For all the advantages of studying isolated myocytes and muscle fibers, an integrated and more realistic analysis of cardiovascular function regards the left ventricle (LV) as a muscle pump coupled to the systemic and venous circulations. In contrast to isolated cardiac muscle, contraction of the intact LV is auxotonic, in that force increases and decreases during ejection of viscous blood into a viscoelastic arterial system. Moreover, attempts to extrapolate results from isolated muscle to the intact LV are hampered by the complexity of chamber geometry and myocardial fiber orientation, which make it difficult to estimate initial fiber length (preload) and the force opposing LV ejection (afterload). Finally, unlike isolated cardiac muscle, ventricular performance is modulated by neurohumoral influences, right and LV interaction, restraining effects of the pericardium, and atrial function. At the organ level, the preceding events are initiated by the electrical activation of the heart and structured by the sequence of events in a heartbeat, the cardiac cycle.
The electrocardiogram (ECG) (Fig. 5–11) records the pattern of electrical activation of the heart on the body surface. Electrical currents generated by differences in potential between depolarized and polarized regions of the heart are conducted through the body, detected by electrodes, and amplified and recorded on calibrated moving paper. The ECG provides important clinical information regarding the electrical orientation of the heart in three-dimensional space, the relative size of the cardiac chambers, and the presence of conduction system defects and provides evidence for a variety of underlying pathologic conditions, such as ischemia, infarction, cardiomyopathy, and hypertrophy.
Nomenclature of the deflections, intervals, and segments of the normal electrocardiogram. Reproduced with permission from Sperelakis N, Banks RO: Essentials of Physiology. 2nd ed. Boston: Little, Brown and Company; 1996.
The SA node is the primary pacemaker of the heart and is located at the junction of the superior vena cava and the right atrium. The action of the SA node is electrically silent, although a measurable conduction time between sinus node discharge and atrial depolarization (denoted by a P wave) can be measured on intracardiac electrograms. Action potentials travel rapidly (1.0-1.5 ms) through the atrial myocardium and generate an atrial contraction. Preferential conduction in specialized bundles of muscle fibers (the internodal tracts of Bachmann, Wenckebach, and Thorel) nearly simultaneously activate the atrial musculature and ensure that the action potential reaches the AV node in a timely fashion. Excitation of the ventricles spreads by means of the AV node and the His-Purkinje system (bundle of His and bundle branches). The impulse travels slowly (0.02–0.05 ms) through the AV node. In contrast, conduction velocity through the Purkinje system is very fast (2.0–4.0 ms). The PR interval includes atrial depolarization, AV nodal conduction, and His-Purkinje activity. Activation of ventricular myocardium (conduction velocity 1.0–2.0 ms) occurs after most of the conduction system is depolarized and is represented by the QRS complex. Ventricular repolarization occurs during the T wave.
The ECG is essentially a voltmeter that measures and records potential differences between pairs of electrodes or leads. Three bipolar leads (I, II, II), three unipolar limb leads (aVR, aVL, aVF), and six precordial leads (V1-V6) record the distribution of the potentials on the frontal and horizontal planes of the heart (see Fig. 5–11). Depolarization and repolarization of the heart results in differences in electrical potential, and the ECG measures these changes in potential over time. The external surface of a depolarized membrane becomes electrically negative relative to quiescent, polarized areas. The direction of the propagated impulse travels from the depolarized to polarized areas. By convention, the direction of the propagation wave toward the positive pole of a bipolar lead system or exploring electrode produces an upright defection and conversely, if the propagation wave is toward the negative pole or away from an exploring electrode, a negative deflection is produced. Depolarization progresses from cell to cell in an orderly fashion from endocardium to epicardium from the apex to base of the heart. In contrast, repolarization does not occur as a propagated wave; nevertheless, it is represented by a single vector that integrates multiple areas of potential difference. Local circuit currents precede the depolarization wavefront, depolarize the adjacent membrane, and bring the membrane to threshold potential; with depolarization, the local circuit currents flow through low-resistance gap junctions (the major component of which is connexin) and depolarize a neighboring cell. Thus, the myocardium functions as a functional syncytium. The ECG is discussed in detail in Chap. 12.
The cardiac cycle describes pressure, volume, and flow phenomena in the ventricles as a function of time. This cycle is similar for both the LV and right ventricle (RV), although there are differences in timing stemming from differences in the depolarization sequence and the levels of pressure in the pulmonary and systemic circulations. For simplicity the cardiac cycle for the left heart during one beat will be described (Fig. 5–12).
