## Abstract

To assess whether preload-adjusted maximal power (PAMP), which is calculated asW˙_{max}/V
(whereW˙_{max} is maximal power and V_{ed} is end-diastolic volume with β = 2) is an index of right ventricular (RV) contractility, we measured RV pressure (P) and volume (V) and pulmonary artery pressure and flow in 10 dogs at baseline and after inotropic stimulation. PAMP was derived from steady-state data, whereas the slope (*E*
_{es}) and intercept (V_{d}) of the end-systolic P-V relationship were derived from data obtained during vena caval occlusion. Inotropic stimulation increased *E*
_{es} (from 0.96 ± 0.25 to 1.62 ± 0.28 mmHg/ml; *P* < 0.001) and V_{d} (from −3.0 ± 17.2 to 12.4 ± 10.8 ml;*P* < 0.05) but not PAMP (from 0.24 ± 0.10 to 0.36 ± 0.22 mW/ml^{2}; *P* = 0.09). We found a strong relationship between the optimal β-factor for preload adjustment and V_{d}. A corrected PAMP, PAMP_{c}= W˙_{max}/(V_{ed} − V_{d})^{2}, which incorporated the V_{d}dependency, was sensitive to the inotropic changes (from 0.23 ± 0.12 to 0.54 ± 0.17 mW/ml^{2}; *P* < 0.001) with a good correlation with *E*
_{es}(*r* = 0.88; *P* < 0.001).

- contractility
- hydraulic
- pressure-volume

ventricular contractility is usually derived from a series of pressure-volume (P-V) loops that are measured during progressively altered cardiac loading conditions. It has been shown that the slope of the end-systolic P-V relation, which is also called the end-systolic or maximal elastance (*E*
_{es}; Refs. 17,18), reflects ventricular contractility independent of preload and afterload. Although *E*
_{es} was first applied to the left ventricle, this index has also been validated for the right ventricle (5). Measuring*E*
_{es} requires invasive techniques because simultaneous pressure and volume measurements are needed as well as the alteration of loading conditions. Hence its assessment is usually restricted to experimental settings and small-scale studies.

To obviate the need for measurement of P-V loops, hydraulic power has been proposed as an alternative way to quantify left ventricular (LV) contractility (2). For the left ventricle, hydraulic power is calculated as the instantaneous product of aortic pressure and flow during steady-state conditions to yield power (W˙) as a function of time (10). Maximal power (W˙_{max}) is the maximal value of this curve (6, 8). An important limitation of W˙_{max} as an index of contractility is its preload dependency (6). However, it has been shown for the left ventricle that W˙_{max} can be corrected for this preload dependency by dividing W˙_{max} by the LV end-diastolic volume (V
) with a β-value of 2 (6, 14). The indexW˙_{max}/V
is usually referred to as the preload-adjusted W˙_{max} (PAMP).

The aim of this study was to validate PAMP as an index for right ventricular (RV) contractility. To do so, we calculated PAMP and*E*
_{es} in mongrel dogs at baseline and after dobutamine infusion. Although the inotropic effect of dobutamine was clearly demonstated by *E*
_{es}, this was not the case for PAMP. To elucidate the apparent discrepancy between these indices of ventricular contractility, we used the experimental data to assess the parameters of a mathematical heart arterial interaction model. Computer simulations revealed a nonlinear relationship between the intercept of the end-systolic P-V relationship (V_{d}) and the β-coefficient that should ideally be used to adjust for preload, with the value of 2 being valid only when V_{d} equals zero. Based on the mathematical model simulations, we propose a correction of PAMP (PAMP_{c}) that takes into account this V_{d}dependency. It is then verified whether PAMP_{c} allows us to detect the effects of inotropic interventions in the animal experiments.

