equation (c) in Fig. 17.5 may be derived. According to this equation, for resistances in parallel, the reciprocal of the total resistance (R _t ) equals the sum of the reciprocals of the individual resistances; that is,

$1/{\text{R}}_{\text{t}}=\left(1/{\text{R}}_1\right)+(1/{\text{R}}_2)+(1/{\text{R}}_3)$

In a few simple illustrations, some of the fundamental properties of parallel hydraulic systems become apparent. For example, if the resistances of the three parallel elements in Fig. 17.6 were all equal, then

${\text{R}}_1={\text{R}}_2={\text{R}}_3$

Therefore, from Eq. 17.8 ,

$1/{\text{R}}_{\text{t}}=3/{\text{R}}_1$

When the reciprocals of these terms are equated,

${\text{R}}_{\text{t}}={\text{R}}_1/3$

Thus the total resistance is less than the individual resistances. For any parallel arrangement, the total resistance must be less than that of any individual component. For example, consider a system in which a tube with very high resistance is added in parallel to a low-resistance tube. The total resistance of the system must be less than that of the low-resistance component by itself because the high-resistance component affords an additional pathway, or conductance, for flow of fluid.

Consider the physiological relationship between the total peripheral resistance (TPR) of the entire systemic vascular bed and the resistance of one of its components, such as the renal vasculature. TPR is the ratio of the arteriovenous (AV) pressure difference (arterial pressure [P _a ] − venous pressure [P _v ]) to the flow through the entire systemic vascular bed (i.e., the cardiac output [Q _t ]). For example, the renal vascular resistance (R _r ) would be the ratio of the same AV pressure difference (P _a − P _v ) to renal blood flow (Q _r ).

In an individual with an P _a of 100 mm Hg, a peripheral P _v of 0 mm Hg, and a cardiac output of 5000 mL/minute, TPR is 0.02 mm Hg/mL/minute, or 0.02 peripheral resistance units (PRUs). Normally, the rate of blood flow through one kidney would be approximately 600 mL/minute. Renal resistance would therefore be 100 mm Hg ÷ 600 mL/minute, or 0.17 PRUs, which is 8.5 times greater than the TPR. In an organ such as the kidney, which weighs only approximately 1% as much as the whole body, the vascular resistance is much greater than that of the entire systemic circulation. Hence, it is not surprising that the resistance to flow would be greater for a component organ, such as the kidney, than for the entire systemic circulation because the systemic circulation has not only one kidney but also many more alternative pathways for blood to flow.

Rheologic Properties of Blood

The viscosity of a given Newtonian fluid at a specified temperature stays constant over a wide range of tube dimensions and flows. However, for a non-Newtonian fluid such as blood, viscosity may vary considerably as a function of tube dimensions and flows. Therefore, the term viscosity does not have a unique meaning for blood. The term apparent viscosity is frequently used for the derived value of blood viscosity obtained under the particular conditions of measurement.

Rheologically, blood is a suspension of formed elements, principally erythrocytes, in a relatively homogeneous liquid, the blood plasma. Because blood is a suspension, the apparent viscosity of blood varies as a function of the hematocrit (ratio of the volume of red blood cells to the volume of whole blood). The viscosity of plasma is 1.2 to 1.3 times that of water. The upper curve in Fig. 17.7 shows that the apparent viscosity of blood with a normal hematocrit ratio of 45% is 2.4 times that of plasma. ^a

a Fig. 17.7 also illustrates that the apparent viscosity of blood, when measured in living tissues, is considerably less than the apparent viscosity of the same blood measured in a conventional capillary tube viscometer.

In severe anemia, blood viscosity is low. As the hematocrit increases, the slope of the curve increases progressively; it is especially steep at the upper range of erythrocyte concentrations (see Fig. 17.7 ).

For any given hematocrit, the apparent viscosity of blood depends on the dimensions of the tube used in estimating the viscosity. Fig. 17.8 demonstrates that the apparent viscosity of blood diminishes progressively as tube diameter decreases to less than approximately 0.3 mm. The diameters of the blood vessels with the highest resistance, the arterioles, are considerably less than this critical value. This phenomenon therefore reduces the resistance to flow in blood vessels that possess the greatest resistance. The influence of tube diameter on apparent viscosity is explained in part by the actual change in blood composition as it flows through small tubes. The composition of blood changes because the red blood cells tend to accumulate in the faster axial stream, whereas plasma tends to flow in the slower marginal layers. Because the axial portions of the bloodstream contain a greater proportion of red cells and this axial portion moves at greater velocity, the red blood cells tend to traverse the tube in less time than plasma does. Furthermore, the hematocrit of the blood contained in small blood vessels is lower than that in blood in large arteries or veins.