Pressure, flow, volume, electrocardiographic and phonocardiographic events constituting the cardiac cycle. Reproduced with permission from Berne RM, Levy MN: Physiology, 2nd ed. St. Louis: Mosby; 1988.
The QRS complex on the surface ECG represents ventricular depolarization. Contraction (systole) begins after an approximately 50-ms delay and results in closure of the mitral valve. The LV contracts isovolumetrically until the ventricular pressure exceeds the systemic pressure; at this time, the aortic valve opens and ventricular ejection occurs. Bulging of the mitral valve into the left atrium during isovolumic contraction causes a slight increase in left atrial pressure (c wave). Shortly after ejection begins, the active state declines, and ventricular pressure begins to decrease. Left atrial pressure rises during ventricular systole (v wave) as blood returns to the left atrium by means of the pulmonary veins. The aortic valve closes when LV pressure falls below aortic pressure; momentum briefly maintains forward flow despite greater aortic than LV pressure. Ventricular pressure then declines exponentially during isovolumic relaxation when both the aortic and mitral valves are closed. This begins ventricular diastole. When ventricular pressure declines below left atrial pressure, the mitral valve opens, and ventricular filling begins. Initially, ventricular filling is very rapid because of the relatively large pressure gradient between the atrium and ventricle. Ventricular pressure continues to decrease after mitral valve opening because of continued ventricular relaxation; its subsequent increase (and the decrease in atrial pressure) slows ventricular filling. Especially at low end-systolic volumes, ventricular early rapid filling can be facilitated by ventricular suction produced by elastic recoil. Ventricular filling slows during diastasis when atrial and ventricular pressures and volumes increase very gradually. Atrial depolarization is followed by atrial contraction; increased atrial pressure (a wave); and a second, late rapid-filling phase. A subsequent ventricular depolarization completes the cycle.
Valve closure and rapid-filling phases are audible with a stethoscope placed on the chest and can be recorded phonocardiographically after electronic amplification. The first heart sound, resulting from cardiohemic vibrations with closure of the AV (mitral, tricuspid) valves, heralds ventricular systole. The second heart sound, which is shorter and composed of higher frequencies than the first, is associated with closure of the semilunar valves (aortic and pulmonic) at the end of ventricular ejection. Third and fourth heart sounds are low-frequency vibrations caused by early, rapid filling and late diastolic atrial contractile filling, respectively. These sounds can be heard in normal children, but in adults usually indicate disease.
An alternative time-independent representation of the cardiac cycle is obtained by plotting instantaneous ventricular pressure and volume (Fig. 5–13). During ventricular filling, pressure and volume increase nonlinearly (phase I). The instantaneous slope of the P–V curve during filling (dP/dV) is diastolic stiffness, and its inverse (dV/dP) is compliance. Thus, as chamber volume increases, the ventricle becomes stiffer. In a normal ventricle, operative compliance is high because the ventricle operates on the flat portion of its diastolic P–V curve. During isovolumic contraction (phase II), pressure increases and volume remains constant. During ejection (phase III), pressure rises and falls until the minimum ventricular size is attained. The maximum ratio of pressure to volume (maximal active chamber stiffness or elastance) usually occurs at the end of ejection. Isovolumic relaxation follows (phase IV), and when LV pressure falls below left atrial pressure, ventricular filling begins. Thus, end-diastole is at the lower right hand corner of the loop, and end-systole is at the upper left corner of the loop. LV P–V diagrams illustrate the effects of changing preload, afterload, and inotropic state in the intact ventricle (see the following).
A. Left ventricular pressure–volume (P–V) loop, the segments of which correspond to events of the cardiac cycle: diastolic ventricular filling along the passive P–V curve (phase I), isovolumetric contraction (phase II), ventricular ejection (phase III), and isovolumetric relaxation (phase IV). B. The ventricle ejects to an end-systolic volume determined by the peak isovolumetric P–V line; an isovolumetric contraction (large arrowheads) from varying end-diastolic volumes (preload). Reproduced with permission from Sperelakis N, Banks RO: Essentials of Physiology. 2nd ed. Boston: Little, Brown and Company; 1996.