## MATERIALS AND METHODS

### Experimental Protocol

This investigation conforms with guidelines established in the*Guide for the Care and Use of Laboratory Animals* [DHEW Publication No. (NIH) 85-23, Revised 1986, Office of Science and Health Reports, DRR/NIH, Bethesda, MD 20205] and was approved by the ethical committee of the Katholieke Universiteit Leuven.

Ten healthy mongrel dogs (body wt 18–24 kg) were included in this study. After the dogs were premedicated with ketamine hydrochloride (10 mg/kg im; Ketalar) and piritramide (1 mg/kg im; Dipidolor, Janssen Pharmaceutica), anesthesia was induced with pentobarbital sodium (10 mg/kg iv; Nembutal). Endotracheal intubation was performed, and the lungs were mechanically ventilated with a 50% mixture of oxygen in air. Anesthesia was maintained with a continuous infusion of pentobarbital sodium (1 mg/kg) and piritramide (1 mg/kg). Arterial blood gases were measured at regular intervals, and ventilation was adjusted accordingly to maintain normocapnia. Normothermia was maintained by means of a heating mattress.

A fluid-filled catheter was inserted into the descending aorta via the right femoral artery to monitor systemic blood pressure and obtain samples for blood-gas analysis. Via a midline sternotomy, the inferior vena cava was dissected, and a band was placed around it for controlled alterations of RV preload. The heart was suspended in a pericardial cradle. The main pulmonary artery (PA) was dissected free from the aorta, and a 16- or 18-mm perivascular flow probe was placed around it and connected to an electromagnetic blood-flow meter (Skalar; Delft, The Netherlands) to provide continuous display of cardiac output. PA pressure (P_{PA}) and right atrial pressure were measured with fluid-filled catheters inserted through purse-string sutures in the RV outflow tract and right auricle, respectively. A 5-Fr microtipped pressure-transducer catheter (Millar Instruments; Houston, TX) and a 5-Fr dual-field conductance catheter (Millar Instruments) were inserted into the right ventricle through small stab wounds in the RV outflow tract. The correct position of the conductance catheter was determined by palpation and confirmed by observation of pressure and segmental volume signals with appropriate phase relationships. In all animals, all five conductance-catheter segments were located within the ventricle and used for analysis. The conductance catheter was connected to a volumetric system (Sigma 5, CD Leycom/CardioDynamics; Zoetermeer, The Netherlands), which was also used to measure blood resistivity. Parallel conductance was determined for each experimental condition by the rapid injection of 7 ml of 10% saline solution into the right atrium (1). For each animal, the α-calibration factor was assessed as the ratio of stroke volume (SV) measured with the conductance catheter (SV_{cond}) and the PA flow probe (SV_{flow}) during steady-state conditions at baseline and was assumed to remain constant throughout the measurements. The correlation between SV_{cond} and SV_{flow} was 0.94 (*P* = 0.0003). Average SV_{cond}/SV_{flow}, i.e., the α-calibration factor, was 1.18 ± 0.22 with a range of 0.75–1.38.

Data were recorded with the open chest and pericardium at steady-state baseline conditions and during transient vena cava occlusion that was induced via a band around the inferior vena cava for ∼10 s. All data were measured with the ventilation suspended at end expiration. After hemodynamic data values had returned to baseline, all measurements (steady state and vena cava occlusion) were repeated during infusion of dobutamine (5–10 μg · kg^{−1} · min^{−1} iv) with the measurements starting at least 15 min after the onset of the infusion.

Data were measured at a sample rate of 250 Hz and stored on hard disk for subsequent analysis using customized software written in Matlab (Mathworks). Individual cardiac cycles were identified using the first minimum that preceded peak RV dP/d*t*, i.e., the onset of RV isovolumic contraction.

