Medical Physiology A Cellular and Molecular Approach, Updated 2nd Ed.


Emile L. Boulpaep


The sequence of mechanical and electrical events that repeats with every heartbeat is called the cardiac cycle. The duration of the cardiac cycle is the reciprocal of heart rate:


For example, for a heart rate of 75 beats/min, the cardiac cycle lasts 0.8 s or 800 ms.

The closing and opening of the cardiac valves define four phases of the cardiac cycle

The cardiac pump is of the two-stroke variety. Like a pump with a reciprocating piston, the heart alternates between a filling phase and an emptying phase. Under normal circumstances, the electrical pacemaker in the sinoatrial node (see Chapter 21) determines the duration of the cardiac cycle, and the electrical properties of the cardiac conduction system and cardiac myocytes determine the relative duration of contraction and relaxation. As long as the heart rate remains unchanged, this pattern remains steady.

The cardiac atria are small chambers. The right atrium receives deoxygenated systemic venous return from the inferior and superior vena cava. The left atrium receives oxygenated blood from the lungs through the pulmonary circulation. Both atria operate as passive reservoirs more than as mechanical pumps. However, they do contract, and this contraction does enhance ventricular filling and cardiac output to a small degree (see the box on the importance of atrial contractions).

The inlet valves of the ventricles are called the AV (atrioventricular) valves. They permit blood to flow in one direction only, from the atria to the ventricles. The valve located between the right atrium and the right ventricle is the tricuspid valve because it has three flaps, or cusps. The valve located between the left atrium and the left ventricle is the mitral valve because it has only two cusps, which resemble a bishop’s miter.

The outlet valves of the ventricles are called semilunar valves. They also allow blood to flow in just a single direction, from each ventricle into a large outflow tract vessel. Both the pulmonary valve, located between the right ventricle and pulmonary artery, and the aortic valve, located between the left ventricle and aorta, have three cusps.

Cardiac valves open passively when upstream pressure exceeds downstream pressure. They close passively when downstream pressure exceeds upstream pressure. The movement of the valve leaflets can be detected by echocardiography (see Chapter 17); their closure makes heart sounds that can be heard with a stethoscope. The stethoscope can also detect leaks in the valves that permit jets of blood to flow backward across the valvular orifice (i.e., regurgitation) as well as stenotic lesions that narrow the valve opening, forcing the blood to pass through a narrower space (i.e., stenosis). During certain parts of the cardiac cycle, blood passing through either regurgitant or stenotic lesions makes characteristic sounds that are called murmurs (see Chapter 17 for the box on this topic).

The cardiac cycle can be artificially divided into phases in any number of ways. However, from the point of view of the ventricles and the positions of their valves, we must consider a minimum of four distinct phases:

Inflow phase. The inlet valve is open and the outlet valve is closed.

Isovolumetric contraction. Both valves are closed, with no blood flow.

Outflow phase. The outlet valve is open and the inlet valve is closed.

Isovolumetric relaxation. Both valves are closed, with no blood flow.

Table 22-1 summarizes these four phases and the key events of the cardiac cycle. Note that the same events occur on the right side of the heart as on the left side.

Table 22-1 Events in the Cardiac Cycle


It is common to separate these phases into two parts. Systole includes phases 2 and 3, when the ventricles are contracting; diastole includes phases 4 and 1, when the ventricles are relaxing. For a heart rate of 75 (cycle duration = 800 ms), systole occupies ~300 ms, and diastole ~500 ms. With increasing heart rate—and thus decreasing cycle length—diastole shortens relatively more than systole does.

For convenience, the events in Table 22-1 start a short time after the beginning of diastole, with the opening of the AV valves and the start of ventricular filling (phase 1).

Changes in ventricular volume, pressure, and flow accompany the four phases of the cardiac cycle

Figure 22-1 illustrates the changes in pressure and volume that occur during the cardiac cycle. The four vertical lines indicate the timing of the four valvular events, which terminate each of the four phases defined previously:

AV valve closure terminates phase 1.

Semilunar valve opening terminates phase 2.

Semilunar valve closing terminates phase 3.

AV valve opening terminates phase 4.


Figure 22-1 Pressures and ventricular volumes during the cardiac cycle. A, Right side of the heart. B, Left side of the heart. The inset shows the placement of catheters used for pressure measurements in the right side of the heart.

The shapes of pressure tracings for the right side of the heart (Fig. 22-1A) and the left side of the heart (Fig. 22-1B) are similar except that the pressures on the right are a scaled-down version of those on the left. In both cases, the tracings begin in the middle of phase 1, that is, the period of decreased filling toward the end of diastole called diastasis (from the Greek dia [apart] + histanai [to stand]). Note that the volume changes in the left ventricle are exactly the same as those in the right ventricle because the cardiac outputs of the right and left sides of the heart are virtually identical (see Chapter 17). For purposes of illustration, we now focus on the records for the left side of the heart (Fig. 22-1B).

Diastasis Period (Middle of Phase 1) During the diastasis, the mitral valve is open, but little blood flows from the left atrium to the left ventricle; ventricular volume slowly rises and approaches a plateau. The pressures in both the left atrium and the left ventricle rise slowly, driven by the pressure in the pulmonary veins, which is only slightly higher. The atrial pressure parallels—and is only slightly above—the ventricular pressure because the mitral valve is wide open, and the flow between the two chambers is minimal. The P wave of the electrocardiogram (ECG; see Chapter 21), which corresponds to atrial excitation, occurs at the end of this phase.

Atrial Contraction (End of Phase 1) Immediately following the P wave is the atrial contraction, which causes a variable amount of blood to enter the left ventricle. In a person at rest, the atrial contraction transfers into the left ventricle a volume of blood that represents less than 20% of the subsequent stroke volume and often only a few percent. During heavy exercise, this figure can rise to 40% (see the box). Atrial contraction causes a slight rise in intra-atrial pressure and a comparable rise in ventricular pressure and volume. All during this period, the aortic pressure decreases as blood flows out to the periphery.

Isovolumetric Contraction (Phase 2) When the ventricles begin to depolarize, as evidenced by the QRS complex on the ECG, systole commences. The ventricles contract, and very soon the pressure in the left ventricle exceeds that in the left atrium (first crossover of blue and orange pressure tracings in Fig. 22-1B, top panel). As a result, the mitral valve closes. The aortic valve has been closed this entire time. Thus, the left ventricle contracts with both mitral and aortic valves closed. Because the blood has no place to go, the result is an isovolumetric contraction that causes the pressure in the left ventricle to rise rapidly, eventually exceeding the pressure in the aorta (first crossover of blue and red tracings) and causing the aortic valve to open.

Ejection or Outflow (Phase 3) As the aortic valve opens, the ejection phase begins. During the first part of phase 3—rapid ejection—ventricular pressure (blue tracing in Fig. 22-1B, top) continues to rise, closely followed by a rapid elevation of aortic pressure, which at first is slightly less (red tracing). Accompanying these rapid pressure increases is a precipitous reduction in ventricular volume (Fig. 22-1B, bottom), as blood flows into the aorta. Aortic pressure continues to rise and eventually exceeds ventricular pressure (second crossover of blue and red tracings in Fig. 22-1B, top) just before both the aortic and ventricular pressures begin to fall. Despite the reversal of the pressure gradient across the aortic valve, the cusps of the aortic valve do not immediately snap shut because of the inertia of blood flow, which imparts considerable kinetic energy to the blood. During the latter part of phase 3—decreased ejection—the decrease in ventricular volume becomes less rapid, and both the ventricular and aortic pressures fall off. During the entire ejection phase, about 70mL of blood flows into the aorta, leaving about 50mL behind in the ventricle. (See Note: Inertial Component of Flow in the Aorta)

Isovolumetric Relaxation (Phase 4) Late in the ejection phase, blood flow across the aortic valve falls to extremely low values, until it actually reverses direction (i.e., retrograde or negative flow). At this point, the aortic valve closes, defining the onset of diastole. As blood flow in the aorta again becomes briefly positive (i.e., forward), there is a small upward deflection in the aortic pressure trace. The result is the dicrotic notch (from the Greek dikrotos, “double-beat”), or incisura, and the subsequent dicrotic wave, which interrupts the generally downward trend of aortic pressure. Because both the aortic and mitral valves are closed and no blood can enter the left ventricle, this is the period of isovolumetric relaxation. Pressure falls rapidly in the left ventricle.

Rapid Ventricular Filling Period (Beginning of Phase 1) When ventricular pressure falls below that in the left atrium (second crossover of blue and orange tracings in Fig. 22-1B), the mitral valve opens. Immediately following mitral valve opening, left ventricular volume begins to increase rapidly (Fig. 22-1B, bottom). During this period of rapid ventricular filling, the left atrial and ventricular pressures evolve in parallel because the mitral valve is wide open. A period of relatively decreased filling follows, the period of diastasis with which we began our discussion. Thus, diastole includes both the rapid ventricular filling period and diastasis. As already noted, the length of diastole decreases with elevations in heart rate. This decrease comes first at the expense of the period of slower ventricular filling (i.e., diastasis).

Importance (and Unimportance) of Atrial Contraction

The relative importance of atrial contraction to overall cardiac function is evident in patients who develop atrial fibrillation (see Fig. 21-14H), an arrhythmia associated with loss of this atrial “kick.” In atrial fibrillation, chaotic electrical activity, bombarding the atria with as many as 500 impulses per minute from all directions, prevents the concerted action of atrial cardiac muscle fibers that is necessary for coordinated atrial contraction. As a result, the atria fibrillate—they look like a wriggling bag of worms. In healthy persons with otherwise normal hearts, the loss of atrial contraction usually causes no symptoms at rest or perhaps only a sensation of an irregular or rapid heartbeat (i.e., the result of atrial fibrillation). However, if the patient already has a compromised myocardium (e.g., from ischemic heart disease, prolonged hypertension, or mitral stenosis), or if the patient is debilitated by dysfunction of other organs (e.g., chronic emphysema), the loss of the atrial contraction may further reduce cardiac output just enough to send the patient into florid congestive heart failure (see the box on cardiac hypertrophy) or even shock (i.e., arterial pressure so low that it compromises perfusion of peripheral tissues). The physician may treat a patient with an otherwise normal heart in a leisurely fashion or perhaps not at all—keeping in mind that patients with atrial fibrillation are at high risk for development of atrial thrombosis and thus possibly cerebral embolism and stroke. In patients with a compromised myocardium, emergency chemical or electrical cardioversion may be necessary.

During rapid ventricular filling, the aortic valve remains closed. Because blood continues to flow out to the periphery, owing to the recoil of the aorta’s elastic wall (see the box on the effect of aortic compliance later in this chapter), the aortic pressure falls. This fall continues during diastasis.

