The debate surrounding the benefits of pulsatile versus nonpulsatile flow during acute and chronic mechanical circulatory support is far from over.1–3 Although several investigators have clearly documented the benefits of pulsatile flow,4–6 many others could not show any differences when using different modes of flow.7–9
To end this long controversy, the first step must be to precisely quantify pressure flow waveforms for direct comparison.1–3,10,11 Without precise quantification, it is impossible to make meaningful comparisons between pulsatile and nonpulsatile flow. This becomes increasingly obvious with different types of pulsatility. Although there are a few adequate formulas to quantify the pressure flow waveforms in the literature, none of them have been adopted and accepted for universal use.1,10–12 The vast majority of investigators used pulse pressure to quantify and compare pulsatile versus nonpulsatile flow.1,3 We have repeatedly suggested that pressure flow waveforms should be quantified in terms of hemodynamic energy levels, not pulse pressure, because generation of pulsatile flow depends upon an energy gradient.1–3,10,11 The energy equivalent pressure (EEP) and surplus hemodynamic energy (SHE) formulas are unique methods that can be used to precisely quantify pressure flow waveforms of different modes of flow.10,11
The objective of this study was to quantify pressure flow waveforms in terms of EEP and SHE levels in an adult mock loop using a pulsatile ventricular assist device (VAD).
Materials and Methods
Precise Quantification of Pressure Flow Waveforms
Energy equivalent pressure.
Shepard's EEP formula is based upon the ratio between the area beneath the hemodynamic power curve (∫ fpdt) and the area beneath the pump flow curve (∫ fdt) during each pulse cycle:10
where f is the pump flow rate, p is the arterial pressure (mm Hg), and dt is the increment in time. The unit of the EEP is mm Hg.
Total hemodynamic energy.
Using Shepard's total hemodynamic energy formula,10
the constant 1,332 changes pressure from units of millimeters of mercury to units of dynes per square centimeters.10
Surplus hemodynamic energy.
Surplus hemodynamic energy is calculated by multiplying the difference between the EEP and the mean arterial pressure (MAP) by 1,332.
For instance, if the MAP is 80 mm Hg and the EEP is 88 mm Hg, the difference is 8 mm Hg, and 8 mm Hg represents 10,656 ergs/cm3 (8 × 1,332 = 10, 656 ergs/cm3).
A Sarns/3 M 70 cc Pierce-Donachy pneumatic pulsatile VAD was used with a Penn State adult mock loop.13,14 A blood analog fluid consisting of 40% glycerin and 60% water (by volume) was used as the test fluid. Pressure transducers (Maxxim Medical, Athens, TX) in the arterial and venous compliance chambers were used to monitor the aortic and venous pressures, respectively. A Transonic ultrasonic flow probe (H14XL with T101 flow meter, Ithaca, NY) was placed downstream of the aortic valve to measure cardiac output. The pressure and flow waveforms were acquired with a National Instruments (Austin, TX) data acquisition board (PCI-6036E) and signal conditioner (SC-2345 carrier with SCC-SG04 and SCC-FT01 modules). Custom-written LabView (National Instruments) software was used to acquire and analyze the waveforms.
The pump flow rate was kept constant at 5 L/min with pump flow rates of 70 and 80 bpm and aortic pressures of 80, 90, and 100 mm Hg, respectively. Pump flows were adjusted by varying the systolic pressure, systolic duration, and the diastolic vacuum of the pneumatic drive unit. The aortic pressure was adjusted by varying the systemic resistance of the mock loop. When the desired MAP was achieved, a 30 second segment of the pressure and flow waveforms was recorded for further analysis. Each set of measurements was repeated eight times. The EEP and SHE were calculated at each experimental stage (n = 8). The difference between the EEP and the mean aortic pressure is the extra energy generated by this device. This difference is approximately 10% in a normal human heart.11
For each response of interest and pump rate combination, an analysis of variance (ANOVA) model was fit to the data to assess differences between aortic pressure groups, and p values for mean difference pressure estimates were adjusted for pairwise comparison testing using Tukey's multiple comparison procedure. The data are presented as the mean ± one standard deviation (SD).
Figures 1–3 represent sample aortic pressure and pump flow waveforms with an aortic pressure of 80 mm Hg, 90 mm Hg, and 100 mm Hg, respectively. Pump flow rates (5 L/min) and pump rates (70 bpm) were constant at all three experimental stages.
The EEP levels were 88.3 ± 0.9 mm Hg, 98.1 ± 1.3 mm Hg, and 107.4 ± 1.0 mm Hg with a pump rate of 70 bpm, and an aortic pressure of 80 mm Hg, 90 mm Hg, and 100 mm Hg, respectively. The percent change from aortic pressure to EEP was 10.4 ± 1.2%, 9.0 ± 1.4 %, and 7.4 ± 1.0% at the same experimental stages. Surplus hemodynamic energy in terms of ergs/cm3 was 11,039 ± 1,236 ergs/cm3, 10,839 ± 1,659 ergs/cm3, and 9,857 ± 1,289 ergs/cm3, respectively. Table 1 summarizes all of our results with different pump rates and aortic pressures.
Our results clearly suggest that this particular pulsatile VAD system generates adequate pulsatile flow and produces significant total hemodynamic energy at each experimental stage. Propulsatile flow investigators have repeatedly insisted that because of this extra energy (EEP and SHE), pulsatile flow maintains more physiologic microcirculation and better end organ recovery.1–3,9,10 However, prononpulsatile flow investigators have claimed that pulsatility does not exist in capillaries, but they have failed to show scientific data or cite any published investigation.1 A recent article on the impact of pulsatile flow upon microcirculation correlates very well with the propulsatile flow investigators' hypothesis. Recently, Baba and associates6 used a total artificial heart to investigate the effects of pulsatile and nonpulsatile flow on microcirculation of the bulbar conjunctiva in a goat model. They have clearly documented that erythrocyte velocity (a 75% decrease with nonpulsatile flow) and the number of perfused capillaries (44% decrease with nonpulsatile flow) was significantly higher with pulsatile flow when compared with nonpulsatile flow.6
If an axial flow pump or a centrifugal pump with a 100% nonpulsatile flow was used in our Penn State adult mock loop, then the integral signs could be removed from the EEP and surplus hemodynamic energy formulas. Because of 100% nonpulsatile flow, the EEP becomes equivalent to the MAP. Therefore, there will be no difference between the EEP and the MAP for 100% nonpulsatile flow. The surplus hemodynamic energy level will also be zero. Only this example demonstrates that axial or centrifugal pumps with 100% nonpulsatile flow have severe disadvantage in terms of hemodynamic energy levels.1
The native heart of patients who are supported by nonpulsatile pumps typically undergoes partial recovery and begins to eject after a few weeks of support. In this particular scenario, even though the nonpulsatile pump produces no pulsatility, there is diminished pulsatility because the native heart is pumping.1 At this stage, we can compare two different types of pulsatile flow, rather than pulsatile versus nonpulsatile flow. To the best of our knowledge, only one centrifugal pump produces 100% nonpulsatile flow regardless of the native heart.9 All others produce some degree of pulsatility. If the patient's native heart has acceptable recovery, then it is also possible to observe physiologic pulsatility with a nonpulsatile pump.15 Therefore, the precise quantification of arterial pressure and pump flow waveforms is a must, not an option, for direct comparisons between the pumping modalities.
The EEP and SHE formulas are adequate to precisely quantify pressure flow waveforms. This particular pulsatile VAD system with the Penn State adult mock loop produces near physiologic hemodynamic energy levels at each experimental stage. Further studies are planned to apply these formulas in an in vivo setting in the near future.
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