### Experimental Data Analysis

#### Steady-state data.

Steady-state data were averaged over at least five cycles and yielded one cardiac cycle of RV pressure and volume, P_{PA}, and PA flow (Q˙_{PA}) data. Heart rate (HR) was obtained from the duration of the average cardiac cycle. Systolic and diastolic pressure and mean P_{PA} values (SBP, DBP, and MAP, respectively) were calculated from P_{PA}. SV was calculated as the area under the flow curve, and cardiac output (CO) was obtained from HR and SV. RV V_{ed} was defined as the maximum value of the calibrated conductance-catheter signal. RV pressure at the onset of RV isovolumic contraction was used as the end-diastolic pressure (P_{ed}). W˙_{max} was calculated as the maximum of the instantaneous product of P_{PA} andQ˙_{PA}. PAMP was calculated asW˙_{max}/V
.

#### Vena cava occlusion data.

Five to ten successive beats were selected from the P-V loops measured during vena cava occlusion. Stroke work (SW) was calculated for each cycle as the area enclosed by the PV loop. The slope of the linear regression equation (M_{w}) on the V_{ed}-SW data yielded preload-recruitable SW. To calculate the slope (*E*
_{es}) and intercept (V_{d}) of the end-systolic points in the P-V plane, we used an iterative method to identify these end-systolic points. We first calculated elastance [*E*(*t*)] as P_{RV}/(V_{RV}− V_{d}), where P_{RV} and V_{RV} were RV pressure and volume, respectively. V_{d} was given the initial value of zero. The points in the P-V plane that corresponded to maximal*E*(*t*) for each cycle were identified as the end-systolic points. Linear regression analysis on these points yielded a first estimate of *E*
_{es} and a new estimate of V_{d}. This procedure was repeated with the resulting V_{d} values until successive values for V_{d} did not differ by >0.1%. For all data, convergence was reached within 3 or 4 iterations.

For each cardiac cycle, W˙_{max} was calculated as the maximum of the instantaneous product of P_{PA} andQ˙_{PA} and plotted against V_{ed} for that cycle. A power law of the form W˙_{max} = αV
was fitted through these points.

### Heart Arterial Interaction Model

#### Model description.

The heart arterial interaction model allows the calculation of RV pressure and volume as well as P_{PA} and PA flow (Fig.1). RV function is described by a time-varying elastance function, whereas the arterial load is represented by a four-element, lumped-parameter windkessel model (16). The pulmonary arterial model parameters are total peripheral resistance (*R*), total arterial compliance (C), total inertance (L), and PA characteristic impedance (*Z*
_{0}). Cardiac parameters are*E*
_{es}, V_{d}, the slope of the diastolic P-V relationship (*E*
_{min}) and RV end-diastolic pressure (P_{ed}), HR, and the time to reach maximal elastance (*t*
_{Ees}). Tricuspid and pulmonary valves are simulated as frictionless, perfectly closing devices that allow forward flow only.

#### Assessing cardiac and arterial model parameters.

In the computer model, *E*(*t*) is implemented in a normalized form [*E*
_{N}(*t*
_{N})], i.e., it is normalized with respect to *E*
_{es} and*t*
_{Ees} (13, 17). We first calculated such an average RV*E*
_{N}(*t*
_{N}) from 25 P-V loops measured in 6 dogs during baseline conditions. For each simulation, RV function is fully determined by the actual *E*(*t*) value calculated from*E*
_{N}(*t*
_{N}) and*E*
_{es},*t*
_{Ees},*E*
_{min}, HR, P_{ed}, and V_{d}values.

For each dog and each condition (baseline or inotropic stimulation), the vena cava occlusion data provide *E*
_{es} and V_{d}, whereas all other cardiac parameters (HR,*t*
_{Ees}, P_{ed}, V_{ed}) follow from the corresponding steady-state data. P_{ed} and HR follow directly from these data.*E*
_{min} is calculated as P_{ed}/(V_{ed} − V_{d}), and*t*
_{Ees} is assumed to be 35% of the duration of the cardiac cycle. The four arterial parameters (*R*, C, *Z*
_{0}, and L) are derived from P_{PA} and Q˙_{PA} measured at baseline via standard parameter-fitting techniques. Q˙_{PA} is used as an input to the four-element windkessel model, and the pressure predicted by the four-element windkessel model is fitted to P_{PA} by adjusting the four model parameters.