The electrocardiogram, phonocardiogram, and echocardiogram follow the cyclic pattern of the cardiac cycle

Accompanying the basic cyclic pattern of cardiac pressure and volume changes are characteristic mechanical, electrical, acoustic, and echocardiographic changes. Figure 22-2 illustrates these events for the left side of the heart and the systemic circulation. Notice that the pressure records in the top panel of Figure 22-2 start with the atrial contraction, that is, slightly later than in Figure 22-1.


Figure 22-2 Mechanical, electrical, acoustic, and echocardiographic events in the cardiac cycle. Top, A repeat of Figure 22-1B, with three modifications: (1) the cardiac cycle begins with atrial contraction; (2) phase 1 of the cardiac cycle has three subparts: rapid ventricular filling, decreased ventricular filling, and atrial systole; (3) phase 3 has two subparts: rapid and decreased ventricular ejection.

Aortic Blood Flow Blood flow from the left ventricle to the ascending aorta rises most rapidly during the rapid ejection phase of the left ventricle. The peaking of aortic flow defines the beginning of the decreased ejection phase. See Figure 22-2, second panel.

Jugular Venous Pulse The third panel of Figure 22-2 includes the jugular venous pulse, for comparison with the timing of other events. We discuss the jugular venous pulse later in the chapter.

Electrocardiogram The ECG (discussed in Chapter 21) begins with the middle of the P wave (atrial depolarization). The QRS complex (ventricular depolarization) is the prelude to the upswing in ventricular pressure. The T wave (ventricular repolarization) occurs in the decreased ejection phase. See Figure 22-2, fourth panel.

Phonocardiogram and Heart Sounds The opening and closing of the valves are accompanied by heart sounds, easily heard through a stethoscope or recorded with a digital stethoscope and stored as a phonocardiogram (Table 22-2). The dominant frequencies of heart sounds are lower (110 to 180Hz) than those of heart murmurs (see Chapter 17), which result from turbulence (180 to 500 Hz). Each of the vertical dotted lines in Figure 22-2 indicates the movement of two valves, one on the right side of the heart and one on the left. Thus, two valves can contribute to a single heart sound, although the two components can often be separated by the ear. The phonocardiogram in Figure 22-2 shows the timing of the two major, or physiological, heart sounds (S1 and S2) as well as two other sounds (S3 and S4) that are occasionally heard. See Figure 22-2, fifth panel.

Table 22-2 The Heart Sounds


The physiological heart sounds S1 and S2 are heard following the closure of the cardiac valves: the mitral and tricuspid valves for S1, and the aortic and pulmonary valves for S2. However, the actual apposition of the valve leaflets (i.e., “slamming the door”) does not produce the sound. Instead, vibrations resulting from sudden tension in the AV valves and the adjacent ventricular walls produce the first heart soundS1. Similarly, vibrations of the large vessel walls and columns of blood produce the second heart soundS2, following closure of the semilunar valves. These vibrations propagate through adjacent tissues to the chest wall, where one can normally hear the first and second heart sounds through a stethoscope. S1 is usually stronger, longer, and of lower frequency than S2.

Although the four vertical lines that define the four phases of the cardiac cycle are similar for the right and left sides of the heart, they do not line up perfectly with one another, as can be seen by comparing A and B of Figure 22-1. For example, the aortic valve usually closes just before the pulmonary valve. This timing difference produces the physiological splitting of the A2 (i.e., aortic) and P2 (i.e., pulmonary) components of the second heart sound. As we shall see later, inspiration accentuates the splitting of S2. Pathological change that accentuates the asynchrony between the left and right sides of the heart (e.g., right bundle branch block) may also lead to splitting of the first heart sound.

With stiffening of the mitral valve, seen in mitral stenosis, the opening of the mitral valve may produce an additional sound, an opening snap (OS), in early diastole just after S2.

A physiological third heart soundS3, is present in some normal individuals, particularly children. S3 occurs in early diastole when rapid filling of the ventricles results in recoil of ventricular walls that have a limited distensibility. An S3 also can be heard in adults when the ventricle is so overfilled at the end of systole that the addition of 70 mL more blood during diastole brings the ventricle into a volume range in which ventricular compliance is very low. The result is an accentuated recoil, heard as an S3. An S3 can originate from the left or the right side of the heart. A gallop rhythm is a grouping of three heart sounds that together sound like hoofs of a galloping horse. Thus, the addition of an S3 to the physiological S1 and S2 creates a three-sound sequence, S1-S2-S3, that is termed a protodiastolic gallop or ventricular gallop.

When present, a fourth heart soundS4, coincides with atrial contraction. It is usually heard in pathological conditions in which an unusually strong atrial contraction occurs in combination with low compliance of the left ventricle. The addition of an S4 produces another three-sound sequence, S4-S1-S2, which is also a gallop rhythm, a presystolic gallop or atrial gallop.

Echocardiogram We discussed echocardiography in Chapter 17. The echocardiogram in the bottom panel of Figure 22-2 shows that the separation between the anterior and posterior leaflets of the mitral valve increases during atrial contraction. The leaflets meet at the beginning of phase 2 and remain together until rapid ventricular filling occurs in the beginning of phase 1, when the separation between leaflets becomes maximal. During the decreased phase of ventricular filling, the leaflets once again move closer together, until the next atrial contraction.

The cardiac cycle causes flow waves in the aorta and peripheral vessels

With the closing and opening of the heart’s exit valves (i.e., pulmonary and aortic valves), blood flow and blood velocity across these valves oscillate from near zero, when the valves are closed, to high values, when the valves are open. Blood flow in the aortic arch actually oscillates between slightly negative and highly positive values (Fig. 22-3A, panel 1). Pressure in the aortic arch typically oscillates between about 80 and about 120 mm Hg (Fig. 22-3B, panel 1) but varies greatly among individuals. Phasic changes in pressure and flow also occur in the peripheral arteries. Arterial pressure is usually measured in a large artery, such as the brachial artery. Because very little pressure drop occurs between the aorta and such a large, proximate artery, the measured systolic and diastolic arterial pressures, as well as the pulse pressure and mean arterial pressure, closely approximate the corresponding aortic pressures. (See Chapter 17.)


Figure 22-3 Flow (A) and pressure (B) profiles in the aorta and smaller vessels.

If blood vessels were rigid tubes, so that the resistance (R) were constant, and if the driving pressure (ΔP) were also constant throughout the cardiac cycle, we could describe blood flow (F) by a simple Ohm’s law–like relationship, as we did in Chapter 17 (see Equation 17-1). However, because blood vessels are compliant (so that R varies with pressure, as in Fig. 19-7B) and because both aortic pressure and flow vary during the cardiac cycle, we cannot describe real arteries in this way. In the field of hydraulics, oscillating flows and pressures have not only an amplitude but also a phase. As a result, the ratio ΔP/F is no longer resistance—a simple, time-independent quantity—but a complex quantity called the mechanical impedance that depends on the classical “resistance” as well as the compliance and inertial properties of the vessels and blood. (See Note: Mechanical Impedance of Blood Flow)

Because of these resistive, compliant, and inertial properties, the pressure and flow waves in vessels distal to the aorta are not quite the same as in the aorta. Instead, the farther the vessels are from the aorta, the more different the pressure and flow waves become.

Aortic Arch During the rapid ejection phase, peak flow through the aortic arch is remarkably high, ~30 L/min (dark beige band in Fig. 22-3A, panel 1). The peak linear velocity is ~100 cm/s, which makes it more likely that the blood will reach the critical Reynolds number value for turbulence (see Chapter 17). The rapid ejection of blood also causes a rapid rise of the pressure in the aorta to above that in the ventricle (Fig. 22-3B, panel 1). Even though the pressure gradient across the valve reverses, the valve does not close, as is evidenced by the continuous flow of blood from the ventricle into the aorta. The reason that flow continues in the forward direction is the inertial component of the blood flow, which represents considerable kinetic energy. Eventually, blood in the aortic arch decelerates sufficiently that the flow becomes zero and eventually negative (producing reflux through the valve). As the aortic valve closes, it produces the dicrotic notch in the aortic pressure trace.

Thoracic-Abdominal Aorta and Large Arteries Just distal to the aortic arch, a transformation of the flow and pressure curves begins to occur. The records in panels 2 to 4 in Figure 22-3A show the flow curves for the abdominal aorta and some of its large branches. Peak systolic flow becomes smaller as one moves from the aorta toward the periphery (i.e., iliac and femoral arteries), as would be predicted because of the branching of the vessels. However, in the abdominal aorta, a new phenomenon is seen. As the elastic aorta—which stored blood during systole—releases blood during diastole, a second peak of flow appears. Note that this diastolic component of flow is larger in the abdominal aorta than in the more distal iliac artery and almost absent in the femoral artery. Of particular importance is the sizable diastolic flow in the carotid and renal arteries (panels 5 and 6). The basis for the diastolic component of flow is the subject of the box titled Effect of Aortic Compliance on Blood Flow.

The cardiac cycle also causes pressure waves in the aorta and peripheral vessels

The pressure curves in Figure 22-3B show that with increasing distance from the heart (panels 1 to 4), the rising portion of the wave becomes steeper and the peak is narrower. Because the peak gradually increases in height and the minimum pressure gradually decreases, the pulse pressure becomes greater. With increasing distance from the heart, an important secondary pressure oscillation appears during diastole (Fig. 22-3B, fourth panel). Thus, although the pressure waves are distorted, they are not damped. Although it might seem counterintuitive that the peak arterial pressure should increase as we get farther from the heart (i.e., Is the blood flowing against a pressure gradient?), it turns out that the mean arterial pressure does fall slightly with increasing distance from the heart.

Effect of Aortic Compliance on Blood Flow

There is a large diastolic component to total blood flow in the large arteries that lie close to the aorta, such as the carotid and renal arteries (panels 5 and 6 in Fig. 22-3A). This sizeable diastolic component is largely the result of the high compliance of the vessel walls and the radial expansion of the vessels that occurs during ventricular ejection. We can reach at least an intuitive understanding of the radial contribution to flow in the aorta and large arteries by examining the ability of the aorta to store and to give up energy during the acceleration and deceleration of flow.

Figure 22-4 compares two branches of a hydraulic system that are identical in radius and length. One branch (branch 1) is rigid and made of glass, the other (branch 2) is elastic and made of rubber. Both branches terminate in a spout with an outflow resistance that is analogous to the resistance of arterioles. We assume that the resistance of the spout is much greater than that of the glass or rubber tube, so that we can ignore small changes in the diameter of the rubber tube on overall resistance. If we apply a steady pressure to both branches, the flows through the two branches are continuous and identical (Fig. 22-4A).


