The controversy over the benefits of pulsatile flow compared with the nonpulsatile flow during cardiopulmonary bypass (CPB) procedures in pediatric as well as in adult cardiac patients continues.1–3 One of the major reasons for this debate is the lack of precise quantification of pressure-flow waveforms.4–6 Without a precise quantification, it is almost impossible to make direct and meaningful comparisons between perfusion modes or different types of pulsatility. To date, pulse pressure was used to quantify pulsatile and nonpulsatile perfusion in more than 90% of all published articles on this topic. Quantification solely in terms of pulse pressure is inadequate, because generation of pulsatile flow depends on an energy gradient.7,8 Quantification of pressure-flow waveforms in terms of hemodynamic energy levels is more precise than using pulse pressure. Therefore, the hemodynamic energy and surplus hemodynamic energy formulas should be considered for direct comparisons between different perfusion modes.
Precise Quantification of Pressure-Flow Waveforms
The following formulas were used to quantify precisely pressure-flow waveforms during pulsatile and nonpulsatile perfusion in this study.
Energy Equivalent Pressure Formula
Shepard’s EEP formula is based on the ratio between the area beneath the hemodynamic power curve (∫ fpdt) and the area beneath the pump flow curve (∫ fdt) during each pulse cycle:7
where f is the pump flow rate, p is the arterial pressure (mm Hg), and dt is the increment in time.
Percent Change in Pressure
Percent change in pressure from mean arterial pressure (MAP) to energy equivalent pressure (EEP) was calculated at each experimental stage.
Surplus Hemodynamic Energy
Surplus hemodynamic energy (SHE) is calculated by multiplying the difference between the EEP and the mean arterial pressure (MAP) by 1,332.
A diagram of the energy equivalent pressure and the surplus hemodynamic energy is shown in Figure 1.
The objective of this investigation was to compare pulsatile versus nonpulsatile perfusion modes in terms of EEP, percent change in pressure (from mean arterial pressure to EEP), and SHE levels during CPB in a simulated neonatal model.
Materials and Methods
The extracorporeal circuit consisted of a Jostra HL-20 heart-lung machine (for both pulsatile and nonpulsatile modes of perfusion), a Capiox Baby RX hollow-fiber membrane oxygenator, a Capiox pediatric arterial filter, 5 feet of arterial tubing, and 6 feet of venous tubing with a quarter-inch diameter. The circuit was primed with a lactated Ringers solution. The total priming volume was 350 ml. A Hoffman clamp was placed at the distal end of the arterial line for maintaining the circuit system pressure of 100 mm Hg (Figure 2).
The pseudo patient was subjected to five pump flow rates in the 400 to 800 ml/min range. When the target pump flow rate was achieved, a 20-second segment of the pressure and flow waveforms with nonpulsatile flow was recorded. Then, the perfusion mode was switched to the pulsatile flow. Pulsatile flow settings in the roller pump were 10% of the base flow, 120 beats per minute of the pump rate, 20% of the pump-head start point, and 80% of the pump-head stop point. The pump start and pump stop timing points are based on the time between two R waves of the ECG, and the pump start and pump stop points are set as a percentage of every circle. In addition to the pump flows, pressure waveforms were recorded at the preoxygenator, postoxygenator, and preaortic cannula sites.
A total of 60 experiments were performed (n = 6 for nonpulsatile and n = 6 for pulsatile) at each of the five flow rates.
In group 1, pump flow = 400 ml/min with nonpulsatile flow (n = 6) and pulsatile flow (n = 6).
In group 2, pump flow = 500 ml/min with nonpulsatile flow (n = 6) and pulsatile flow (n = 6).
In group 3, pump flow = 600 ml/min with nonpulsatile flow (n = 6) and pulsatile flow (n = 6).
In group 4, pump flow = 700 ml/min with nonpulsatile flow (n = 6) and pulsatile flow (n = 6).
In group 5, pump flow = 800 ml/min with nonpulsatile flow (n = 6) and pulsatile flow (n = 6).
Pressure transducers (Maxxim Medical, Athens, TX) were used to monitor the preoxygenator, postoxygenator, and precannula pressures. A Transonic ultrasonic flow probe was placed at the outlet of the oxygenator to monitor the pump flow continuously.
A 20-second segment of the pressure and flow waveforms was recorded for analysis. 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.