### Model Simulations: Steady-State Data and Vena Cava Occlusion

All model parameters were assessed for all animals at baseline and inotropic stimulation (20 reference parameter sets). SBP, DBP, and SV as predicted by the simulations were compared with the actual measured values.

For each animal and each condition, P_{RV}, V_{RV}, P_{PA}, and Q˙_{PA} were calculated with the reference parameter set and with four lower values of P_{ed}(in steps of 0.5 mmHg) to simulate the effect of reduced preload (i.e., V_{ed} reduction due to vena cava occlusion) over the same range as measured in the animals. For each cycle,W˙_{max} was calculated and plotted as a function of V_{ed}. Similar to the animal experimental data, a power law of the form P_{max} = αV
was fitted through these points.

### Impact of V_{d} on PAMP

The computer-simulation data revealed that the optimal β-coefficient to be used to correct W˙_{max} for preload is not a constant. In contrast, there was a strong relationship between V_{d} and the β-coefficient. To study the impact of V_{d} on PAMP, we first used the computer-simulation data to show that when W˙_{max} is plotted against V_{ed} − V_{d}, there is again a power-law relationship between both although with a constant β-value that appears to be close to 2. Second, based on these findings, we derived a corrected PAMP index, PAMP_{c}, which is defined asW˙_{max}/(V_{ed} − V_{d})^{2}. Third, it was verified whether the relationship between V_{d} and the β-coefficient, which was observed in the computer-simulation data, was also present in the experimental data. Fourth, PAMP_{c} was applied to the experimental data.

### Statistical Analysis

Experimental data are given as means ± SD. Baseline data were compared with inotropic stimulation data using two-tailed, paired*t*-tests (SigmaStat 2.0, Jandel Scientific). Power-law fitting (the V_{ed} − W˙_{max} relation) was done using SigmaPlot 3.0 (Jandel Scientific) nonlinear regression tools. Linear relations between indices were studied by linear correlation and/or linear regression analysis (SigmaPlot 3.0).

## RESULTS

### Experimental Data Analysis

Hemodynamic data at baseline and with inotropic stimulation are shown in Table 1. Inotropic stimulation significantly increased blood pressure with a trend toward increased CO (Fig. 2). The improved RV systolic function was partly due to an increase in V_{ed}(*P* = 0.07) but also to an increase in contractility as indicated by *E*
_{es} and the slope of the preload-recruitable, stroke-work relationship (*M*
_{w}; Fig. 2). Besides an increase in the slope of the end-systolic P-V relation, we also found an increase in its intercept V_{d} (Fig. 2). W˙_{max} was higher after inotropic stimulation (see Table 1), but after correction for V_{ed}, PAMP was not different between baseline and inotropic stimulation (Fig. 3) groups.

### Model Simulations

Figure 4 shows the correspondence between measured and simulated SBP, DBP, and SV. Pressures were underestimated by 12 and 8% on average for SBP and DBP, respectively. Linear regression analysis yielded *y* = −2.46 + 0.95*x*; *r*
^{2} = 0.94 for SBP and*y* = −0.74 + 0.97*x*;*r*
^{2} = 0.98 for DBP. On average, predicted SV was <5% higher than measured SV with the regression line given by*y* = 0.98 + 0.99*x*;*r*
^{2} = 0.97.