Figure 22-4 A and B, Effect of pulsatile pressure on flow through a compliant vessel. In C, the gold arrows indicate movements analogous to systole, and the violet arrows, diastole.

However, if we apply the pressure in square pulses, the flows in the two branches are quite different (Fig. 22-4B). The flow through the glass tubing instantly rises to a maximum value with the onset of the pressure wave and then instantly falls to zero when the driving pressure falls to zero. Thus, the plot of flow through the glass tube perfectly mirrors the plot of the applied square wave pressure. The flow through the rubber tube has a very different profile. During the interval of peak pressure, the rubber vessel gradually dilates, storing a volume of fluid. Therefore, the flow rises slowly to its maximum value. During the interval of the cycle when the driving pressure falls to zero, the expanded rubber vessel delivers its stored volume downstream, resulting in some forward flow despite the absence of any pressure head. The time-averaged outflow from the rubber tube exceeds that from the glass tube.

The aorta and large vessels behave like the rubber tube in Figure 22-4B. The oscillating pressure head in our model (i.e., between zero and a peak value) represents ventricular pressure. The maintenance of flow during interruption of the pressure head is equivalent to the continuing flow from the aorta during diastole.

Figure 22-4C shows an alternative mechanical model, that of a Windkessel (German for “wind chamber”), in which we replace the compliance of a distensible rubber tube with the compressible air within a chamber above the blood.

The two models illustrated in Figure 22-4B and C show how compliant blood vessels can convert discontinuous flow into a more continuous flow. The so-called Windkessel action of the arterial system considerably improves the efficiency of the pump (i.e., the heart) because the vessels are able to convert the phasic flow peaks of the pump into a more continuous flow.

Terminal Arteries and Arterioles In the smallest arteries, the flows must be small. Here, the trend toward an increased peak pressure reverses. Instead, the pulse wave gets damped out for two reasons. First, because we are dealing with many parallel vessels with a large aggregate wall area, the aggregate compliance increases, damping the pressure wave. Second, because these smaller arteries have a smaller radius and thus a far greater resistance, the mean arterial pressure must fall in proportion to the much higher resistance. Thus, in contrast to the situation in the larger arteries, damping predominates over distortion.

Capillaries By the time the blood reaches the capillaries, the damping is so severe that pulsations (i.e., pressure oscillations) do not normally occur—blood flow is continuous. The pulmonary capillaries are an exception; their upstream vessels are short, and they have low resistance and high compliance. The pulsation of systemic capillaries occurs only in cases of markedly increased pulse pressure, such as in patients with aortic regurgitation or hyperthyroidism, or in cases of generalized peripheral vasodilation.

Distortion of pressure waves is the result of their propagation along the arterial tree

Imagine that you are listening to a patient’s heart with a stethoscope while simultaneously feeling the pulse of the radial artery near the wrist. For each heartbeat that you hear, you feel a radial pulse. You know that the peak pressure in the left ventricle occurs about midway between the first and second heart sounds, but the delay between the midpoint of the two heart sounds and the peak of the radial pulse is only ~0.1 s. Red blood cells take several seconds to flow from the heart to the wrist. Why, then, are you able to feel the pulse so soon after the heartbeat?

The answer is that the blood vessels conduct the palpable pulse as a pressure wave. The linear velocity of red blood cells—carried in the blood by convection—ranges from approximately 1 m/s in the aorta to vanishingly small values in the capillaries (see Chapter 19). However, the pressure wave travels at a velocity of 5 to 6 m/s in the aorta, increasing to 10 to 15 m/s in the small arteries.

The following example illustrates the difference between the velocity of a pressure wave and that of convection. Imagine that two people are submerged in a river, floating downstream (convection). Now the person upstream makes a sound under water. The sound waves (an example of a pressure wave) travel to the person downstream with a velocity that is far greater than the velocity of the river.

We could illustrate how pressure waves propagate along arteries by replotting the arterial pressure profiles from Figure 22-3B and stacking them one on top of the other. The four pressure waves in Figure 22-5actually represent data obtained simultaneously in a dog with four catheters, the first placed in the aortic arch and the last three placed precisely 10 cm downstream from the previous one. The downstream propagation of the wave through the larger arteries is accompanied by a serious distortion of the pressure profile: it gets narrower and taller as we move downstream.


Figure 22-5 Arterial pressure waves. These simultaneous pressure records are from a dog, with catheters placed at 0, 10, 20, and 30 cm from the aortic arch. As the wave moves down the vessel, the upstroke is delayed, but the peak is higher.

Effect of Frequency on Wave Velocity and Damping The pressure wave moving from the aorta to the periphery is actually an ensemble of many individual waves, each with its own frequency. Higher frequency waves travel faster and undergo more damping than low-frequency waves (Fig. 22-6A). Recombination of these waves at a more peripheral site thus produces a new wave with a shape that is a distorted version of the original aortic wave. (See Note: Distortion of Propagated Waves)


Figure 22-6 Propagation of pressure waves. In A and B, the flow is from left to right. The left pair of pressure waves is at the same early time, whereas the right pair of pressure waves is at the same late time. If on the right (i.e., end of the vessel) we summate waves of different frequency at the same instant in time, then the composite wave is distorted (like the green femoral artery curve in Fig. 22-5).

Effect of Wall Stiffness on Wave Velocity As the pressure wave reaches vessels that have a stiffer wall (e.g., greater ratio of wall thickness to vessel diameter), the velocity of the wave increases (Fig. 22-6B). Conversely, with a more compliant vessel, some of the energy of the pressure pulse goes into dilation of the vessel, so that the pressure wave spreads out and slows down. Because aging causes a decrease in vessel compliance (i.e., distensibility), the velocity of propagation actually increases. (See Note: Distortion of Propagated Waves)

Pressure waves in veins do not originate from arterial waves

We have seen earlier in this chapter that flow in capillaries is usually not pulsatile. Nevertheless, blood flow in systemic capillaries can exhibit slow oscillations, unrelated to the cardiac cycle. The action of upstream vasomotor control elements in arterioles and precapillary sphincters can cause fluctuations. In addition, changes in tissue pressure (e.g., caused by muscle contraction) can compress capillaries and cause further fluctuations in capillary flow. Pulmonary capillaries are especially susceptible to changes in the surrounding alveolar pressure (see Chapter 31).

Although systemic veins have pressure waves, these waves do not originate from arterial waves propagating through the capillary beds, which are nonpulsatile. Three mechanisms can contribute to the venous pulse: (1) retrograde action of the heartbeat during the cardiac cycle, (2) the respiratory cycle, and (3) the contraction of skeletal muscles.

Effect of the Cardiac Cycle A large vein close to the heart, such as the jugular vein, has a complex pulse wave (Fig. 22-7A) synchronized to the cardiac cycle. The three maxima, or peaks, in the jugular pulse wave are labeled a, c, and v. The three minima, or dips, are labeled av, x, and y. These pressure transients reflect events in the cardiac cycle:

The a peak is caused by the contraction of the right atrium.

The av minimum is due to relaxation of the right atrium and closure of the tricuspid valve.

The c peak reflects the pressure rise in the right ventricle early during systole and the resultant bulging of the tricuspid valve—which has just closed—into the right atrium.

The x minimum occurs as the ventricle contracts and shortens during the ejection phase, later in systole. The shortening heart—with tricuspid valve still closed—pulls on and therefore elongates the veins, lowering their pressure.

The v peak is related to filling of the right atrium against a closed tricuspid valve, which causes right atrial pressure to rise. As the tricuspid valve opens, the v peak begins to wane.

The y minimum reflects a fall in right atrial pressure during rapid ventricular filling, as blood leaves the right atrium through an open tricuspid valve and enters the right ventricle. The increase in venous pressure after the y minimum occurs as venous return continues in the face of reduced ventricular filling.


Figure 22-7 Venous pressure changes. In A, the time scale is a single cardiac cycle. The relative heights of the peaks and valleys are variable. In B and C, the time scale surrounds one protracted inspiration (i.e., several heartbeats); the y-axis shows the mean jugular venous pressure. (B, Data from Brecher GA: Venous Return. New York, Grune & Stratton, 1956. C, Data from Pollack AA, Wood EH: Venous pressure in the saphenous vein at the ankle in man during exercise and changes in posture. J Appl Physiol 1949; 1:649-662.)

Effect of the Respiratory Cycle Poiseuille was the first to observe that the pressure in the jugular vein becomes negative during inspiration (Fig. 22-7B). During inspiration, the diaphragm descends, causing intrathoracic pressure (and therefore the pressure inside the thoracic vessels) to decrease and intra-abdominal pressure to increase (see Chapter 27). Consequently, the venous return from the head and upper extremities transiently increases, as low-pressure vessels literally suck blood into the thoracic cavity. Simultaneously, the venous flow decreases from the lower extremities because of the relatively high pressure of the abdominal veins during inspiration. Therefore, during inspiration, pressure in the jugular vein falls while pressure in the femoral vein rises.

Effect of Skeletal Muscle Contraction (“Muscle Pump”) The contraction of skeletal muscle can also affect pressure and flow in veins. Large veins in the lower limbs are equipped with valves that prevent retrograde movement of blood (see Chapter 17). When a person is at rest and in the recumbent position, all venous valves are open and venous blood flow toward the heart is continuous. Standing causes the venous pressure in the foot to rise gradually to the hydrostatic pressure dictated by the vertical blood column from the foot to the heart (Fig. 22-7C). If the person begins to walk, the combination of the pumping action of the leg muscles on the leg veins and the action of the venous valves as hydrostatic relay stations causes the venous pressure in the foot to decrease. Each step causes both a small oscillation and a small net decrease in foot vein pressure. Once foot vein pressure has bottomed out, each step simply causes a small pressure oscillation. Walking causes a net decrease in pressure in both the superficial and deep foot veins as well as in the corresponding capillaries. When the exercise ceases, the venous pressure again rises.


The heart is a system of two pumps linked in series. The muscular wall of the left ventricle is thicker and more powerful than that of the right. The interventricular septum welding the two pumps together is even thicker. The thick muscular walls of the ventricles are responsible for exerting the heart’s pumping action.

The heart does not depend on a rhythm generator in the brain, like the central pattern generators (see Chapter 16) that drive other rhythmic behaviors, such as respiration, locomotion, chewing, and shivering. Instead, pacemaker cells within the heart itself initiate cardiac excitation. When the heart is in a normal sinus rhythm, the pacemaker cells setting the rate are located in the sinoatrial (SA) node of the right atrium (see Chapter 21). The action potential then spreads through atrial myocytes and specialized tracts or bundles. The impulse cannot cross from the atria to the ventricles except through the atrioventricular (AV) node. The AV node inserts a time delay into the conduction that is essential to allow the ventricles to finish filling with blood before contraction and ejection occur. From the AV node, the impulse spreads through the bundle of His and then the right and left bundle branches; the left bundle branch divides in an anterior and posterior fascicle. Finally, the system of Purkinje fibers excites the ventricular myocytes, where the impulse propagates from cell to cell through gap junctions.