A P–V loop can also be described for atrial events. During ventricular ejection, descent of the ventricular base lowers atrial pressure and thus assists in atrial filling. Filling of the atria from the veins results in a v wave on the atrial and venous pressure tracing. When the mitral and tricuspid valves open, blood stored in the atria empties into the ventricles. Atrial contraction, denoted by an a wave on the atrial pressure tracing, actively assists ventricular filling. The resultant atrial P–V diagram has a figure-of-eight configuration with a clockwise V loop, representing passive filling and emptying of the atria, and a counterclockwise A loop, representing active atrial contraction. Thus, the atria function as a reservoir, a conduit for venous flow (during ventricular systole and diastole, respectively), and a booster pump for ventricular filling late in diastole.39
Pressure–Volume Relationships in the Isolated Heart
Isolated, perfused, isovolumically contracting hearts are useful preparations to study preload dependency of ventricular performance (the Frank-Starling relation) and fully relaxed, end-diastolic P–V relationships without the confounding, uncontrolled changes in either neurohumoral activation or coronary perfusion. These preparations are especially well-suited for quantifying end-systolic elastance (stiffness), a relatively load-independent index of ventricular function. The time-varying elastance model of ventricular contraction is based on the experimental observations in which ventricular volume and loading are altered under conditions of unvarying contractility (Fig. 5–14). At any time, t, following the onset of contraction, the relation between pressure (P) and volume (V) is linear according to the relation: P(t) = E(t) – [V(t) – V0], where E is the time-varying elastance and V0 is the volume at zero pressure or dead volume; this relation becomes progressively steeper until it reaches a maximum at end-systole. Thus, the ventricle behaves like a spring with a stiffness (elastance) that increases during contraction and decreases during relaxation. The slope of the end-systolic P–V relationship, end-systolic elastance (Ees) changes directly as a function of acute changes in contractility without a change in dead volume (V0). Appropriate changes in Ees are also observed with increases in beating frequency (eg, force–frequency relationships).
The time-varying elastance concept. A. A series of variably loaded pressure–volume areas (PVAs). The relationship between pressure and volume at any time, t, during variably loaded contractions under constant contractility conditions is linear and reaches a maximum value at end-systole. Filled circles connected by straight lines occur at the same time, t, during contraction. Emax is the line connecting the points at end-systole (ES). B. Elastance, E (t) increases at each time, t, during contraction until it reaches maximal values at ES. Increasing contractility increases the slope at any time, t, including ES (Emax). C,D. The concept that the ventricle behaves as an elastic spring with a stiffness (elastance) that increases during systole and decreases during diastole (EES or Emax). D, diastole; P, pressure; S, systole; V, volume; V0, dead volume. Reproduced with permission from Fuster V, Alexander RW, O’Rourke RA, et al: Hurst’s The Heart. 11th ed. New York: McGraw-Hill; 2004.
The elastance concept has been extended to the study of ventricular mechanoenergetics by proposing that the P–V area (PVA) bounded by the LV P–V loop is a measure of the total mechanical energy of LV contraction.40,41 The PVA concept is shown schematically in Fig. 5–15. The total mechanical energy of contraction can be considered to consist of two components: (1) external work, the area enclosed within the P–V loop; and (2) potential energy stored in the ventricular spring at ES: that is, the area between the end-systolic pressure relation on the left and the end-diastolic P–V relation on the right.
A. Schematic of the VO2–PVA (pressure–volume area) concept. In ejecting contractions, PVA = EW + PE; in isovolumic contraction, PVA = PE only. B. Correlation of PVA with VO2. P–V, pressure– volume; VO2, oxygen consumption. Reproduced with permission from Fuster V, Alexander RW, O’Rourke RA, et al: Hurst’s The Heart. 11th ed. New York: McGraw-Hill; 2004.
The myocardial oxygen consumption (MVO2)–PVA relationship is obtained by measuring P–V area loops and LV MVO2 at several steady-states. There is a highly linear correlation (r > .98) between LV VO2/beat and PVA/beat over a wide range of experimental conditions (see Fig. 5–15, bottom), indicating the accuracy of the PVA as a measure of total mechanical energy. The VO2 intercept of the VO2–PVA relationship is the unloaded VO2 (PVA–independent VO2), which in an isovolumically contracting heart, corresponds to a point at which LV peak pressure is 0 mm Hg (Fig. 5–16). At this point, except for a low level of cross-bridge cycling caused by shape changes, there is neither mechanical energy produced nor energy expended for cross-bridge cycling.41 The VO2 under unloaded conditions reflects energy used for E–C coupling and basal metabolism; the latter can be eliminated experimentally by arresting the heart. In this manner, changes in E–C coupling energy consumption have been detected as shifts in the unloaded VO2. Oxygen consumption used by the contractile apparatus for cross-bridge cycling is PVA-dependent VO2, which increases linearly and directly with PVA. Because PVA-dependent VO2 is the energy input and the PVA is the total energy output of the contractile machinery, the inverse slope of the VO2–PVA relationship is a dimensionless measure of the thermodynamic efficiency of the contractile machinery. Unlike efficiency expressed as the external work/total VO2, efficiency expressed by the VO2–PVA relationship is relatively insensitive to load. The VO2–PVA relationship is sensitive to metabolic changes and impacts the efficiency of ATP production.
