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Ventricular–Arterial Interaction in Patients With Heart Failure and a Preserved Ejection Fraction
Case Study
A 79-year-old obese male (body mass index 33 kg/m 2 ) with prior history of coronary bypass grafting, hypertension, and diabetes presented with New York Heart Association functional class III dyspnea progressive over 1 year. Vitals included a blood pressure of 162/90 mmHg and heart rate of 55 bpm. Jugular venous pressure and cardiac auscultation was normal, and there was no edema. Stress echocardiogram demonstrated no ischemia with a normal ejection fraction (60%), estimated right ventricular systolic pressure of 24 mmHg, and an E/e′ ratio of 8. Due to persistent unexplained dyspnea, hemodynamic exercise catheterization was pursued, which demonstrated a high pulmonary capillary wedge pressure (PCWP) at rest along with central aortic hypertension ( Fig. 6.1 ). With exercise, there was a pathologic increase in left-sided filling pressures and impaired cardiac output response consistent with underlying heart failure with preserved ejection fraction (HFpEF).
As we will explain in this chapter, these central hemodynamic derangements were coupled with abnormal pulsatile, reflective, and steady-state load that failed to demonstrate the normal expected degree of reduction in arterial load during exercise. Abnormalities in vascular stiffness and their coupling with ventricular systolic and diastolic performance during rest and exercise contribute to exercise intolerance and symptom manifestation in HFpEF.
Introduction
The mammalian cardiovascular system evolved to provide adequate flow at physiologic pressures both at rest and under a broad range of demands. Since blood flow is pulsatile, changes in cardiac output are accompanied by alterations in the arterial pulse wave amplitude and peak systolic pressure. To prevent wide fluctuations in blood pressure that otherwise can lead to vascular and end-organ damage, the heart and arteries are compliant so that pulse and peak pressures can be buffered, while systemic diastolic pressures are augmented. For the vasculature, this compliance is largely contained within the proximal conduit vessels, while within the heart it is described by the end-systolic stiffness (elastance, inverse of compliance) that is achieved during contraction. The normal heart develops a ventricular systolic stiffness of about 2.0 mmHg/mL of ventricular end-systolic volume, while arterial stiffness is about 1.5 mmHg/mL. These low values allow relatively large changes in volume in both the heart and the vascular bed to be achieved with only modest changes in ejected pressures.
With advancing age, both ventricular and arterial stiffness increase, and these changes may be further exacerbated by common disorders such as hypertension, obesity, diabetes mellitus, and renal disease. Since the heart and arterial systems are coupled, such stiffening results in amplification of systolic and pulse pressures during ejection, faster pressure decay during diastole, and enhanced cardiac systolic loading. Stiff arteries facilitate rapid transit of the flow and pressure pulse through the vasculature, increasing the velocity at which these waves encounter regions of impedance mismatch (e.g., distal arteriolar narrowings), which then increases the amplitude of reflected pressure waves, further exacerbating systolic load. The net effect is an adverse impact on cardiac systolic and diastolic function, with increased myocardial oxygen consumption required to provide the body with blood flow, impaired cardiovascular reserve function, labile systemic blood pressures, and diminished coronary flow reserve. Aging is also associated with endothelial dysfunction, which may in part relate to mechanical stiffening, as reduced wall distensibility can itself compromise endothelial-dependent responses to shear stress stimulation and vasorelaxation. One can consider this adverse interaction between heart and arteries as a form of coupling disease that ultimately limits the ability of the integrated cardiovascular system to respond to stress.
Abnormal ventricular-arterial stiffening and thus coupling may play an important role in patients with heart failure symptoms but with apparent preservation of systolic function. Such patients are typically older and female, with histories of chronic hypertension and a high prevalence of diabetes, obesity, and renal dysfunction. They often develop marked systolic hypertension under conditions of stress, and both their arterial and ventricular systolic pressures are very sensitive to blood volume status. While abnormal diastolic function is thought to contribute to heart failure symptoms by increasing congestion, it cannot explain the observed increases in systemic pressures, nor does it fully underlie limitations of cardiac reserve. Here, we review the pathophysiology of ventricular-arterial stiffening and its role in the syndrome of heart failure with a preserved ejection fraction (HFpEF).