The right ventricle contracts like a bellows, whereas the left ventricle contracts like a hand squeezing a tube of toothpaste

The two ventricles share a common envelope of spiral and circular muscle layers. The arrangement of the spiral bundles ensures that ventricular contraction virtually wrings the blood out of the heart, although incompletely. The apex contracts before some of the basal portions of the ventricle, a sequence that propels blood upward to the aortic and pulmonary valves.

The mechanical action of the right ventricle resembles that of a bellows used to fan a fire (Fig. 22-8A). Although the distance between the free wall and the septum is small, the free wall has such a large surface area that a small movement of the free wall toward the septum ejects a large volume.


Figure 22-8 Comparison of the dynamics of the left and right ventricles.

The mechanism of emptying of the right ventricle involves three motions. First, the longitudinal axis of the right ventricle shortens when spiral muscles pull the tricuspid valve ring toward the apex. Second, the free wall of the right ventricle moves toward the septum in a bellows-like motion. Third, the contraction of the deep circular fibers of the left ventricle forces the septum into a convex shape, so that the septum bulges into the right ventricle. This bulging of the septum stretches the free wall of the right ventricle over the septum. These three motions are well suited for ejection of a large volume but not for development of a high pressure. The right ventricle ejects the same blood volume as the left ventricle does, but it does so at much lower intraventricular pressures.

The mechanical action of the left ventricle occurs by a dual motion (Fig. 22-8B). First, constriction of the circular muscle layers reduces the diameter of the chamber, progressing from apex to base, akin to squeezing a tube of toothpaste. Second, contraction of the spiral muscles pulls the mitral valve ring toward the apex, thereby shortening the long axis. The first mechanism is the more powerful and is responsible for the high pressures developed by the left ventricle. The conical shape of the lumen gives the left ventricle a smaller surface-to-volume ratio than the right ventricle and contributes to the ability of the left ventricle to generate high pressures.

The contraction of the atria normally makes only a minor contribution to the filling of the two ventricles when the subject is at rest (see the box on the importance of atrial contraction). However, the contraction of the atria is a useful safety factor in at least two circumstances. During tachycardia, when the diastolic interval—and thus the time for passive filling—is short, the atrial contraction can provide a much-needed boost. Atrial contraction is also useful in certain pathological conditions. For example, when a narrowed (i.e., stenotic) AV valve offers substantial resistance to the flow of blood from atrium to ventricle, the atrial pump can make an important contribution to ventricular filling.

The right atrium contracts before the left, but the left ventricle contracts before the right

When the cardiac cycle was introduced earlier in the chapter, we assumed that the events on the right and left sides of the heart happen simultaneously. However, as we have already noted in our discussion of the splitting of heart sounds, the timing of the two sides of the heart is slightly different (Fig. 22-8C).

Atrial Contraction Because the SA node is located in the right atrium, atrial contraction begins and ends earlier in the right atrium than in the left (Fig. 22-8C, Contraction).

Initiation of Ventricular Contraction Ventricular contraction starts slightly earlier on the left side, and the mitral valve closes before the tricuspid valve. However, this timing difference in the closure of the AV valves (Fig. 22-8C, Valve movements) is so small that it is unusual to hear a split S1. On the other hand, the right ventricle has a briefer period of isovolumetric contraction because it does not need to build up as much pressure to open its semilunar (i.e., outflow) valve and to initiate ejection. Thus, the pulmonary valve opens slightly ahead of the aortic valve.

Ventricular Ejection Ejection from the right ventricle lasts longer than that from the left. The semilunar valves do not close simultaneously. The aortic valve, with its higher downstream pressure, closes before the pulmonary valve. Therefore, the pulmonary valve—with its lower downstream pressure—opens first and closes last. This timing difference in the closure of the semilunar valves explains the normal physiological splitting of S2 (Fig. 22-8C, Sounds). During inspiration, the relatively negative intrathoracic pressure enhances filling of the right side of the heart, causing it to have a larger end-diastolic volume and therefore more blood to eject. The additional time required for right ventricular ejection postpones the closure of the pulmonary valve (P2), broadening the physiological splitting of S2.

Ventricular Relaxation Isovolumetric relaxation is briefer in the right side of the heart than in the left. The pulmonary valve closes after the aortic valve, and the tricuspid valve opens before the mitral valve. Therefore, the right ventricle begins filling before the left.

Measurements of ventricular volumes, pressures, and flows allow clinicians to judge cardiac performance

Definitions of Cardiac Volumes The cardiac output is the product of heart rate and stroke volume (see Chapter 17). The stroke volume (SV) is the difference between ventricular end-diastolic volume (EDV) and ventricular end-systolic volume (ESV), that is, the difference between the maximal and minimal ventricular volumes. EDV is typically 120 mL, and the ESV is 50 mL, so that


The ejection fraction (EF) is a dimensionless value, defined as the stroke volume normalized to the end-diastolic volume:


In our example, the EF is (70 mL)/(120 mL) or ~0.6. The value should exceed 55% in a healthy person. Whereas the ejection fractions of the left and right ventricles are as a rule equal, clinicians normally measure left ventricular ejection fraction (LVEF).

Measurements of Cardiac Volumes Clinicians routinely measure the volume of the cardiac chambers by means of angiography or echocardiography (see Chapter 17). One-dimensional (or M-mode) echocardiography allows one to assess left ventricular performance in terms of linear dimensions and velocities by providing measurements of velocity of the posterior left ventricular wall, fractional shortening of the left ventricular circumference, and rate of fractional circumferential shortening. Two-dimensional echocardiography makes it possible to determine several ventricular volumes:

Left ventricular end-diastolic volume (LVEDV)

Left ventricular end-systolic volume (LVESV)

Stroke volume (SV = LVEDV − LVESV)

Left ventricular ejection fraction (LVEF = SV/LVEDV)

Measurement of Ventricular Pressures For right-sided heart catheterizations, clinicians use a Swan-Ganz catheter, which consists of three parallel tubes of different lengths. The longest is an end-hole catheter with a balloon flotation device that directs the tip in the direction of the blood flow. The other two tubes are side-hole catheters that terminate at two points proximal to the tip. The physician advances the catheter percutaneously through a large systemic vein, into the right side of the heart, and then into the pulmonary circulation where the tip of the longest tube literally wedges in a small pulmonary artery. Because a continuous and presumably closed column of blood connects the probe’s end and the left atrium, the wedge pressure is taken as an index of left atrial pressure. For left-sided heart catheterizations, clinicians insert a simple catheter percutaneously into an artery and then advance the catheter tip upstream to the left side of the heart. Table 22-3 lists some of the most important pressure values for the right and left sides of the heart.

Table 22-3 Comparison of Pressures in the Right and Left Circulations


Measurement of Flows The cardiologist can calculate flow from changes in ventricular volume, as measured by echocardiography and the Doppler ultrasound technique (see Chapter 17), both of which measure the flow of blood in the outflow tract (i.e., aorta). Figure 22-8D and E illustrates the profiles of outflow pressure and velocity for the two ventricles. Although the two ventricles expel on average the same amount of blood in a single cardiac cycle, the peak velocity is much higher in the left ventricle. In addition, the velocity rises far more rapidly in the left ventricle, indicating greater acceleration of the blood during ejection. The pressure wave is about five times larger in the left ventricle than in the right, and the rate at which the pressure rises (ΔPt) is more rapid in the left ventricle.

The pressure-volume loop of a ventricle illustrates the ejection work of the ventricle

In Figure 22-1, we saw separate plots of ventricular pressure against time and volume against time. If, at each point in time, we now plot pressure against volume, the result is a pressure-volume loop, as shown in Figure 22-9 for the left ventricle. This loop is a “phase plot” that describes the relationship between left ventricular pressure and left ventricular volume during the cardiac cycle. Notice that although timedoes not explicitly appear in this plot, as we make one complete counterclockwise cycle around the loop, we sequentially plot pressure and volume at all time points of the cardiac cycle. However, the distance between two points on the loop is not proportional to elapsed time.


Figure 22-9 Pressure-volume loop of the left ventricle.

In examining this pressure-volume loop, we will arbitrarily start at point A in Figure 22-9 and then consider each of the segments of the loop (e.g., AB, BC, and so on) before again returning to point A. Although we use the left ventricle as an example, a similar analysis applies to the right ventricle.

Segment AB Point A represents the instant at which the mitral valve opens. At this point, left ventricular volume is at its minimal value of ~50 mL, and left ventricular pressure is at the fairly low value of ~7 mm Hg. As the mitral valve opens, the ventricle begins to fill passively because atrial pressure is higher than ventricular pressure. During interval AB, ventricular pressure falls slightly to ~5 mm Hg because the ventricular muscle is continuing to relax during diastole. Thus, despite the rapid entry of blood, ventricular pressure falls to its lowest value in the cardiac cycle.

Segment BC During a second phase of ventricular filling, volume rises markedly from ~70 to ~120 mL, accompanied by a rather modest increase in pressure from ~5 to ~10 mm Hg. The modest rise in pressure, despite a doubling of ventricular volume, reflects the high compliance (C = ΔVP) of the ventricular wall during late diastole. The relationship between pressure and volume during segment BC is similar to that in blood vessels (see Chapter 19).

Segment CD Point C represents the closure of the mitral valve. At this point, ventricular filling has ended and isovolumetric contraction—represented by the vertical line CD—is about to begin. Thus, by the definition of isovolumetric, ventricular volume remains at 120 mL while left ventricular pressure rises to ~80 mm Hg, about equal to the aortic end-diastolic pressure.

Segment DE Point D represents the opening of the aortic valve. With the outlet to the aorta now open, the ventricular muscle can begin to shorten and to eject blood. During this period of rapid ejection, ventricular volume decreases from ~120 to ~75 mL. Notice that as contraction continues during interval DE, the ventricular pressure rises even farther, reaching a peak systolic value of ~130 mm Hg at point E.