Schematic of the relationship between VO2/beat and the pressure–volume area (PVA). The diagonal VO2–PVA line is obtained by mechanical unloading of the ventricle; the y intercept is the unloaded VO2. Mechanically unloaded VO2 is subdivided into VO2 for residual cross-bridge cycling and that for nonmechanical VO2 (BDM, which inhibits cross-bridge cycling). The latter consists of VO2 for excitation–contraction (E–C) coupling and basal metabolism. P–V, pressure–volume; VO2, oxygen consumption. Reproduced with permission from LeWinter MM, Suga H, Watkins MW: Cardiac Energetics: From Emax to Pressure-Volume Area. Norwell: Kluwer Academic Publishers; 1995.
Determinants of Left Ventricular Function
Measures of Ventricular Performance
Measures of overall ventricular performance typically include cardiac output (the quantity of blood delivered to the circulation, calculated as the stroke volume and heart rate), stroke volume (quantity of blood ejected/beat, which equals the ventricular end-diastolic volume minus the end-systolic volume), and stroke work (the product of pressure and stroke volume, which equals the area bounded by the ventricular PVA and which can be approximated in the clinical setting as ([Mean LV systolic – Diastolic pressure] × Stroke volume × 0.0136). Cardiac output responds to changes in the oxygen requirements of tissues, for example, as occurs with exercise. The extraction of nutrients by tissue can be expressed as the arteriovenous difference across the tissue. According to the Fick principle, the consumption of a particular nutrient (eg, oxygen) by a tissue equals the rate of delivery of that nutrient: that is, the cardiac output times the arteriovenous difference of that nutrient. Changes in cardiac output necessary to meet the metabolic needs of the tissues can be produced by changes in the stroke volume, heart rate, or both. Changes in stroke volume are mediated by altered loading conditions, inotropic state, and heart rate. Thus, factors that influence the strength of contraction in isolated muscle are the same factors that determine cardiac output. The stroke volume expressed as a function of the end-diastolic volume is the ejection fraction (EF). Thus, EF = (End-diastolic volume – End-systolic volume)/End-diastolic volume.
The influence of preload on measures of ventricular performance defines the LV function curve, known as the Frank-Starling curve.42,43 Increasing LV end-diastolic volume increases stroke volume in ejecting beats and increases peak LV pressure in isovolumic beats. The modulation of ventricular performance by changes in preload, termed heterometric regulation, operates on a beat-by-beat basis and is responsible for matching outputs of the right and LVs, as with changes in posture and breathing. The Frank-Starling curve also represents an important compensatory mechanism that maintains LV stroke volume (vis-à-vis increasing LV end-diastolic volume) when LV shortening is impaired, owing either to myocardial contractile dysfunction or to excessive afterload. The atria also exhibit a Frank-Starling curve that becomes clinically important during exercise and when there is resistance to early diastolic LV filling.
Because a representative fiber length (ie, preload) is difficult to determine in the LV, changes in the myocardial fiber length are estimated from changes in either the LV end-diastolic volume or LV end-diastolic pressure. In the clinical setting, end-diastolic pressure and pulmonary capillary wedge pressure are used frequently as measures of preload. However, the passive P–V relationship, analogous to the passive length–tension curve in isolated muscle, is not linear but exponential. Thus, the ratio of change in LV pressure to volume is greater at higher than at lower LV volumes. Not surprisingly, under certain circumstances, ventricular pressure can inaccurately reflect the ventricular volume. Moreover, changes in ventricular volume can erroneously be inferred from changes in cardiac pressures, which can result only from alterations in ventricular compliance. For example, whereas chronic volume overload can shift the ventricular diastolic pressure relationship rightward so that volume is increased at a normal end-diastolic pressure, chronic pressure overload can shift the diastolic P–V relationship leftward and for the same end-diastolic pressure result in a smaller ventricular volume. Compliance of the LV is affected by pericardial pressure, RV pressure and volume, and coronary artery perfusion (turgor) in addition to changes in the intrinsic elastic properties of the LV.