### Impact of V_{d} on PAMP

The power-law β-coefficients derived from the vena cava occlusion simulations varied from 0.98 to 3.89. Linear correlation analysis showed a strong correlation between V_{d} and the β-coefficient (*r*
^{2} = 0.86), but the relationship between V_{d} and the β-coefficient was better described with an exponential function (*r*
^{2}= 0.95) as shown in Fig. 5. The power-law fitting for the simulated baseline and inotropic stimulation case for*dog 7* are given as an example in Fig.6. In the standard V_{ed}− W˙_{max} plot, the β-coefficient varied from 1.7 to 2.3. Plotting the same W˙_{max} data as a function of V_{ed} − V_{d} and fitting the power law yielded β-values of 1.95 and 1.97. The average β-coefficient for all dogs and all conditions obtained in this way was 1.99 ± 0.06 (range, 1.85–2.17). We thus propose PAMP_{c} =W˙_{max}/(V_{ed} − V_{d})^{2} as a PAMP_{c} index. Although more subject to scatter, the β-V_{d} relationship was also present in the experimental data (see Fig. 5).

Figure 3 also shows that PAMP_{c} was higher in the inotropic stimulation group than in the baseline group. The correlation between*E*
_{es} and PAMP is weak (*r* = 0.36; not significant) in contrast to the correlation between*E*
_{es} and PAMP_{c} (*r* = 0.88; *P* < 0.001; Fig.7).

## DISCUSSION

We have applied two indices known to reflect LV contractility (i.e., *E*
_{es} and PAMP) to the right ventricle. Pharmacological inotropic stimulation yielded the anticipated increase in *E*
_{es} but not PAMP. A parameter study on a computer model that simulated heart arterial interaction revealed that V_{d}, the intercept of the end-systolic P-V relation, determines the β-power coefficient that should be used to correctW˙_{max} for preload. The value β = 2 [as proposed by Kass and co-workers (6)] only applies when V_{d} is negligible. Using computer simulations as well as experimental data, we have shown that this problem is overcome whenW˙_{max} is corrected asW˙_{max}/(V_{ed} − V_{d})^{2}, an index that we have indicated as PAMP_{c}.

In this open-chest, open-pericardium experimental study, inotropic stimulation led to an increase in P_{PA} and a borderline increase in CO. It was shown earlier (5) that when afterload is constant, *E*
_{es} reflects changes in RV performance, which is confirmed by our results. In our study, there was a trend toward an increased total pulmonary vascular resistance (*P* = 0.08) and decreased pulmonary vascular compliance (*P* = 0.09), i.e., an increased afterload. These effects may explain the trend toward an increased RV end-diastolic volume.

To our knowledge, PAMP has not previously been studied in the right ventricle. Our data demonstrate that PAMP, defined asW˙_{max}/V
, is not an accurate index of RV contractility. The concept of PAMP is based on the observation that the power-law relationship W˙_{max}= αV_{ed2} can be fitted throughW˙_{max} − V_{ed} data points obtained under altered loading conditions. Optimal preload correction is then obtained by dividing W˙_{max} by V
, and α is a measure of contractility. It is important, however, that for a generally applicable index of contractility, the coefficient used in the index (i.e., β) is constant and independent of physiological variables. In both our experimental data and computer simulations, the β-factor (usually assumed to be 2) was not constant but varied with V_{d} (range, 0.98–3.89). As a consequence,W˙_{max}/V
cannot be used as such to quantify contractility; in some cases (when V_{d} < 0), the factor 2 will be too high, whereas in other cases (when V_{d} > 0), it will be too low. The β-V_{d}relationship can be expressed as β = 2.008e^{0.0216Vd}}. It is only when V_{d}= 0 that the β-value approximates 2.