The physical forces responsible for the drift of erythrocytes toward the axial stream and away from the vessel walls when blood is flowing at normal rates are not fully understood. One factor is the great flexibility of red blood cells. At low flow rates, like those in the microcirculation, rigid particles do not migrate toward the central axis of a tube, whereas flexible particles do. The concentration of flexible particles near the tube’s central axis is enhanced by an increase in the shear rate.

The apparent viscosity of blood diminishes as the shear rate is increased ( Fig. 17.9 ), a phenomenon called shear thinning. The greater the amount of flow, the greater the rate that one lamina of fluid shears against an adjacent lamina. The greater tendency for erythrocytes to accumulate in the axial laminae at higher flow rates is partly responsible for this non-Newtonian behavior. However, a more important factor is that at very slow flow rates, the suspended cells tend to form aggregates; such aggregation increases blood viscosity. As flow is increased, this aggregation decreases, and so does the apparent viscosity of blood (see Fig. 17.9 ).

The tendency for erythrocytes to aggregate at low flow rates depends on the concentration of the larger protein molecules in plasma, especially fibrinogen. For this reason, changes in blood viscosity with flow rate are much more pronounced when the concentration of fibrinogen is high. In addition, at low flow rates, leukocytes tend to adhere to the endothelial cells of the microvessels and thereby increase the apparent viscosity of the blood.

The deformability of erythrocytes is also a factor in shear thinning, especially when the hematocrit is high. The mean diameter of human red blood cells is approximately 7 µm, but they are able to pass through openings with a diameter of only 3 µm. As blood with densely packed erythrocytes flows at progressively greater rates, the erythrocytes become more and more deformed. Such deformation diminishes the apparent viscosity of blood. The flexibility of human erythrocytes is enhanced as the concentration of fibrinogen in plasma increases ( Fig. 17.10 ). If the red blood cells become hardened, as they are in certain spherocytic anemias, shear thinning may diminish.

Arterial Elasticity

The systemic and pulmonary arterial systems distribute blood to the capillary beds throughout the body. The arterioles are high-resistance vessels of this system that regulate the distribution of flow to the various capillary beds. The aorta, the pulmonary artery, and their major branches have a large amount of elastin in their walls, which makes these vessels highly distensible (i.e., compliant). This distensibility serves to dampen the pulsatile nature of blood flow that results as the heart pumps blood intermittently. When blood is ejected from the ventricles during systole, these vessels distend, and during diastole, they recoil and propel the blood forward ( Fig. 17.11 ). Thus the intermittent output of the heart is converted to a steady flow through the capillaries.

The elastic nature of the large arteries also reduces the work of the heart. If these arteries were rigid rather than compliant, the pressure would rise dramatically during systole. This increased pressure would require the ventricles to pump against a large load (i.e., afterload) and thus increase the work of the heart. Instead, as blood is ejected into these vessels, they distend, and the resultant increase in systolic pressure, and thus the work of the heart, is reduced.

IN THE CLINIC

As people age, the elastin content of the large arteries is reduced and replaced by collagen. This reduces arterial compliance ( Fig. 17.12 ). Thus with age, systolic pressure increases, as does the difference between systolic and diastolic blood pressure, called the pulse pressure (described in the next section).

Determinants of Arterial Blood Pressure

Arterial blood pressure is routinely measured in patients, and it provides a useful estimate of their cardiovascular status. Arterial pressure can be defined as (P_ _a ) , which is the pressure averaged over time, and as systolic (maximal) and diastolic (minimal) arterial pressure within the cardiac cycle ( Fig. 17.13 ). The difference between systolic and diastolic pressure is termed pulse pressure.

The determinants of arterial blood pressure are arbitrarily divided into “physical” and “physiological” factors. The two physical factors, or fluid mechanical characteristics, are fluid volume (i.e., blood volume) within the arterial system and the static elastic characteristics (compliance) of the system. The physiological factors are cardiac output (which equals heart rate × stroke volume) and peripheral resistance.