Afterload in the intact heart can be considered as the tension in the LV wall that resists ventricular ejection (wall stress during systole) or as the arterial input impedance (the ratio of instantaneous change in pressure to instantaneous change in flow). Although forces within the ventricular wall are difficult to measure, initial estimates of systolic wall stress can be derived from application of the Laplace relationship in which wall tension = (P · r)/2 h, where P refers to pressure, r to ventricular radius, and h to wall thickness. More complex derivations based on various geometric assumptions are used to calculate end-systolic wall stress. Input impedance is a complex function of arterial pressures, elasticity, vessel dimension, and blood viscosity, which requires measurement of instantaneous aortic pressure and flow and is therefore impractical to measure in the clinical setting. Because of its simplicity, aortic pressure is often used as a surrogate for afterload. An increase in afterload causes a decrease in stroke volume and the velocity of LV shortening. The resulting stress–shortening and stress–velocity curves are analogous to those obtained from variably afterloaded isotonic contractions in isolated muscle.
The ideal method of measuring the inotropic state in the intact LV should incorporate the variables of force, length, velocity, and time; be independent of external loading conditions; and relate to physicochemical processes at the sarcomeric level. Because of these constraints, changes in inotropic state are usually defined operationally by shifts of the various ventricular function curves, which, by definition, are independent of loading conditions. For example, a drug with positive inotropic activity (eg, dobutamine) shifts the Frank-Starling curve (analogous to the length–shortening curve in papillary muscle preparations) upward and to the left, and changes in the stress–shortening relationship (analogous to the force–velocity curve) upward and to the right.
The rate of pressure development in the LV during isovolumic systole (dP/dt) is used frequently as an index of the inotropic state. Although LV + dP/dtmax provides a measure of the rate of tension development and of myocardial contractility, this index is preload dependent, caused in part by length-dependent changes in myofilament Ca2+ sensitivity. However, LV + dP/dtmax is largely independent of afterload, provided that the maximum rate of increase occurs before aortic valve opening. Although changes in the maximal rate of increase of ventricular pressure are highly sensitive to acute changes in contractility and are useful to assess directional changes in inotropic state, absolute dP/dtmax is not as useful for assessment of basal contractility as are the ejection phase indices, such as LVEF (stroke volume/end-diastolic volume ×100). Furthermore, dP/dtmax cannot be corrected for changes in muscle mass produced by LV hypertrophy, in which case it is best to compare peak stress, which incorporates pressure, volume, mass, and geometry. Because of the direct influence of preload on dP/dt—dP/dt at a common developed pressure (LV systolic minus diastolic pressure) and the slope of the dP/dt end-diastolic volume curve (preload recruitable stroke work; see the following) have been proposed as preload independent indices of the inotropic state.40
End-systolic P–V points from ejecting beats obtained from variably preloaded or afterloaded contraction fall reasonably close to the isovolumetric P–V line for a given inotropic state (vide supra). Thus, changes in the inotropic state, independent of the loading conditions can be identified by changes in the slope of the end-systolic P–V relationship (Ees). By acutely altering loading conditions (eg, transient vena caval occlusions or phenylephrine boluses), a family of PVAs is obtained (single-beat methods designed for clinical use have been proposed). End-systole can be defined as end ejection or as the time of maximal elastance (the maximal P–V ratio) during systole. In the normal heart, these two points are closely related in time. In practice, the end-systolic P–V relationship (ESPVR) is constructed by connecting the end-systolic points of each loop; the relationship is relatively linear and defines the properties of the chamber when maximally activated.40
However, Ees does have a modest degree of load dependence, likely caused by the load dependence of activation. Moreover, the linear ESPVR is really curvilinear, particularly at the extremes of the contractile state. The effects of nonlinearity are particularly important when the P–V relationship is acquired over a narrow range of pressures and volume. A single slope in the latter instance will not uniquely characterize the ESPVR and therefore the contractile state. In addition, the extrapolated V0 is unlikely to represent dead volume. Finally, V0 is not entirely independent of inotropic state. Thus, more than Ees is needed to compare two contractile states; interpretation must take into account V0, and analysis of covariance, or a multiple linear regression analysis with dummy variables is desirable.40
Other considerations for the use of P–V relations to characterize contractility are (1) specialized and invasive instruments are necessary for its measurement; (2) methods used to alter load should be free of inotropic effects; (3) because changes in autonomic tone and heart rate can complicate analysis, loading changes should be as rapid as possible; (4) arrhythmias may occur and complicate the analysis; (5) changes in coronary perfusion pressure that can alter the P–V relationship occur with changes in load; and (6) changes in mass and geometry of the ventricle make changes in the ESPVR ambiguous. In addition to Ees, preload recruitable stroke work (slope of the end-diastolic volume–stroke work relationship) and the slope of the end-diastolic volume–dP/dtmax relationship are derived as indices of contractility from P–V analysis. Each of these approaches is linear and afterload independent. Preload recruitable stroke work is independent of heart size, and the slope of the end-diastolic volume–dP/dtmax is more sensitive to inotropic state than is Ees.40
Heart rate is normally determined by the interplay between the intrinsic automaticity of the SA node and the activity of the autonomic nervous system. Increasing heart rate causes a small but measurable increase in the inotropic state through the force–frequency relationship. In addition, heart rate is a major determinant of cardiac output. However in a normal heart, pacing between heart rates of 60 and 160 beats/min has little effect on cardiac output because the diminished diastolic filling time offsets the modest increase in inotropic state.