A linear mixed-effects model that accounts for the correlation among repeated measurements was fit to the data to assess differences in MAP, EEP, percent change in pressure (from MAP to EEP), and SHE between pump flow rates, perfusion modes, and experimental sites. The between-subjects factors for the model were flow (400, 500, 600, 700, and 800 ml/min) and perfusion modes (pulsatile vs. nonpulsatile). The within-subjects factor for the model was stage (preoxygenator, postoxygenator, and precannula sites). The Tukey multiple comparison procedure was used to adjust p values for post hoc pairwise comparisons. All analyses were performed with the SAS statistical software package (SAS Institute Inc., Cary, NC). All results are expressed as mean ± standard deviation (SD).
Mean arterial pressures were slightly higher in the pulsatile group compared with the nonpulsatile group at preoxygenator, postoxygenator, and precannula sites.
Table 1 represents the MAP results. Figures 3 and 4 represent pulsatile and nonpulsatile pressure-flow waveforms with a pump flow of 600 ml/min.
Energy Equivalent Pressure Levels
Energy equivalent pressure levels were significantly higher in the pulsatile group compared with the nonpulsatile group at all experimental stages at all different pump flows (p < 0.001). Table 2 and Figure 5 describe the results in detail.
Percent Change in Pressure (from MAP to EEP)
Pulsatile flow generated a significantly higher percent of change in pressure compared with the nonpulsatile flow at preoxygenator, postoxygenator, and precannula sites at different pump flow rates (p < 0.001). Detailed results are presented in Table 3 and Figure 6.
Surplus Hemodynamic Energy Levels
Surplus hemodynamic energy levels were significantly higher in the pulsatile group compared with the nonpulsatile group at all pump flow rates at all experimental sites (p < 0.001). Table 4 and Figure 7 summarize the results in detail.
Results of this investigation suggest that by simply changing the nonpulsatile flow to the pulsatile flow at identical experimental settings, it is possible to produce higher hemodynamic energy levels. Our results also confirmed that a membrane oxygenator and the length of the arterial tubing are the main sources for “energy loss” during extracorporeal circulation. Therefore, it is crucial to select suitable circuit components if the pulsatile flow is used. In the past decade, we have tested almost all pediatric heart-lung machines and pediatric oxygenators (FDA approved as well as non-FDA approved) with pulsatile and nonpulsatile perfusion.2,4,6,9–11 Jostra HL-20 heart-lung machine and Capiox Baby RX oxygenator were chosen for this study because of their previous satisfactory performances with either perfusion modes.2,4,11
The main idea behind this investigation was to determine all possible test conditions with pulsatile and nonpulsatile perfusion in a pseudo-patient before using this system clinically. The components of this circuit were identical to our clinical set-up. We will start using this set-up with pulsatile flow in our clinical CPB patients at the Penn State Children’s Hospital.
The difference between the EEP and MAP is the “extra” energy that equals the surplus hemodynamic energy (Figure 1). Under 100% nonpulsatile flow conditions, the difference will be zero. So, there will be no “extra” energy. In this study, however, under nonpulsatile perfusion, particularly at preoxygenator and postoxygenator sites, there was significant surplus hemodynamic energy, because roller pumps (even under 100% nonpulsatile condition) still generate some degree of pulsatility. Therefore, it is even more important to precisely quantify pressure-flow waveforms for direct and meaningful comparisons between perfusion modes. On the basis of our extensive data on this topic, we have repeatedly suggested that investigators who compare pulsatile versus nonpulsatile perfusion should also use EEP and SHE in addition to the pulse pressure for direct and meaningful comparisons.1–6
The pressure transducers used in this study are intended for clinical use. To limit the effect of fluid-filled catheters, the transducers were attached directly to the bypass loop using stopcocks. The transducers have a reported flat dynamic frequency response of more than 100 Hz. We believe that this setup permits the faithful recording and reproduction of pressure waveforms. This system has been used widely in our laboratory to accurately monitor pressures both in vivo and in vitro. No comparisons were made between the pressure transducers used in this study and catheter-tip pressure transducers.
Is there any evidence in the literature that this “extra energy” generated under pulsatile flow conditions has any impact on vital organ recovery in pediatric models? We have already documented that pulsatile perfusion maintains more physiological vital organ function during acute and chronic extracorporeal circulation in animal models as well as in pediatric patients.12–15
The Jostra HL-20 roller pump generated significantly higher EEP, percent change in pressure, and SHE levels in the pulsatile mode when compared with the nonpulsatile mode at all five pump flow rates. SHE and EEP formulas are adequate to precisely quantify pressure-flow waveforms during extracorporeal circulation. Further research to investigate possible mechanisms between extra hemodynamic energy levels and vital organ function is warranted.
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Copyright © 2006 by the American Society for Artificial Internal Organs
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