Segment EF Point E represents the instant at which the ventricular muscle starts to relax. During this period of decreased ejection, ventricular pressure falls from ~130 to ~100 mm Hg. Nevertheless, blood continues to leave the ventricle, and ventricular volume falls from ~75 mL at point E to ~50 mL at point F. Point F represents end-systolic volume and pressure. Notice that the ventricle does not shrink to zero volume at the end of systole. In total, 120 − 50 or 70 mL of blood has left the ventricle during systole (i.e., between points D and F). Therefore, the stroke volume is substantially less than the maximum ventricular volume (i.e., EDV). The ejection fraction in this example is ~60%, which is in the normal range. Ejection occurs against aortic pressures ranging between 80 and 130 mm Hg. Therefore, ejection is not “isotonic” (see Chapter 9).

Segment FA Point F represents the closing of the aortic valve. At this point, ejection has ended and isovolumetric relaxation is about to begin. The ventricular volume remains at 50 mL, while left ventricular pressure falls from ~100 mm Hg at point F to ~7 mm Hg at point A. At the end of isovolumetric relaxation, the mitral valve opens and the cardiac cycle starts all over again with ventricular filling.

The six segments of the pressure-volume loop correspond to different phases of the cardiac cycle:

Phase 1, the inflow phase, includes segments AB and BC.

Phase 2, isovolumetric contraction, includes segment CD.

Phase 3, the outflow phase, includes segments DE and EF.

Phase 4, isovolumetric relaxation, includes segment FA.

Segments CDEF represent systole, whereas segments FABC represent diastole.

The “pumping work” done by the heart is a small fraction of the total energy the heart consumes

The heart does its useful work as a pump by imparting momentum to the blood and propelling it against the resistance of the periphery.

Work, in its simplest definition, is the product of the force applied to an object and the distance the object moves (W = force × distance). In considering pressure-volume work, we must revise this definition. Imagine that we have a volume of blood in a syringe. If we apply a constant force to the plunger, that is, if we apply a constant pressure to the blood, the plunger moves a certain distance as we eject the blood through a needle, thereby reducing blood volume by an amount ΔV. How much work have we done? For pressure moving a fluid, the external work is


If the aortic pressure were constant, the work done with each heartbeat would be simply the product of the aortic pressure (P) and the stroke volume (ΔV = SV = EDV − ESV).

The pressure-volume relationships in Figure 22-10 illustrate the pressure-volume work of the left ventricle. The surface below the segment ABC (i.e., filling phase) is the work done by the blood (previously contained in the venous reservoirs and atria at a low pressure) on the ventricle (Fig. 22-10A). The surface below DEF (i.e., ejection phase) is the work done by the heart on the blood during the ejection (Fig. 22-10B). The difference between the areas in parts A and B of Figure 22-10—that is, the area within the single-cycle loop—is the net external work done by the heart (Fig. 22-10C). (See Note: "Pumping Work" Done by the Heart)


Figure 22-10 External work of the left ventricle.

The pressure-volume diagram for the right ventricle has the same general shape. However, the area (i.e., net external work) is only about one fifth as large because the pressures are so much lower.

The area of the loop in Figure 22-10C—that is, the pressure-volume work (P · V)—ignores the speed at which the ventricle pumps the blood (i.e., acceleration that the heart imparts to the blood, or the time it takes to complete one cardiac cycle). Thus, the work per beat should also include the kinetic energy (½ mv2) that the heart imparts to the ejected blood:


Of its total external work, the heart delivers only a relatively small fraction as kinetic energy. Moreover, the total external work is itself only a small portion of the total energy that the heart actually expends. Like other muscles, the heart not only shortens and performs classical work (e.g., isotonic contraction) but also maintains active tension without shortening (i.e., isometric contraction; see Chapter 9). During the isovolumetric contraction, the ventricle develops and maintains a high pressure without performing any total external work—just as we perform no useful work when we hold a weight without lifting it. However, in both isometric exercises, the muscle breaks down adenosine triphosphate (ATP) as long as it maintains isometric tension; the energy ends up as heat. This type of energy cost in heart muscle is called tension heat, which is proportional to the product of the tension of the ventricular wall (T) and the length of time (Δt) that the ventricle maintains this tension (i.e., tension-time integral). In the case of the heart, the pressure against which the ventricle must pump is a major determinant of the wall tension. (See Note: Tension Heat)

The total energy transformed in one cardiac cycle is the sum of the total external work done on the blood and the tension heat:


k is a proportionality constant that converts T • Δt into units of energy. The tension heat is the major determinant of the total energy requirements of the heart. Total external work represents a relatively small fraction (3%) of the total energy needs of the heart at rest, rising to as much as 10% during exercise. The heat developed as part of the tension-time integral remains the major component of the total energy consumption, even during exercise.

The tension heat is not only far more costly for the heart than the pressure-volume work but also of considerable practical interest for the patient with coronary artery disease who wishes to step up cardiac output during increased physical activity. The major burden for such an individual may be not so much the total external work expended in driving additional blood through the circulation (i.e., increasing the cardiac output) but rather an increase in tension heat (k · T · Δt). Thus, it is advantageous to the patient to have a low wall tension (T)—that is, a low blood pressure. It is also advantageous for the patient not to spend too much time (Δt) in systole. The heart spends a greater fraction of its time in systole when the heart rate is high. Thus, the cardiac patient is better off to increase cardiac output at low pressure and low heart rate (i.e., a low T · Δt product). The only option left is to increase stroke volume.

The ratio of the ventricle’s total external work (P · V + ½mv2) to the total energy cost (i.e., W/E) is the heart’s mechanical efficiency. Note that the mechanical efficiency has nothing to do with how effective the ventricle is at expelling blood (i.e., ejection fraction).


In Chapter 9, we examined the general features of muscle contraction and compared the properties of skeletal, cardiac, and smooth muscle. In this chapter, we examine how some of the features of cardiac muscle underlie cardiac performance.

The entry of Ca2+ from the outside triggers Ca2+-induced Ca2+ release from the SR, initiating contraction of cardiac myocytes

Excitation-contraction (EC) coupling in cardiac ventricular myocytes is similar to EC coupling in skeletal muscle (see Chapter 9). One major difference is that in the case of skeletal muscle, the initiating event is the arrival of an action potential at the neuromuscular junction, the release of acetylcholine, and the initiation of an end-plate potential. In the ventricular myocyte, action potentials in adjacent myocytes depolarize the target cell through gap junctions (see Chapter 21) and thereby generate an action potential.

As in a skeletal muscle fiber (see Chapter 9), the depolarization of the plasma membrane in the ventricular myocyte invades T tubules that run radially to the long axis of the myocyte. Unlike skeletal muscle cells, cardiac myocytes also have axial T tubules that run parallel to the long axis of the cell and interconnect adjacent radial T tubules.

Another major difference in EC coupling between cardiac and skeletal muscle is in the way that the L-type Ca2+ channels (Cav1.2, dihydropyridine receptors) in the T-tubule membrane activate the Ca2+ release channels made up of four RYR2 molecules in the sarcoplasmic reticulum (SR) membrane. In skeletal muscle, the linkage is mechanical and does not require Ca2+ entry per se. If you place skeletal muscle in a Ca2+-free solution, the muscle can continue contracting until its intracellular Ca2+ stores become depleted. In contrast, cardiac muscle quickly stops beating. Why?

In cardiac muscle, Ca2+ entry through the L-type Ca2+ channel is essential for raising of [Ca2+]i in the vicinity of the RYR2 on the SR. A subset of Cav1.2 channels may be part of caveolae. This trigger Ca2+activates an adjacent cluster of RYRs in concert, causing them to release Ca2+ locally into the cytoplasm (Ca2+-induced Ca2+ release). Such single events of Ca2+-induced Ca2+ release can raise [Ca2+]i as high as 10 μM in microdomains of ~1 μm in diameter. These localized increases in [Ca2+]i appear as calcium sparks when they are monitored with a Ca2+-sensitive dye by confocal microscopy. If many L-type Ca2+channels open simultaneously, the spatial and temporal summation of many elementary Ca2+ sparks leads to a global increase in [Ca2+]i.

The basic structure of the thin and thick filaments in cardiac muscle is the same as in skeletal muscle (see Fig. 9-5). After [Ca2+]i increases, Ca2+ binds to the cardiac isoform of troponin C (TNNC1; see Table 9-1), and the Ca2+-TNNC1 complex releases the inhibition of the cardiac isoform of troponin I (TNNI3) on actin. As a result, the tropomyosin (TPM1) filaments bound to cardiac troponin T (TNNT2) on the thin filament shift out of the way (see Fig. 9-6), allowing myosin to interact with active sites on the actin. ATP fuels the subsequent cross-bridge cycling (see Fig. 9-7). Because the heart can never rest, cardiac myocytes have a very high density of mitochondria and thus are capable of sustaining very high rates of oxidative phosphorylation (i.e., ATP synthesis).

The cross-bridge cycling causes thick filaments to slide past thin filaments, generating tension. When we discussed the mechanics of skeletal muscle in Chapter 9, we introduced the concept of a length-tension diagram (see Fig. 9-9), which is a plot of muscle tension as a function of muscle length. The length parameter in such a plot can be either the length of the whole skeletal muscle or the length of a single sarcomere. For heart muscle, which wraps around the ventricle, the length parameter can be either the ventricular volume, which is analogous to whole-muscle length, or the sarcomere length. The sarcomere, stretching from one Z line to another, is the functional unit in both skeletal and cardiac muscle.

Phosphorylation of phospholamban and of troponin I speeds cardiac muscle relaxation

With the waning of the phase 2 plateau of the cardiac action potential (see Fig. 21-4B), the influx of Ca2+ through L-type Ca2+ channels decreases, lessening the release of Ca2+ by the SR. By itself, halting of Ca2+ entry and release can only prevent a further increase in [Ca2+]i. The actual relaxation of the contractile proteins depends on three processes: (1) extrusion of Ca2+ into the extracellular fluid, (2) re-uptake of Ca2+ from the cytosol by the SR, and (3) dissociation of Ca2+ from troponin C. The last two of these processes are highly regulated.

Extrusion of Ca2+ into the Extracellular Fluid Even during the plateau of the action potential, the myocyte extrudes some Ca2+. After the membrane potential returns to more negative values, the extrusion processes gain the upper hand and [Ca2+]i falls. In the steady state (i.e., during the course of several action potentials), the cell must extrude all the Ca2+ that enters the cytosol from the extracellular fluid through L-type Ca2+ channels. As in most other cells (see Chapter 7), this extrusion of Ca2+ into the extracellular fluid occurs by two pathways: (1) a sarcolemmal Na-Ca exchanger (NCX1), which operates at relatively high levels of [Ca2+]i; and (2) a sarcolemmal Ca2+ pump (cardiac subtype 1, 2, and 4 of PMCA), which may function at even low levels of [Ca2+]i. However, PMCA contributes only modestly to relaxation. Because PMCA is concentrated in caveolae, which contain receptors for various ligands, its role may be to modulate signal transduction.