Supposing that V_{d} is known, one could think of correcting W˙_{max} by dividing by V
, with the β-coefficient calculated as 2.008e^{0.0216Vd}. This approach, however, does not work. For example, in Fig. 6 (*dog 7*), using the optimal value for the β-value,W˙_{max}/V
= 0.322 at baseline and W˙_{max}/V
= 0.086; i.e., a lower α-value is found during inotropic stimulation. A more consistent approach is to plot W˙_{max} as a function of V_{ed} − V_{d}. Doing so, the optimal value for the β-coefficient becomes constant with the computer simulations yielding an average value of 1.99 ± 0.06 (range, 1.85–2.17). Therefore, when PAMP is modified asW˙_{max}/(V_{ed} − V_{d})^{β}, the V_{d} dependency of the β-factor disappears: the β-value is indeed approximately constant, and the formula is generally applicable with β = 2. For the same example of Fig. 6, W˙_{max}/(V_{ed} − V_{d})^{2} becomes 0.090 at baseline and 0.331 during inotropic stimulation, respectively, which is consistent with the anticipated effect of inotropic stimulation. This corrected PAMP is indicated as PAMP_{c}. Application on the experimental data revealed a significant difference between baseline and inotropic stimulation. Furthermore, in contrast to PAMP, PAMP_{c} shows an excellent correlation with *E*
_{es}(*r* = 0.88; *P* < 0.001).

At present, we do not have the necessary experimental data to study PAMP and the relationship of the β-value with V_{d} for the left ventricle. However, model simulations with parameters for the LV and systemic arterial system (data not shown) revealed the same tendency. Note, however, that a nonconstant β-coefficient for the left ventricle was earlier reported by Kass and co-workers (11). They proposed to use β = 1 for the normal left ventricle, whereas β = 2 should be more appropriate for dilated ventricles. This finding is in line with our data, as it can be assumed that in the normal left ventricle V_{d} is small (and may even be negative), whereas with LV enlargement, V_{d}shifts to the right and requires higher β-values.

In our study, the increase in *E*
_{es} was accompanied by a trend toward rightward shifts of V_{ed} and the intercept of the end-systolic P-V relation. Variation of RV V_{d} over the cardiac cycle was reported by several authors (3, 5, 9), but a rightward shift in V_{d} with dobutamine infusion has to our knowledge not been documented for the right ventricle. RV volumes were measured with a conductance catheter, which is an indirect technique whereby the measured signals require calibration (parallel conductance and α-slope factor) to obtain volumes. As such, the observed rightward shifts in V_{ed} and V_{d} may be due to physiological phenomena induced by dobutamine infusion and the borderline increase in RV afterload but also to variations or measurement errors in parallel conductance or α-value. In this study, the α-value was higher than reported in other studies (4, 15), where transit-time flow probes were used as a reference (on the order of 0.7–0.8). We may thus have underestimated absolute volumes. At the same time, however, this discrepancy illustrates the difficulties in measuring absolute volumes with the conductance-catheter technique and explains our rationale for investigating indices for RV contractility based on hydraulic power. Unfortunately, these indices require correction for preload, which is approximated as RV end-diastolic volume. As such, they cannot be fully uncoupled from volume measurements and the inherent measuring uncertainties. The sensitivity to measuring errors in absolute volumes is particularly problematic when PAMP is calculated asW˙_{max}/V
, because *1*) small volume changes are squared, and *2*) the β-factor is not constant and equal to 2 but instead varies with V_{d}. By using W˙_{max}/(V_{ed} − V_{d})^{2}, these important limitations are attenuated, as volume shifts due to changes or erroneous measurement of parallel conductance have the same effect on V_{d} and V_{ed} (and thus do not affect V_{ed} − V_{d}), and the β-factor is approximately constant and equal to 2. By plotting W˙_{max} as a function of V_{ed} − V_{d}, V_{d} can be seen as a “correction volume,” and the exact position of the P-V loop on the volume axis is less important for calculating PAMP. Within the linearized time-varying elastance concept, V_{d} is the theoretical volume (derived from linear extrapolation) for which the ventricle does not generate any pressure. It is only when filled to volumes higher than V_{d} that the ventricle generates pressure. Therefore, within the linear time-varying elastance concept, V_{ed} − V_{d} is perhaps a more appropriate marker of ventricular preload.