Pathophysiology of Left Ventricular-Arterial Coupling
The influence of increases in ventricular systolic and vascular stiffness on net cardiovascular function is best depicted in the pressure-volume (P-V) plane. Ventricular end-systolic chamber stiffness is expressed as end-systolic elastance (Ees) ( Fig. 6.2 ), defined by the slope of the end-systolic P-V relationship. Ventricular afterload can be represented as aortic input impedance, derived from Fourier analysis of aortic pressure and flow waves. Impedance is expressed in the frequency domain and thus is more difficult to match with optimal measures of ventricular systolic function, which are typically determined in the time domain. One approach to this problem involves the development of a vascular parameter that shares units applicable to the heart, namely effective arterial elastance (Ea). Ea combines both mean and pulsatile loading, providing a lumped parameter that reflects the net impact of arterial vascular load on the heart. This index was developed and validated by Sunagawa et al. in the −1980s, and then applied and verified in humans by Kelly et al. The latter group confirmed that the simple ratio of end-systolic pressure to stroke volume (Pes/SV) could serve to estimate Ea in both hypertensive and normal humans. Graphically, Ea can be depicted as the absolute value of the slope of a line linking the coordinate points of end-systolic volume [Ves], Pes, and end-diastolic volume [Ved], P = 0 (see Fig. 6.2 ).
Coupling of heart and artery is often then depicted by the interaction of these two relations and expressed as a ratio of Ea/Ees. The intersection of these lines determines Pes and Ves (see Fig. 6.2 ). As shown in Fig. 6.3 , the Ea/Ees ratio is fairly preserved with normal aging to maintain optimal efficiency, declining somewhat in women, while both the numerator (vascular load) and the denominator (ventricular stiffness) increase. Ea is dominated by mean ventricular load, namely systemic vascular resistance (SVR), but it is also altered by artery stiffening to increase pulse pressure. Blood pressure pulsatility rises with aging, and this is reflected in the P-V diagram in the elderly individual (see Fig. 6.2 B) and the HFpEF patient (see Fig. 6.2 C) by the rise in systolic pressure throughout ejection ((arrows). Since Ea is determined by the ratio Pes/SV, the greater the disparity between Pes and mean arterial pressure (i.e., the more pulsatile or stiff the arterial system), the higher Ea will be relative to mean resistance load. Ea also varies directly with heart rate, since for any given cardiac SV, the systolic pressure will increase or decrease proportionally with the number of strokes (i.e., cardiac output). It is important to keep these factors in mind when interpreting data reporting on Ea and Ea/Ees coupling.
Effective coupling of heart to artery can be defined in several ways. One is the optimal transfer of blood from heart to periphery without excessive increases in pulse or systolic pressures. One can mathematically express optimal coupling as the interaction that best enhances the work performed by the heart on the body (i.e., optimal external work). Lastly, one must consider the efficiency of the heart in performing this work—the energy consumption required to effect this external work. All of these are reasonable definitions, although prior experimental and clinical studies have tended to focus on the latter two: optimizing external work and efficiency. For this, one can both predict and observe experimentally that an Ea/Ees coupling ratio of 0.6 to 1.2 achieves near optimal work and efficiency. This range is normally maintained under various physiologic stresses, as shown elegantly some years back in exercising animals. It can become very high, particularly in dilated cardiomyopathy, where depressed heart function (low Ees) is coupled to a high arterial impedance (high Ea).