Re-uptake of Ca2+ by the SR Even during the plateau of the action potential, some of the Ca2+ accumulating in the cytoplasm is sequestered into the SR by the cardiac subtype of the Ca2+ pump SERCA2a (see Chapter 5). Phospholamban (PLN), an integral SR membrane protein with a single transmembrane segment, is an important regulator of SERCA2a. In SR membranes of cardiac, smooth, and slow-twitch skeletal muscle, unphosphorylated PLN can exist as a homopentamer that may function in the SR as an ion channel or as a regulator of Cl channels. The dissociation of the pentamer allows the hydrophilic cytoplasmic domain of PLN monomers to inhibit SERCA2a. However, phosphorylation of PLN by any of several kinases relieves phospholamban’s inhibition of SERCA2a, allowing Ca2+ resequestration to accelerate. The net effect of phosphorylation is an increase in the rate of cardiac muscle relaxation. PLN knockout mice have uninhibited SERCA2a Ca2+ pumps and thus an increased velocity of muscle relaxation. (See Note: Phospholamban)

Phosphorylation of PLN by protein kinase A (PKA) explains why β1-adrenergic agonists (e.g., epinephrine), which act through the PKA pathway (see Chapter 3), speed up the relaxation of cardiac muscle.

Dissociation of Ca2+ from Troponin C As [Ca2+]i falls, Ca2+ dissociates from troponin C (see Chapter 9), blocking actin-myosin interactions and causing relaxation. β1-Adrenergic agonists accelerate relaxation by promoting phosphorylation of troponin I, which in turn enhances the dissociation of Ca2+ from troponin C.

The overlap of thick and thin filaments cannot explain the unusual shape of the cardiac length-tension diagram

We discussed passive and active length-tension diagrams for skeletal muscle in Chapter 9 (see Fig. 9-9C and D). We obtain a passive length-tension diagram by holding a piece of resting skeletal or cardiac muscle at several predefined lengths and measuring the tension at each length (Fig. 22-11A, green and violet curves). We obtain the active length-tension diagram by stimulating the muscle at each predefined length (i.e., isometric conditions) and measuring the increment in tension from its resting or passive value (turquoise and brown curves).


Figure 22-11 Length-tension diagram. Compared with A, in B end-diastolic volume (EDV) on the x-axis is used as an index of sarcomere length. (Because EDV was difficult to measure before the days of echocardiography, Starling actually used left atrial pressure as an index of the degree of filling.) Starling measured pressure on the y-axis as an index of tension. Thus, systolic pressure replaces active tension, and diastolic pressure replaces passive tension. In C, left atrial pressure on the x-axis is used as an index of sarcomere length, and stroke work (systolic pressure × ejected volume) on the y-axis is used instead of tension.

The passive length-tension diagrams for skeletal and cardiac muscle are quite different. The passive tension of a skeletal muscle (Fig. 22-11A, green curve) is practically nil until the length of the sarcomere exceeds 2.6 μm. Beyond this length, passive tension rises slowly. On the other hand, the passive tension of cardiac muscle (violet curve) begins to rise at much lower sarcomere lengths and rises much more steeply. As a result, cardiac muscle will break if it is stretched beyond a sarcomere length of 2.6 μm, whereas it is possible to stretch skeletal muscle to a sarcomere length of 3.6 μm.

The reason for the higher passive tension is that the noncontractile (i.e., elastic) components of cardiac muscle are less distensible. The most important elastic component is the giant protein titin, which acts as a spring that provides the opposing force during stretch and the restoring force during shortening (see Fig. 9-4B).

The active length-tension diagrams also differ between skeletal and cardiac muscle. The active tension of skeletal muscle (turquoise curve) is high and varies only modestly between sarcomere lengths of 1.8 and 2.6 μm (Fig. 22-11A). In Chapter 9, we accounted for the shape of this curve as caused by the degree of myofilament overlap. In cardiac muscle (brown curve), active tension has a relatively sharp peak when the muscle is prestretched to an initial sarcomere length of ~2.4 μm. As the prestretched sarcomere length increases from 1.8 to 2.4 μm, active tension rises steeply. We cannot account for this rise by an increase in the overlap of thick and thin filaments because the filament dimensions of cardiac and skeletal muscle are similar. Rather, the rise in tension at longer sarcomere lengths in cardiac muscle probably has two general causes. (1) Raising the sarcomere length above 1.8 μm increases the Ca2+ sensitivity of the myofilaments. One mechanism controlling the Ca2+ sensitivity may be interfilament spacing between thick and thin filaments because fiber diameter varies inversely with fiber length. As we stretch the muscle to greater sarcomere lengths, the lateral filament lattice spacing is less than in an unstretched fiber so that the probability of cross-bridge interaction increases. Increased cross-bridge formation in turn increases the Ca2+ affinity of TNNC1, thereby recruiting more cross-bridges and therefore producing greater force. Another mechanism could be that as the muscle elongates, increased strain on titin either alters lattice spacing or alters the packing of myosin molecules within the thick filament. (2) Raising the sarcomere length above 1.8 μm increases tension on stretch-activated Ca2+ channels, thereby increasing Ca2+ entry from the extracellular fluid and thus enhancing Ca2+-induced Ca2+release.

As cardiac sarcomere length increases above 2.4 μm, active tension declines precipitously, compared with the gradual fall in skeletal muscle. Once again, this fall-off does not reflect a problem in the overlap of thin and thick filaments. Instead, titin increases the passive stiffness of cardiac muscle and may also impede development of active tension at high sarcomere lengths.

Starling’s law states that a greater fiber length (i.e., greater ventricular volume) causes the heart to deliver more mechanical energy

Long before the development of the sliding filament hypothesis and our understanding that active tension should depend on sarcomere length, Ernest Starling in 1914 anticipated the results of Figure 22-11A,using an isolated heart-lung preparation. Starling’s law states that “the mechanical energy set free in the passage from the resting to the active state is a function of the length of the fiber.” Therefore, the initial length of myocardial fibers determines the work done during the cardiac cycle. Figure 22-11B shows the results of experiments that Starling performed on the intact heart. Starling assumed that the initial length of the myocardial fibers is proportional to the end-diastolic volume (EDV). Further, he assumed that tension in the myocardial fibers is proportional to the systolic pressure. Therefore, starting from a volume-pressure diagram, Starling was able to reconstruct an equivalent length-tension diagram (Table 22-4).

Table 22-4 Equivalent Units for Converting Between a Three-Dimensional Heart and a Linear Muscle Fiber


His diagram for diastole (Fig. 22-11B, purple curve), which shows a rising pressure (tension) with increased EDV (fiber length), is very similar to the early part of the passive length-tension diagram for cardiac muscle (Fig. 22-11A, violet curve). His diagram for systole (Fig. 22-11B, red curve) is more or less equivalent to the ascending phase of the active length-tension diagram for cardiac muscle (Fig. 22-11A, brown curve). Therefore, Starling’s systole curve shows that the heart is able to generate more pressure (i.e., deliver more blood) when more is presented to it.

ventricular performance curve (Fig. 22-11C) is another representation of Starling’s length-tension diagram, but it is one a clinician can obtain on a patient. A ventricular performance curve shows, on the y-axis, stroke work (P• ΔV, see Equation 22-4), which includes Starling’s systolic pressure (itself an estimate of muscle tension), plotted against left atrial pressure, which corresponds to Starling’s end-diastolic volume (itself an estimate of muscle length). What we learn from performance curves obtained on living subjects is that Starling’s law is not a fixed relationship. For instance, the norepinephrine released during sympathetic stimulation—which increases myocardial contractility (as we will see later in this chapter)—steepens the performance curve and shifts it upward and to the left (Fig. 22-11C, brown arrow). Similar shifts occur with other positive inotropic agents (e.g., cardiac glycosides), that is, drugs that increase myocardial contractility. Note also that ventricular performance curves show no descending component because sarcomere length does not increase beyond 2.2 to 2.4 μm in healthy hearts.

The velocity of cardiac muscle shortening falls when the contraction occurs against a greater opposing force (or pressure) or at a shorter muscle length (or lower volume)

The functional properties of cardiac muscle—how much tension it can develop, how rapidly it can contract—depend on many factors but especially on two properties intrinsic to the cardiac myocyte.

1. Initial sarcomere length. For the beating heart, a convenient index of initial sarcomere length is end-diastolic volume. Both initial sarcomere length and EDV are measures of the preload imposed on the cardiac muscle just beforeit ejects blood from the ventricle during systole. Starling’s law, in which the independent variable is EDV, focuses on preload.

2. Force that the contracting myocytes must overcome. In the beating heart, a convenient index of opposing force is the arterial pressure that opposes the outflow of blood from the ventricle. Both opposing force and arterial pressure are measures of the afterload the ventricular muscle must overcome as it ejects blood during systole. Experiments on isotonic contractions focus on the afterload, factors that the ventricle can sense only after the contraction has begun.

Figure 22-12 shows how one might measure the velocity of shortening in a way that is relevant for a cardiac muscle facing both a preload and an afterload. In Figure 22-12A, the muscle starts off at rest, stretched between a fixed support (bottom of the muscle) and the left end of a lever (top of the muscle). A weight attached to the other end of the lever, but resting on a table, applies stretch to the muscle—to the extent allowed by the screw, which adjusts the “stop” of the lever’s left end. Thus, the combination of the weight and the screw determines initial sarcomere length (i.e., preload). At this time, the muscle cannot sense the full extent of the weight. The more we stretch the muscle by retracting the screw, the greater the preload. When we begin to stimulate the muscle, it develops a gradually increasing tension (Fig. 22-12C, lower blue curve), but the length between the fixed point and the left end of the lever (Fig. 22-12A) remains constant. That is, the muscle cannot shorten. Therefore, in the first phase of the experiment, the muscle exerts increasing isometric tension.


Figure 22-12 Effect of preload and afterload on velocity of shortening. In A, the developed tension is not yet sufficient to lift the weight (i.e., afterload). In B, the muscle, which has now developed sufficient tension to lift the weight, shortens against a constant afterload. In C, the slope of the blue curve (ΔLt) is the velocity of shortening. The velocities of shortening for three different afterloads (tensions) in D are plotted as the three colored points of the lower curve in E. In F, the x-axis has the longest lengths on the left, so that “time” runs from left to right (arrows). Note that the family of curves is enclosed by the envelope created by the curve for the greatest initial length.