PAMP is potentially advantageous over *E*
_{es} to characterize ventricular contractility, as its calculation does not require multiple P-V loops recorded under altered loading conditions. Moreover, for the left ventricle, W˙_{max} can be computed from noninvasively measured arterial pressure (tonometry) and flow (Doppler echocardiography) (7). With a noninvasive estimate of V_{ed}, PAMP would be a noninvasive measure of LV contractility. Our data, however, indicate that W˙_{max}is best corrected for preload using (V_{ed} − V_{d})^{2}. Because V_{d} can only be determined from multiple P-V loops measured under altered loading conditions, the measurement of PAMP requires steady-state pressure and flow in the PA as well as RV P-V loops measured during caval vein occlusion. One can therefore only conclude that*E*
_{es} is easier to obtain than corrected PAMP and that there is no simple index for LV or RV contractility based on one single (P-V loop) measurement during steady-state conditions.

This study is partly based on computer simulations where RV function is characterized by a time-varying elastance model, whereas the pulmonary arterial system is represented by a four-element windkessel model. The model was developed for LV heart arterial interaction studies and was validated using LV and aortic experimental data (12). Application of the model to the right ventricle is subject to some limitations. With the four-element windkessel model, we were able to fit the low-frequency (<5th harmonic) contents of the impedance spectra derived from measured P_{PA} andQ˙_{PA}, but the model often failed to accurately match the higher harmonics. Owing to this limitation, the model can predict low-frequency information such as SBP or DBP, SV, or maximalQ˙_{PA}, but the waveforms are poorly predicted. Another limitation is the use of the time-varying elastance concept for the right ventricle. Several authors (3, 9) reported an increase in the slope and a leftward shift of the intercept of the P-V relationship during RV contraction. In the experimental data analysis as well as the computer simulations, we assumed constant V_{d}. These computer-model limitations, however, do not affect the outcome of this study. All observations based on the computer-model simulations were confirmed by the experimental data and vice versa, although often with more scatter in the experimental data. Besides the above-mentioned computer-model limitations, the higher scatter is also due to the fact that simulated interventions are based on varying a single condition (e.g., filling pressure), which is an idealized situation that is virtually impossible to achieve in vivo. We also acknowledge that our experimental model has a number of potential limitations. First, it is relatively invasive, and pentobarbital sodium may have a negative inotropic effect. However, the latter should not have any influence on our results, because inotropic state was modulated in the experimental protocol. Second, the presence of an open chest and pericardium may have caused larger volume changes than a closed chest/closed pericardium setting, but we would not expect this to influence our main observation that V_{d} is required for an adequate correction of PAMP. Finally, the experimental data do not allow for assessment of the sensitivity of PAMP_{c} to HR or afterload changes. The present findings therefore cannot automatically be transposed to normal human beings; further studies in closed chest/closed pericardium conditions are required to validate our observations.

In summary, experimental data showed that in contrast to*E*
_{es}, PAMP could not quantify the hemodynamic effect of inotropic stimulation of the right ventricle. This is due to the rightward shift in V_{d}, the intercept of the end-systolic P-V relation, and to the fact that W˙_{max}should be adjusted for preload by dividing it by (V_{ed}− V_{d})^{2}. Therefore, correct computation of PAMP requires data measured under altered loading conditions. This restriction limits the clinical value of PAMP for characterizing RV contractility.

## Acknowledgments

This study was supported by Grant 1.5.208.99 from the Fonds voor wetenschappelijk onderzoek-Vlaanderen (FWO-Vlaanderen) to P. F. Wouters. P. Segers is the recipient of a postdoctoral grant from FWO-Vlaanderen. H. A. Leather is the recipient of a doctoral grant from FWO-Vlaanderen.

## Footnotes

Address for reprint requests and other correspondence: P. Segers, Hydraulics Laboratory, Institute Biomedical Technology, Ghent Univ., Sint-Pietersnieuwstraat 41, 9000 Ghent, Belgium (E-mail:patrick.segers{at}rug.ac.be).

The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked “

*advertisement*” in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.June 13, 2002;10.1152/ajpheart.00340.2002

- Copyright © 2002 the American Physiological Society