Methods to evaluate arterial stiffness
Measurement of a vascular load is complex and not easily summarized by instantaneous single values of peak systolic and nadir diastolic pressures as is commonly used in clinical practice. Arterial pressures oscillate in a time-varying pulsatile fashion around a mean arterial blood pressure, and therefore a time-varying pulsatile load is more difficult to quantify by such static pressure measurements. In addition, arterial stiffness and pulsatile load are dynamic measures that vary with both heart rate and baseline ambient pressure in the systemic circulation. Furthermore, in terms of quantifying the vascular load on the left ventricle, measuring the central aortic pressure is more important than brachial blood pressure (which is subject to pulse pressure amplification of brachial systolic pressure).
For accurate measurement of pulsatile hemodynamics, pressure waveforms over time must be recorded with high fidelity. The gold standard measure is by invasive arterial pressure measurement, but this is obviously not practical for routine use. Applanation tonometry has enabled high-quality noninvasive measurement of radial or carotid arterial waveforms (although calibration to absolute pressures using independently measured brachial blood pressure is necessary). Since the carotid artery is anatomically close to the central aorta, carotid tonometry derived pressure can be used as a direct surrogate of central aortic pressure, whereas radial tonometry requires a mathematical transfer function to derive central aortic pressure from validated algorithms.
Once central aortic pressure is obtained, the pulse pressure (systolic-diastolic pressure) and stroke volume can be used to estimate total arterial compliance (TAC) (stroke volume/pulse pressure) as a measure of pulsatile load. This is the simplest and most commonly used method to estimate compliance, but it remains an estimate as the arterial circulation is not truly a chamber, and there is continuous outflow from the arterial capacitor during systole. Alternate methods of calculating TAC include: (1) an iterative pulse pressure method, where compliance is altered until the derived pulse pressure from a two-element Windkessel model matches measured pulse pressure and (2) analysis of the diastolic decay (during which there is no arterial inflow), which can be fit to an exponential curve to derive the Resistance*Compliance (RC) constant. Since SVR can be readily calculated by modeling the circulation as a steady-state resistor [SVR = mean arterial − atrial pressure/cardiac output], TAC can therefore be solved from this measured RC time constant. However, the stroke volume/pulse pressure estimate remains a widely used approximation of these more complex methods and has been validated.
As mentioned, effective arterial elastance provides a lumped measure of steady-state and pulsatile arterial load derived in the pressure volume plane and is also dependent on heart rate. This can be calculated by end-systolic pressure/stroke volume, with end-systolic pressure often approximated by 0.9 *systolic pressure. It is important to note that although Ea is inversely proportional to TAC, it is not merely the simple inverse of TAC but in fact incorporates the principal elements of vascular load, including peripheral resistance, total arterial compliance, characteristic impedance, and systolic and diastolic time intervals. In other words, Ea is affected by TAC (and varies inversely with it), but Ea is not in itself equal to 1/TAC.
Although the three-element Windkessel model is useful to study the arterial circulation, reflected waves also contribute to arterial load but are not well represented by such modeling. Pulse wave analysis can estimate augmentation of systolic pressure by reflected waves based on the initial inflection point in the systolic arterial pressure tracing from which the augmentation index can be calculated (augmentation pressure/pulse pressure). Wave separation analysis can further quantify the reflected load by separating the time-varying central aortic waveform into a forward wave and backward wave to calculate the reflection index (backward wave/forward wave). This wave separation can be done using either directly measured aortic flow in the ascending aorta (measured either invasively by a Doppler flow wire or noninvasively by echocardiographic pulsed wave Doppler or phase contrast magnetic resonance imaging [MRI]) or most commonly by estimation/modeling of flow patterns from the pressure tracing configuration alone. Such flow patterns when combined with instantaneous pressure also allow more complex measures of aortic characteristic impedance, which is typically represented in the frequency domain, and therefore is less intuitive to relate to existing pressure measures in the time domain.
Another commonly used measure of arterial stiffness that is widely utilized is measurement of pulse wave velocity (which should increase as arterial stiffness worsens). This can be measured from knowledge of the time delay of pressure wave onset between the carotid and radial/femoral artery sites measured by applanation tonometry. The distance between the two measurement sites, along with the time delay of onset, can be used to directly calculate pulse wave velocity as distance/time.
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