When the muscle has built up enough tension, it can now begin lifting the weight off the table (Fig. 22-12B). This phase of the contraction is termed afterloaded shortening. The tension now remains at a fixed afterload value (flat portion of lower curve in Fig. 22-12C), but the muscle gradually shortens (rising portion of upper blue curve). Therefore, in the second phase of the experiment, the muscle exerts isotonic contraction. From the slope of the upper curve in Figure 22-12C, we can compute the velocity of shortening at a particular afterload.

This experiment roughly mimics the actions of ventricular muscle during systole. Initially, during its isometric contraction, our hypothetical muscle increases its tension at constant length, as during the isovolumetric contraction of the cardiac cycle shown in segment CD of Figure 22-9. The initial length corresponds to EDV, the preload. Later, the muscle shortens while overcoming a constant force (i.e., generating a constant tension), as during the ejection phase of the cardiac cycle shown in segment DEF of Figure 22-9. The tension corresponds to arterial pressure, the afterload.

What happens if we vary the afterload (i.e., change the weight)? As we already observed in our discussion of skeletal muscle, it is easier to lift a feather than a barbell. Thus, with a heavier weight, the muscle develops a lot of tension but shortens slowly (Fig. 22-12D, red tracings). Conversely, with a lighter weight, the muscle develops only a little bit of tension but shortens rapidly (purple curves).

If we plot the velocities of shortening in Figure 22-12D as a function of the three different afterloads being lifted, we obtain the purple, blue, and red points on the load-velocity curve in Figure 22-12E. The velocity of muscle shortening corresponds to outflow velocity of the ventricle (Fig. 22-8D, E). Thus, at higher opposing arterial pressures, the outflow velocity should decrease. The black curve in Figure 22-12E applies to a muscle that we stretched only slightly in the preload phase (i.e., low preload in Fig. 22-12A). The red curve in Figure 22-12E shows a similar load-velocity relationship for a muscle that we stretched greatly in the preload phase (i.e., high preload). In both cases, the velocity of shortening increases as the tension (i.e., afterload) falls.

When the afterload is so large that no shortening ever occurs, that afterload is the isometric tension, shown as the point of zero velocity on the x-axis of Figure 22-12E. As expected from Starling’s law, the greater the initial stretch (i.e., preload), the greater the isometric tension. In fact, at any velocity (Fig. 22-12E, dashed horizontal line), the tension is greater in the muscle that was stretched more in the preload phase (red curve)—a restatement of Starling’s law.

In summary, at a given preload (i.e., walking up the black curve in Fig. 22-12E), the velocity of shortening for cardiac muscle becomes greater with lower afterloads (i.e., opposing pressure). Conversely, at a given afterload—that is, comparing the black and red curves for any common x value (Fig. 22-12E, dashed vertical line)—the velocity of shortening for cardiac muscle becomes greater with a greater preload (i.e., sarcomere length).

Finally, the curves in Figure 22-12E do not represent a fixed set of relationships. Positive inotropic agents shift all curves up and to the right. Thus, a positive inotropic agent allows the heart to achieve a given velocity against a greater load or to push a given load with a greater velocity.

Another way of representing how velocity of shortening depends on the initial muscle length (i.e., preload) is to monitor velocity of shortening during a single isotonic contraction. If we first apply a large preload to stretch a piece of muscle to an initial length of 9.0 mm (Fig. 22-12F) and then stimulate it, the velocity instantly rises to a peak value of ~8.5 mm/s; it then gradually falls to zero as the muscle shortens to 7.5 mm. If we start by applying a smaller preload, thereby stretching the muscle to an initial length of 8.5 or 8.0 mm, the peak velocity falls. Thus, initial length determines not only the tension that cardiac muscle can generate but also the speed with which the muscle can shorten.

Increases in heart rate enhance myocardial tension

Heart muscle tension has a special dependence on the frequency of contraction. If we stimulate isolated heart muscle only a few times per minute, the tension developed is much smaller than if we stimulate it at a physiological rate of 70/min. The progressive rise of tension after an increase in rate—the positive staircase phenomenon—was first observed by Henry Bowditch in 1871. Underlying the staircase phenomenon is an increase in SR Ca2+ content and release. The larger SR Ca2+ content has three causes. First, during each action potential plateau, more Ca2+ enters the cell through Cav1.2 L-type Ca2+channels, and the larger number of action potentials per minute provides a longer aggregate period of Ca2+ entry through these channels. Second, the depolarization during the plateau of an action potential causes the Na-Ca exchanger NCX1 to operate in the reverse mode, allowing Ca2+ to enter the cell. At higher heart rates, these depolarizations occur more frequently and are accompanied by an increase in [Na+]i, which accentuates the reversal of NCX1, both of which enhance Ca2+ uptake. Third, the increased heart rate stimulates SERCA2a, thereby sequestering in the SR the Ca2+ that entered the cell because of the first two mechanisms. The mechanism of this stimulation is that the rising [Ca2+]i, through calmodulin, activates CaM kinase II, leading to phosphorylation of PLN, enhancing SERCA2a. (See Note: Cardiac Currents Carried by Electrogenic Transporters)

Contractility is an intrinsic measure of cardiac performance

Now that we know that the performance of the heart depends on such factors as degree of filling (i.e., preload), arterial pressure (i.e., afterload), and heart rate, it would be useful to have a measure of the heart’s intrinsic contractile performance, independent of these extrinsic factors. Contractility is such a measure.

Contractility is a somewhat vague but clinically useful term that distinguishes a better performing heart from a poorly performing one. In a patient, it is difficult to assess cardiac performance by use of the approaches in Figures 22-11 and 22-12. One clinically useful measurement of contractility is the ejection fraction (see earlier). However, according to Starling’s law, ejection depends on end-diastolic volume (i.e., preload), which is external to the heart. Two somewhat better gauges of contractility are the rate of pressure development during ejection (ΔPt) and the velocity of ejection. Both correlate well with the velocity of shortening in Figure 22-12E and F, and they are very sensitive guides to the effect of inotropic interventions. (See Note: Contractility in Patients)

A third assessment of contractility focuses on the physiological relationship between pressure and volume during the cardiac cycle. In the era of echocardiography, these volume data are now reasonably easy to obtain. We return to the ventricular pressure-volume loop that we introduced in Figure 22-9 and redraw it as the purple loop in Figure 22-13A. In this example, the end-diastolic volume is 120 mL. Point D′on the loop represents the relationship between pressure and volume at the end of the isovolumetric contraction, when the aortic valve opens. If we had prevented the aortic valve from opening, ventricular pressure would have continued to rise until the ventricle could generate no additional tension. In this case, the pressure would rise to point G′, the theoretical maximum isovolumetric pressure. We could repeat the measurement at very different EDVs by decreasing or increasing the venous return. Point G represents the maximum isovolumetric pressure for an EDV below 120 mL (orange loop), and point G′ for an EDV above 120 mL (green loop). The gold dashed line through points G, G′, and G′ in Figure 22-13Awould describe the relationship between pressure and EDV under isometric (i.e., aortic valve closed) conditions—the equivalent of an isometric Starling curve (e.g., brown curve in Fig. 22-11A). The steeper this line, the greater the contractility.


Figure 22-13 Assessment of contractility by use of a ventricular pressure-volume loop. The purple pressure-volume loop is the normal curve in Figure 22-9. In A, at the same normal state of cardiac contractility, the red loop is generated by decreasing EDV, and the green loop is generated by increasing EDV. The slope of the line through the points at the end of systole (F, F’, and F”) represents the end-systolic pressure-volume relation (ESPVR).

It is impossible to measure maximum isovolumetric pressures in a patient because it is hardly advisable to prevent the aortic valve from opening. However, we can use the end-systolic pressure at point F′, on the normal pressure-volume loop with an EDV of 120 mL (purple loop). For an EDV below 120 mL (orange loop), the corner point would slide down and to the left (F). Conversely, for an EDV above 120 mL (green loop), the corner point would slide upward and to the right (F”). The corner points of many such pressure-volume loops fall along a line—the end-systolic pressure-volume relation (ESPVR)—that is very similar to that generated by the points G, G′, and G”. (See Note: Using ESPVR in Lieu of an Isometric Starling Curve)

Effect of Changes in Contractility The ESPVR is a clinically useful measure of contractility. Enhancing the contractility increases the slope of the ESPVR line, just as it increases the steepness of the ventricular performance curves (Fig. 22-11C, brown arrow). For example, imagine that—with the same EDV and aortic pressure as in the control situation (purple area and gold ESPVR line in Fig. 22-13B)—we increase contractility. We represent increased contractility by steepening the ESPVR line (from gold to red dashed line in Fig. 22-13B). The result is that ejection continues from point D′ to a new point F (red loop in Fig. 22-13B) until the left ventricular volume reaches a much lower value than normal. In other words, enhanced contractility increases stroke volume. Decreasing contractility would flatten the slope of the ESPVR and decrease stroke volume.

Effect of Changes in Preload (i.e., Initial Sarcomere Length) A pressure-volume loop nicely illustrates the effect of an increased preload (i.e., increased filling or EDV) without changing contractility. Starting from the control situation (Fig. 22-13C, purple area), increase of the EDV shifts the isovolumetric segment to the right (CD on the red loop). Because the volume change along segment DEF is larger than for the control situation, stroke volume increases—as predicted by Starling’s law.

Effect of Changes in Afterload A pressure-volume loop also illustrates the effect of an increased afterload (i.e., increase in aortic pressure). Starting from the control situation (Fig. 22-13D, purple area), increase of the aortic pressure shifts the upper right corner of the loop from point D′ (purple loop) to D (red loop) because the ventricle cannot open the aortic valve until ventricular pressure reaches the higher aortic pressure. During the ejection phase—assuming that contractility (i.e., slope of the ESPVR) does not change—the ventricle necessarily ejects less blood until segment DEF intersects the ESPVR line. Therefore, an increase in afterload (at constant contractility) causes the loop to be taller and narrower, so that stroke volume and ejection fraction both decrease. However, if we were to increase contractility (i.e., increase the slope of the ESPVR), we could return the stroke volume to normal.

Positive inotropic agents increase myocardial contractility by raising [Ca2+]i

Modifiers of contractility can affect the dynamics of cardiac muscle contraction, independent of preload or afterload. These factors have in common their ability to change [Ca2+]i. When these factors increase myocardial contractility, they are called positive inotropic agents. When they decrease myocardial contractility, they are called negative inotropic agents.

Positive Inotropic Agents Factors that increase myocardial contractility increase [Ca2+]i, either by opening Ca2+ channels, inhibiting Na-Ca exchange, or by inhibiting the Ca2+ pump—all at the plasma membrane.