Mean Arterial Pressure

To estimate ${\overline{\text{P}}}_{\text{a}}$ from an arterial blood pressure tracing, the area under the pressure curve is divided by the time interval involved (see Fig. 17.13 ). Alternatively, ${\overline{\text{P}}}_{\text{a}}$ can be approximated from the measured values of systolic pressure (P _s ) and diastolic pressure (P _d ) by means of the following formula:

${\overline{\text{P}}}_{\text{a}}={\text{P}}_{\text{d}}+({\text{P}}_{\text{s}}-{\text{P}}_{\text{d}}/3)$

Consider that ${\overline{\text{P}}}_{\text{a}}$ depends on only two physical factors: mean blood volume in the arterial system and arterial compliance ( Fig. 17.14 ). Arterial volume (V _a ), in turn, depends on the rate of inflow (Q _h ) into the arteries from the heart (cardiac output) and on the rate of outflow (Q _r ) from the arteries through the resistance vessels (peripheral runoff). These relationships are expressed mathematically as

${\text{dV}}_{\text{a}}/\text{dt}={\text{Q}}_{\text{h}}-{\text{Q}}_{\text{r}}$

where dV _a /dt is the change in arterial blood volume per unit of time. If Q _h exceeds Q _r , arterial volume increases, the arterial walls are stretched further, and pressure rises. The converse happens when Q _r exceeds Q _h . When Q _h equals Q _r , P _a remains constant. Thus increases in cardiac output raise ${\overline{\text{P}}}_{\text{a}}$ , as do increases in peripheral resistance. Conversely, decreases in cardiac output or peripheral resistance decrease ${\overline{\text{P}}}_{\text{a}}$ .

Arterial Pulse Pressure

Arterial pulse pressure is systolic pressure minus diastolic pressure. It is principally a function of just one physiological factor, stroke volume, which determines the change in arterial blood volume (a physical factor) during ventricular systole. This physical factor, in addition to a second physical factor (arterial compliance), determines the arterial pulse pressure (see Fig. 17.14 ).

Stroke Volume

As described previously, ${\overline{\text{P}}}_{\text{a}}$ depends on cardiac output and peripheral resistance. During the rapid ejection phase of systole, the volume of blood introduced into the arterial system exceeds the volume that exits the system through the arterioles. Arterial pressure and volume therefore peak; the peak arterial pressure is systolic pressure. During the remainder of the cardiac cycle (i.e., ventricular diastole), cardiac ejection is zero, and peripheral runoff now greatly exceeds cardiac ejection. The resultant decrement in arterial blood volume thus causes pressure to fall to a minimum, which is diastolic pressure. Fig. 17.15 illustrates the effect of stroke volume on pulse pressure when arterial compliance is constant.

Arterial Compliance

Arterial compliance (C _a ), the ratio of blood volume to mean blood pressure (see Eq. 19.1 ), also affects pulse pressure. This relationship is illustrated in Fig. 17.16 . When cardiac output and TPR are constant, a decrease in arterial compliance results in an increase in pulse pressure. Diminished arterial compliance also imposes a greater workload on the left ventricle (i.e., increased afterload), even if stroke volume, TPR, and ${\overline{\text{P}}}_{\text{a}}$ are equal in the two individuals.

Total Peripheral Resistance and Arterial Diastolic Pressure

As previously discussed, if the heart rate and stroke volume remain constant, an increase in TPR causes ${\overline{\text{P}}}_{\text{a}}$ to increase. When arterial compliance is constant, an increase in TPR leads to proportional increases in systolic and diastolic pressure so that the pulse pressure is unchanged ( Fig. 17.17 A ). However, arterial compliance is not linear. As ${\overline{\text{P}}}_{\text{a}}$ increases and the artery is stressed, compliance decreases (see Fig. 17.17 B ). Because of the decrease in arterial compliance with increased P _a , pulse pressure increases when P _a is elevated.

Effect of Arterial Compliance on Myocardial Energy Consumption

The increased cardiac energy requirement imposed by a rigid arterial system is illustrated in Fig. 17.18 . In the data depicted in Fig. 17.18 , the cardiac output from the left ventricle either was allowed to flow through the natural route (the aorta) or was directed through a stiff plastic tube to the peripheral arteries. In this experiment, the TPR values were virtually identical, regardless of which pathway was selected. The results showed that for any given stroke volume, myocardial oxygen consumption was substantially greater when the blood was diverted through the plastic tubing than when it flowed through the aorta. The increased oxygen consumption indicates that the left ventricle has to expend significantly more energy to pump blood through a less compliant conduit than through a more compliant conduit.