Diastole and Diastolic Function
Diastole is the summation of processes by which the heart keeps latent its ability to generate force and shorten, and returns to its precontractile state. Diastolic properties of the ventricle are complex and multifactorially determined and are related to the speed and synchrony of myocardial relaxation and inactivation, loading conditions, viscoelasticity, heart rate, atrial function, and ventricular interaction. Diastole occurs in a series of energy-consuming steps beginning with release of calcium from TnC, detachment of actin–myosin cross-bridges, SERCA2a-induced calcium sequestration into the SR, NCX-induced extrusion of calcium from the cytoplasm, and return of the sarcomere to its resting length. Adequate ATP must be present for these processes to occur at a sufficient rate and extent.
The P–V relationship during early diastole reflects the lusitropic (relaxation) state of the heart, analogous to the inotropic (contraction) state measured during systole. The rate of LV relaxation can be estimated from the maximal rate of pressure decay (–dP/dtmax) and indices (eg, relaxation half-time [RT1/2]) that are related to the time necessary for ventricular relaxation, but these measurements are highly dependent on the prevailing load of the intact circulation. In contrast, τ, the time constant of LV relaxation during isovolumic relaxation, provides a more accurate, less load-dependent measure of relaxation; τ is shortened by β-adrenergic stimulation (cyclase-dependent phosphorylation of phospholamban and TnI) and prolonged with β-adrenergic antagonists.44 Although several mathematical models of the exponential decay of LV pressure exist, a simple monoexponential model that declines to zero is frequently used: P(t) = Poe – t/ τ where P(t) is the LV pressure at any time, t; τ is the relaxation constant; Po is the LV pressure at the onset of relaxation; and e is the base of the natural logarithm. The natural logarithmic transformation of both sides of the equation yields lnP = –1/T + ln Po. Thus, τ is derived by obtaining the negative of the reciprocal of the slope of lnP (t) versus time, t, from aortic valve closure to mitral valve opening (isovolumic relaxation). High-fidelity catheter tip micromanometers are necessary for accurate measurement of –dP/dt max and τ.
In addition to relaxation, the passive viscoelasticity of the ventricle, dependent both on intracellular and extracellular structures, is a major determinant of diastolic function. During contraction, cytoskeletal proteins such as titin and microtubules are deformed by actin–myosin cross-bridge cycling and sarcomere contraction, which act like viscoelastic springs during diastole.45 This reclaimed potential energy constitutes a recoiling force that helps restore the myocardium to its resting configuration. In addition, ECM proteins such as collagen contribute to the establishment of resting force and length.
Chamber stiffness is quantified from the relationship between diastolic LV pressure and volume. LV diastolic pressure can be changed either by a volume-dependent change in operating stiffness (equal to the slope of a tangent drawn to the P–V curve at any point) or by a volume-independent change in the overall chamber stiffness because of a change in properties either intrinsic (eg, hypertrophy) or extrinsic (eg, pericardial) to the ventricle (Fig. 5–17, Table 5–1). Operating stiffness changes throughout filling, such that stiffness (dP/dV) is less at smaller volumes and greater at larger volumes. Because the diastolic P–V relationship is generally exponential, the relationship between dP/dV and pressure is linear. The slope of this relationship is called the modulus of chamber stiffness (kc) and has been used to quantitate chamber stiffness. Thus, when chamber stiffness is increased, the P–V curve shifts to the left, the slope of the dP/dt versus pressure relationship becomes steeper, and kc is increased.