Cardiac Hypertrophy

Either volume overload or pressure overload can mechanically compromise the heart. A volume overload is an excessive EDV (i.e., preload). For example, a large arteriovenous shunt would volume overload both the left and right sides of the heart. The increased EDV leads to an increase in stroke volume (Fig. 22-13C), which elevates cardiac output. Systemic arterial pressure usually remains normal. A pressure overload is an excessive pressure in the ventricle’s outflow tract (i.e., afterload). For the left side of the heart, the problem would be an increase in systemic arterial pressure (i.e., hypertension). The increased aortic pressure leads to a decrease in stroke volume (Fig. 22-13D). However, because of a compensatory increase in heart rate, cardiac output usually remains normal. When, over time, the adaptive process of hypertrophy becomes inadequate to cope with demand, the result is mechanical dysfunction and, ultimately, heart failure (see the box titled Cellular Basis of Heart Failure).

Because cells of the adult heart are terminally differentiated, stimuli that might be mitogenic in other cells cannot elicit cell division in the heart but rather cause the cardiac myocytes to hypertrophy and increase muscle mass. Elite athletes develop physiological hypertrophy, whereby the cardiac cells increase proportionally both in length and in width. Volume overload leads to eccentric hypertrophy characterized by increases of myocyte length out of proportion to width. Pressure overload causes concentric hypertrophy with a relatively greater increase in myocyte width.

A host of events may trigger hypertrophy, including various hypertrophic factors, increases in [Ca2+]i, and mechanical forces.

Hypertrophic Factors

Agents implicated in cardiac hypertrophy include the cardiac cytosolic protein myotrophin (Myo/V1) and the cytokine cardiotrophin 1 (CT-1) as well as catecholamines, angiotensin II, endothelin 1, insulin-like growth factor 2, transforming growth factor β, and interleukin 1. Catecholamines and angiotensin II both activate the MAP kinase cascade. Farther downstream the signal transduction pathway, the transcriptional response to hypertrophic stimuli includes the zinc finger transcription factor GATA4 and perhaps also the transcription factors SRF and Sp1 as well as the TEF-1 family. (See Chapter 4.)


Elevated [Ca2+]i may be both a trigger for hypertrophy and part of signal transduction pathways that lead to hypertrophy. [Ca2+]i in heart cells is probably elevated initially during chronic volume or pressure overloads, just as [Ca2+]i would be elevated in a normal heart that is working hard. Elevated [Ca2+]imay activate calcineurin, a Ca2+-dependent phosphatase (see Chapter 3). After being dephosphorylated by calcineurin, the transcription factor NF-AT3 can enter the nucleus and bind to GATA4 (see earlier), which transcriptionally activates genes responsible for hypertrophy. Mice that express constitutively activated forms of calcineurin develop cardiac hypertrophy and heart failure.

Mechanical Factors

Mechanical stretch induces the expression of specific genes. The mechanical sensor that triggers cardiac hypertrophy may be MLP (muscle LIM protein), part of the myocardial cytoskeleton. Stretch activates a phosphorylation cascade of protein kinases: Raf-1 kinase, extracellular signal–regulated kinase (ERK), and a separate subfamily of the MAP kinases called SAPKs (for s tretch-a ctivated p rotein k inases). These various kinases regulate gene expression by activating the transcription factor AP-1 (see Chapter 4).

The pathways we have just discussed lead to several changes in gene expression within cardiac myocytes during hypertrophy. In addition to synthesizing many housekeeping proteins, hypertrophic cardiac myocytes undergo other changes that are more specific for contraction. Some of the most striking changes include reduced levels of the mRNA encoding three critical proteins in the membrane of the SR: (1) the Ca2+ release channel, (2) phospholamban, and (3) the SR Ca2+ pump (SERCA2). In addition, cardiac hypertrophy is associated with increased levels of mRNA for the skeletal α-actin, which is normally expressed in fetal but not in adult heart. Hypertrophic hearts also have increased expression of the angiogenic factor VEGF (see Chapter 20).

Although a hypertrophied myocardial cell may be able to do more work than a nonhypertrophied cell, it has a lower “contractility” when normalized to its cross-sectional area. Why should hypertrophied cardiac muscle not be as good as normal muscle? Possibilities include alterations in the transient increases in [Ca2+]i during the cardiac action potential and alterations in the expression of the contractile filaments, particularly the myosin isoenzymes.

1. Adrenergic agonists. Catecholamines (e.g., epinephrine, norepinephrine) act on β1 adrenoceptors to activate the α subunit of Gs-type heterotrimeric G proteins. The activated αs subunits produce effects by two pathways. First, αsraises intracellular levels of cyclic adenosine monophosphate (cAMP) and stimulates protein kinase A (see Chapter 3), which can then act by the mechanisms summarized in Table 22-2to increase contractility and speed relaxation. Second, αs can directly open L-type Ca2+ channels in the plasma membrane, leading to an increased Ca2+ influx during action potentials, increased [Ca2+]i, and enhanced contractility.

2. Cardiac glycosides. Digitalis derivatives inhibit the Na-K pump on the plasma membrane (see Chapter 5) and therefore raise [Na+]i. We would expect the increased [Na+]i to slow down the Na-Ca exchanger NCX1, to raise steady-state [Ca2+]i, and to enhance contractility. Recent evidence suggests that cardiac glycosides may also increase [Ca2+]i by a novel pathway—increasing the Ca2+ permeability of Na+channels in the plasma membrane.

3. High extracellular [Ca2+]. Acting in two ways, elevated [Ca2+]o increases [Ca2+]i and thereby enhances contractility. First, it decreases the exchange of external Na+ for internal Ca2+. Second, more Ca2+enters the myocardial cell through L-type Ca2+ channels during the action potential.

4. Low extracellular [Na+]. Reducing the Na+ gradient decreases Ca2+ extrusion through NCX1, raising [Ca2+]i and enhancing contractility.

5. Increased heart rate. As we noted in introducing the staircase phenomenon, an increased heart rate increases SR stores of Ca2+ and also increases Ca2+ influx during the action potential.

Cellular Basis of Heart Failure

Heart failure is among the most common causes of hospitalization in developed countries for people aged 65 years or older and is a leading cause of death. People whose hearts cannot sustain an adequate cardiac output become breathless (because blood backs up from the left side of the heart into the lungs) and have swollen feet and ankles (because blood backs up from the right side of the heart and promotes net filtration in systemic capillaries; see the box on edema in Chapter 20). On the cellular level, decreased contractility in heart failure could be a result of cardiac hypertrophy (see the box titled Cardiac Hypertrophy), reflecting alterations in the transient increases of [Ca2+]i, the expression of the contractile filaments, or both.

Changes in [Ca2+]i physiology could reflect altered properties of the Cav1.2 L-type Ca2+ channel in the plasma membrane or the Ca2+ release channel RYR2 in the SR membrane. In an animal model of hypertension-induced cardiac hypertrophy that leads to heart failure, the Cav1.2 channels exhibit an impaired ability to activate RYR2 through Ca2+-induced Ca2+ release. A distortion of the microarchitecture in hypertrophic cells, and thus a distortion of the spacing between Cav1.2 channels and RYR2, could be responsible for impaired coupling. Each of the four RYR2 molecules in the Ca2+ release channel associates with a molecule of calstabin 2 (also known as the FK506-binding protein FKBP12.6) that, together with other proteins, forms a macromolecular complex regulating the Ca2+ release channel. Depletion of calstabin 2 in heart failure results in leaky RYR2 channels that continually release Ca2+ into the cytosol. High [Ca2+]i makes the heart prone to delayed afterdepolarizations (see Chapter 21), ventricular arrhythmias, and sudden death. (See Note: Triggered Activity)

Changes in the expression of contractile proteins can reduce contractility. Two isoforms of myosin heavy chain, αMHC and βMHC, are present in heart (see Table 9-1). The speed of muscle shortening increases with the relative expression of αMHC. In human heart failure, the amount of αMHC mRNA, relative to total MHC mRNA, falls from ~35% to ~2%.

An interesting animal model of heart failure is the knockout mouse that lacks the gene encoding MLP, the muscle LIM protein (see the box on cardiac hypertrophy). MLP-deficient mice have the same disrupted cytoskeletal architecture seen in failing hearts. In addition, these mice have a dilated cardiomyopathy. Although humans with failing hearts are generally not deficient in MLP, the evidence from these knockout mice suggests that the MLP system could play a role in certain forms of cardiomyopathy.

Negative Inotropic Agents Factors that decrease myocardial contractility all decrease [Ca2+]i.

1. Ca2+ channel blockers. Inhibitors of L-type Ca2+ channels (see Chapter 7)—such as verapamil, diltiazem, and nifedipine—reduce Ca2+ entry during the plateau of the cardiac action potential. By reducing [Ca2+]i, they decrease contractility.

2. Low extracellular [Ca2+]. Depressed [Ca2+]o lowers [Ca2+]i, both by increasing Ca2+ extrusion through NCX1 and by reducing Ca2+ entry through L-type Ca2+ channels during the plateau of the cardiac action potential.

3. High extracellular [Na+]. Elevated [Na+]o increases Ca2+ extrusion through NCX1, thereby decreasing [Ca2+]i.


Books and Reviews

Kobayashi T, Solaro RJ: Calcium, thin filaments, and the integrative biology of cardiac contractility. Annu Rev Physiol 2005; 67:39-46.

Moss RL, Fitzsimons DP: Frank-Starling relationship: Long on importance, short on mechanism. Circ Res 2002; 90:11-13.

Simmerman HKB, Jones LR: Phospholamban: Protein structure, mechanism of action, and role in cardiac function. Physiol Rev 1998; 78:921-947.

Wehrens XH, Marks AR: Molecular determinants of altered contractility in heart failure. Ann Med 2004; 36(Suppl 1):70-80.

Zimmer H-G: The isolated perfused heart and its pioneers. News Physiol Sci 1998; 13:203-210.

Journal Articles

Collins SP, Arand P, Lindsell CJ, et al: Prevalence of the third and fourth heart sound in asymptomatic adults. Congest Heart Fail 2005; 11:242-247.

Fukuda N, Wu Y, Nair P, Granzier HL: Phosphorylation of titin modulates passive stiffness of cardiac muscle in a titin isoform–dependent manner. J Gen Physiol 2005; 125:257-271.

Gómez AM, Valdivia HH, Cheng H, et al: Defective excitation-contraction coupling in experimental cardiac hypertrophy and heart failure. Science 1997; 276:800-806.

Sarnoff JS, Berglund E: Ventricular function. I: Starling’s law of the heart studied by means of simultaneous right and left ventricular function curves in the dog. Circulation 1954; 9:706-718.

Wiggers CJ, Katz LN: The contours of ventricular volume curves under different conditions. Am J Physiol. 1922; 58:439-475.


If you find an error or have any questions, please email us at Thank you!