IN THE CLINIC

Arterial pulse pressure provides valuable information about a person’s stroke volume, provided that arterial compliance is essentially normal. Patients who have severe congestive heart failure or who have suffered a severe hemorrhage are likely to have a very low arterial pulse pressure because their stroke volumes are abnormally small. Conversely, individuals with large stroke volumes, as in aortic valve regurgitation, are likely to have an increased arterial pulse pressure. Similarly, well-trained athletes at rest tend to have large stroke volumes because their heart rates are usually low. The prolonged ventricular filling times in these individuals induce the ventricles to pump a large stroke volume, and hence their pulse pressure is large.

Peripheral Arterial Pressure Curves

The radial stretch of the ascending aorta brought about by left ventricular ejection initiates a pressure wave that is propagated down the aorta and its branches. The pressure wave travels much faster (≈4–12 m/second) than the blood itself does. This pressure wave is the “pulse” that can be detected through palpation of a peripheral artery.

IN THE CLINIC

In chronic hypertension, a condition characterized by a persistent elevation in TPR, the arterial pressure-volume curve resembles that shown in Fig. 17.17 B. Because arteries become substantially less compliant when P _a rises, an increase in TPR causes systolic pressure to be more elevated than diastolic pressure. Diastolic pressure is elevated in such individuals but ordinarily not more than 10 to 40 mm Hg above the average normal level of 80 mm Hg. Not uncommonly, however, systolic pressure is elevated by 50 to 100 mm Hg above the average normal level of 120 mm Hg.

The velocity of the pressure wave varies inversely with arterial compliance. In general, transmission velocity increases with age, which confirms the observation that the arteries become less compliant with advancing age. Velocity also increases progressively as the pulse wave travels from the ascending aorta toward the periphery. This increase in velocity reflects the decrease in vascular compliance in the more distal portions than in the more proximal portions of the arterial system.

The P _a contour becomes distorted as the wave is transmitted down the arterial system. This distortion in the pressure wave contour of the human arterial tree is demonstrated as a function of age and of recording site in Fig. 17.19 . Damping of the high-frequency components of the arterial pulse is caused largely by the viscoelastic properties of the arterial walls. The pulse pressure wave travels more rapidly in older people than in the younger people, as a consequence of reduced compliance. Several factors—including wave reflection and resonance, vascular tapering, and pressure-induced changes in transmission velocity—contribute to peaking of the P _a wave.

Blood Pressure Measurement in Humans

Most commonly, blood pressure is estimated indirectly by means of a sphygmomanometer. In hospital intensive care units, needles or catheters may be introduced into the peripheral arteries of patients to measure arterial blood pressure directly by means of strain gauges. When blood pressure readings are taken from the arm, systolic pressure may be estimated by palpation of the radial artery at the wrist (palpatory method). While pressure in the cuff exceeds the systolic level, no pulse is perceived. As pressure falls just below the systolic level ( Fig. 17.20 A ) , a spurt of blood passes through the brachial artery under the cuff during the peak of systole, and a slight pulse is felt at the wrist.

The auscultatory method is a more sensitive and therefore more precise technique for measuring systolic pressure, and it also enables diastolic pressure to be estimated. The practitioner listens with a stethoscope applied to the skin of the antecubital space over the brachial artery. While the pressure in the cuff exceeds systolic pressure, the brachial artery is occluded, and no sounds are heard (see Fig. 17.20 B ). When the inflation pressure falls just below the systolic level (120 mm Hg in Fig. 17.20 A ), a small spurt of blood escapes the occluding pressure of the cuff, and slight tapping sounds (called Korotkoff sounds ) are heard with each heartbeat. The pressure at which the first sound is detected represents systolic pressure. It usually corresponds closely to the directly measured systolic pressure. As the inflation pressure of the cuff continues to fall, more blood escapes under the cuff per beat and the sounds become louder. When the inflation pressure approaches the diastolic level, the Korotkoff sounds become muffled. When the inflation pressure falls just below the diastolic level (80 mm Hg in Fig. 17.20 A ), the sounds disappear; the pressure reading at this point indicates diastolic pressure. The origin of the Korotkoff sounds is related to the discontinuous spurts of blood that pass under the cuff and meet a static column of blood beyond the cuff; the impact and turbulence generate audible vibrations. Once the inflation pressure is less than diastolic pressure, flow is continuous in the brachial artery, and sounds are no longer heard (see Fig. 17.20 C ).