TABLE 5–1.Factors Influencing Left Ventricular Chamber Stiffness |Favorite Table|Download (.pdf) TABLE 5–1. Factors Influencing Left Ventricular Chamber Stiffness
|Physical properties of the LV |
| LV chamber volume and mass |
| Composition of the LV wall |
| Viscosity, stress relaxation, and creep |
|Intrinsic factors |
| Myocardial relaxation |
| Coronary turgor |
|Extrinsic factors |
| Pericardial restraint |
| RV interaction |
| Atrial contraction |
| Pleural and mediastinal pressure |
A. End-diastolic pressure–volume relationship (EDPVR) in two ventricles with differing passive diastolic properties. Chamber stiffness is dP/dV at any point on the EDPVR. The stiffer chamber on the left has a steeper overall slope. B. The same data plotted as pressure versus chamber stiffness. Because of the exponential nature of EDPVR, the relation between chamber stiffness and pressure is a straight line whose slope is the chamber stiffness constant (kc) that characterizes the overall slope of the EDPVR. A similar relationship holds for stress and strain. Reproduced with permission from Fuster V, Alexander RW, O’Rourke RA, et al: Hurst’s The Heart. 11th ed. New York: McGraw-Hill; 2004.
Diastolic chamber stiffness, similar to the systolic chamber stiffness index, Ees, is dependent on both material (myocardial) stiffness and ventricular chamber characteristics (eg, volume, mass). Myocardial stiffness is quantified from the relationship between diastolic LV wall stress (ε) and strain (σ). Strain is the deformation of the muscle produced by an applied force and is expressed as the percent change in length from the unstressed length. At any given strain throughout diastole, myocardial stiffness is equal to the slope (dσ/dε) of a tangent drawn to the stress–strain curve at that strain. Because the stress–strain relationship is generally exponential, the relationship between (dσ/dε) and stress is linear. The slope of this relationship is the modulus of myocardial stiffness (Km) and has been used to quantitate myocardial stiffness. Thus, when myocardial stiffness is increased, the stress–strain relationship shifts to the left, the slope of the (dσ/dε) versus stress relationship becomes steeper, and Km increases.
The end-diastolic P–V relationship (EDPVR) is constructed by connecting the end-diastolic points (lower right hand) of a series of PVAs; the relationship is nonlinear and defines the passive properties of the chamber when it is fully relaxed. The nonlinearity of the EDPVR results from the different types of structural proteins being stretched over the range of pressures and volumes. Thus, at the low end of the relationship, where operative stiffness is low, stiffness is caused by compliant elastin and sarcomeric titin. As volume increases and operative stiffness increases, the slack length of collagen and titin are exceeded, and stretch is resisted. At the other extreme (subphysiologic volumes), negative pressures are required to reduce volume (diastolic suction); however, negative pressures are rarely recorded in vivo, and less stringent criteria to establish the presence of diastolic suction are required. It is important to recall that changes in intrathoracic pressure, pericardial constraint, and ventricular interaction all influence the EDPVR. Analytic limitations similar to the ESPVR are present for the EDPVR; that is, comparisons of EDPVR should account for covariance of the parameters.40
A variety of curve fits for EDPVR using nonlinear regression analysis have been proposed, but single value indices of stiffness, such as the stiffness constant, have met with limited success. Chamber stiffness (kc) and myocardial stiffness (Km) provide load and chamber size-independent parameters of passive chamber and myocardial properties, respectively; however, when comparing hearts of different sizes, a simple approach is to measure the volume at a specified pressure.40
In isolated muscle, loading conditions represent the force applied to muscle before and after (preload and afterload, respectively) the onset of contraction. In the intact ventricle, preload and afterload are also determined by the volume status of the individual and the characteristics of the arterial and venous circulations (pulmonary and systemic circulations for the RV and LV, respectively). Thus, loading conditions are not only important direct determinants of ventricular performance, but they also function indirectly by coupling the ventricle to the vascular system.
Ventricular contraction transfers blood from the venous to the arterial side of the circulation, and arterial and venous capacitances (the change in volume per change in pressure, dV/dP) determine the respective pressures that result from the shift in blood volume. These pressures determine the driving force across the peripheral resistance (where resistance equals pressure gradient for flow divided by the cardiac output) and are primarily responsible for venous return to the heart.
The venous return curve describes the inverse relationship between venous pressure and cardiac output (Fig. 5–18A, B). In contrast to convention, the venous return curve plots the independent variable (cardiac output) on the vertical axis and the dependent variable (venous pressure) on the horizontal axis. The x-intercept is the mean circulatory pressure (ie, that pressure in the vascular system in the absence of cardiac pumping). The mean circulatory pressure is a function of the capacitance of the vascular system and the total blood volume. The plateau of the venous return curve and the y-intercept represents the maximal obtainable cardiac output as venous pressure is reduced. In the normal heart, cardiac output is limited by venous return, and the operating venous pressure is near the plateau of the venous return curve.
Venous pressure–cardiac output curves. The equilibrium point is defined by the intersection of the ventricular function curve with the venous return curve. A. Volume loading and venoconstriction shift the venous return curves to the right, resulting in an equilibrium point with a higher cardiac output and higher mean circulatory pressure. Volume depletion and venodilation shift the curve to the left, resulting in an equilibrium point with a lower cardiac output and a smaller mean circulatory pressure. B. The effects of arteriolar constriction and dilation on the venous return curves are more complex. At B (moderate heart failure), cardiac output is preserved at the expense of venous pressure. C. Sympathetic nerve stimulation causes a leftward shift of the ventricular function curve, resulting in an equilibrium point with a lower venous pressure and higher cardiac output. At C (severe heart failure), cardiac output is decreased and venous pressure further increased. D. Chronic heart failure causes rightward shifts of both the ventricular function and venous return curves. Reproduced with permission from Berne RM, Levy MN: Physiology, 2nd ed. St. Louis: Mosby; 1988.
Coupling of the venous system of the heart is graphically represented in Fig. 5–18C. In this analysis, the intersection of the ventricular function (Frank-Starling) curve and the venous return curve represents the steady-state operating values of cardiac output and venous pressure. At this equilibrium point, the ability of the venous system to provide venous return at a given pressure is matched with the ability of the ventricle to pump that venous return when distended to the same pressure.
Increased blood volume and venoconstriction shift the venous function curve upward and to the right, increasing the mean circulatory pressure and the maximal cardiac output (Fig. 5–18D). The venous system contains the major fraction of blood in the vascular system because of the greater capacitance of veins than of arteries. As a result, venoconstriction shifts significant quantities of blood from the peripheral to central circulation. Because arteries contain only a small percentage of the total blood volume, their contractile state does not affect the mean circulatory pressure. Moreover, because venous pressure varies inversely with systemic vascular resistance, arteriolar constriction (increased afterload) shifts the curve downward and to the left without changing the mean circulatory pressure. Conversely, arteriolar dilation shifts the curve upward and to the right. An increased inotropic state shifts the ventricular function curve to the left without significantly altering the venous return curve. Conversely, in chronic heart failure, there is a rightward shift of the ventricular function curve and because of renal salt and water retention, a parallel rightward shift of the vascular function curve. In this way, cardiac output is initially maintained at the expense of increased venous pressure and congestion. If the compensatory mechanisms fail, venous pressure increases further, and cardiac output falls.
Ventriculoarterial coupling can also be expressed in the P–V framework (Fig. 5–19). Arterial properties are represented by effective arterial elastance (EA), which incorporates the mean resistance and pulsatile features of the arterial load. EA is estimated by PES/SV, where PES is the end-systolic pressure and SV is the stroke volume. The EA/Ees ratio has been used as an index of ventriculoarterial coupling and has been shown to be a critical determinant of pump performance and efficiency. With increases in EA, stroke work initially increases, reaches a plateau, and then decreases. Maximum stroke work occurs when arterial and ventricular properties are equal (ie, when EA = Ees). Similar changes with increases in EA occurs with ventricular efficiency, defined as external stroke work/MVO2/beat, and are maximum when EA = Ees/2. Therefore, in this conceptual framework, energetically optimal ventriculoarterial coupling exists when the EA/Ees ratio ranges from 0.5 to 1.0.
Ventriculoarterial coupling using the arterial elastance-LV end-systolic relation (EA/EES). A. Baseline. B. Nitroprusside increases EES with little change in EA, resulting in a decrease in EA/EES. C. Angiotensin II increases EA without a change in EES, resulting in an increase in EA/EES. LV, left ventricular. Reproduced with permission from LeWinter MM, Suga H, Watkins MW: Cardiac Energetics: From Emax to Pressure-Volume Area. Norwell: Kluwer Academic Publishers; 1995.