The Twin-Pulse Life Support System (T-PLS) is a novel pulsatile extracorporeal life support system (ECLS) developed in Korea.1,2 This new device received Korea Food and Drug Administration approval in 2004 and has been used in clinics to save patients’ lives. It has an actuator and two blood sacs, and the reciprocating actuator pushes blood sacs alternatively. The unique feature of T-PLS is in the pulsatile flow generation compared with other ECLSs. Recent studies show that the T-PLS achieves higher levels of tissue perfusion of the kidney during short-term extracorporeal circulation and provides more blood flow to coronary artery than a nonpulsatile blood pump.3,4 The mean perfusion pressure (48.7–52.0 vs. 39.8–45.5 mm Hg; p = 0.023–0.48) and the tissue perfusion flow of the kidney (65.8–88.3 vs. 5–64 ml·min–1·100 g–1; p = 0.026–0.45) were higher in T-PLS than in a nonpulsatile centrifugal pump during cardiopulmonary bypass.3 The blood flow and velocity of the left anterior descending coronary artery were approximately 1.5 times and 2 times higher in T-PLS than in nonpulsatile centrifugal pump, respectively.4
However, these results lack the quantification of pulsatility and thus make it hard to analyze the relationship between pulsatility and its effects on hemodynamic performances. There has long been controversy regarding the clinical benefits of pulsatile and nonpulsatile devices,5–9 but the lack of the precise quantification of pulsatility made it hard to compare benefits of pulsatile and nonpulsatile flow. Thus, in these debates, the quantification of pulsatility is essential.10 Although there is no universally adopted quantification method, the energy equivalent pressure (EEP) suggested by Undar et al. is thought to be adequate to quantify pressure-flow waveforms of different modes of flow, because EEP takes hemodynamic energy into account in describing the quantification of pulsatility.10–14
The purpose of this study was to analyze pulsatility performance in vitro according to circuit configurations and to suggest an optimal circuit configuration that can produce the highest performance. The evaluation of pump performance is held in a mock circulation system that can offer stable hemodynamic conditions. We set up three different circuits with T-PLS and a membrane oxygenator. In all cases, the EEP and surplus hemodynamic energy (SHE) quantifications were studied.
Materials and Methods
Device and Circuit Setups
The T-PLS (NewHeartBio, Seoul, Korea) is a pulsatile ECLS with two blood sacs and a reciprocating actuator (Figure 1). The reciprocating actuator pushes two sacs alternatively and produces outflow twice per beat. A membrane oxygenator (MO) is accompanied with the T-PLS and is located at the inlet or outlet site of the T-PLS according to the circuit setup.
Three different circuit setups were considered: serial, parallel A, and parallel B types (Figure 2). In the serial type circuit setup, the MO and T-PLS are connected serially. The blood flow drained to the T-PLS is pumped out to the MO and returned to another T-PLS blood sac, and the blood is then pumped out to the artery. In the parallel A type, two blood sacs are connected parallel; the blood flow enters the MO first, passes through blood sacs, and is pumped out to the artery. Similarly, in the parallel B type, the blood flow passes through blood sacs first, and then is pumped out to the MO. In this circuit setup, the blood flow pumped by the T-PLS passes the MO before entering the artery. Thus, the pulsatility generated by T-PLS is reduced by passing through the MO. Thus, these differences in circuit setup show the different performances in pulsatility.
In Vitro Test and Precise Quantification of Pressure-Flow Waveforms
The performance of the T-PLS was tested in a mock circulation system with three circuit setups. In serial and parallel B types, a closed-type MO (Capiox S10, hollow fiber membrane oxygenator, Terumo Corp., Tokyo, Japan) was used. In the parallel A type, a low-resistance type MO (Capiox CX230, hollow fiber membrane oxygenator, Terumo Corp.) was used. A blood analog fluid consists of 40% glycerin and 60% water. The pressures were measured at the site of the arterial, venous compliance chamber with pressure transducers (PSHBC1250, Sensys, Busan, Korea). The flow rate was measured at the output cannula connected to the arterial compliance chamber with an ultrasonic flow meter (TS410 flow meter, ME11PXL probe, Transonic, Ithaca, NY). The pressure and flow waveforms were acquired with a data-acquisition board, and MATLAB (Mathworks, Natick, MA) software was used to analyze the waveforms.
In each circuit setup, the T-PLS was manipulated with 30, 40, and 50 bpm conditions and with afterloads (mean aortic pressure) of 80, 90, 100, 110, and 120 mm Hg, respectively. The aortic pressure was regulated by varying the systemic resistance of the mock circulation system. A 30-second waveform of pressure and flow rate was acquired; each measurement was repeated six times.
The following EEP formula for quantification of pulsatility uses pressure and flow waveforms:
Where f is the pump flow rate (l/min), p is the arterial pressure (mm Hg), and dt is the increment in time.10,13 The SHE formula is as follows:
Where MAP is the mean aortic pressure (mm Hg).10
The pressure and flow waveforms in serial, parallel A, and parallel B circuit setups are shown in Figures 3–5, respectively. Each setup made a different shape of pulses. In serial circuit setup, unlike in parallel setups, one pump beat makes a big pulse and a smaller one (Figure 3). A big pulse is made when the actuator of the T-PLS pushes the blood sac connected to the output cannula; a smaller one is formed by indirect flow that passes the MO and a blood sac to output cannula when the actuator pushes another blood sac. However, in parallel A and B circuits, both blood sacs are connected to output cannula through an MO or not (Figure 2b,c). Thus, one reciprocating movement of actuator produces double pumping out of the blood.
The percent changes from MAP to EEP and the pump outputs with varying MAP and pump rate are plotted in Figures 6–8. With different circuit setups, the T-PLS showed different pulsatility performances to the same afterloads (MAPs) and pump rates. The average percent changes from MAP to EEP were 13.2% ± 3.2%, 10.0% ± 1.6%, and 7.0% ± 1.1% with serial, parallel A, and parallel B circuit setups, respectively. The average SHE levels were 17,404 ± 3,750 ergs/cm3, 13,170 ± 1,486 ergs/cm3, and 9,192 ± 1,122 ergs/cm3 in each circuit setup. The pump outputs were 1.87 ± 0.29 l/min, 3.09 ± 0.75 l/min, and 3.06 ± 0.56 l/min in serial, parallel A, and parallel B circuit setups, respectively. The detailed results are summarized in Table 1 and 2 with different circuit setups and pump rates.
The dependency of the percent change on MAP and pump rate is shown in all circuit configurations. The percent changes are reduced with increasing MAP and decreasing pump rate. However, the EEP is higher in the serial setup than in parallel setups. In parallel setups, although each produces a similar pump output, the EEP is higher in the parallel A setup than in the parallel B setup.
These results show that a particular blood pump could generate different pulsatility and different shapes of pulse according to circuit configurations. The EEP changes of the serial and parallel A circuits are 13.2% ± 3.2% and 10.0% ± 1.6%; these are higher than 10% in the normal human heart.14 The difference between parallel A and B circuits is the position of the MO. In the parallel A circuit, the MO is placed in a prepump position, thus the blood outflow is directly connected to output cannula. However, in the parallel B circuit, the MO is placed at a postpump position, thus the pulsatility of the T-PLS is diminished to 7.0% ± 1.1% by the MO.
Parallel A and B circuits produce higher pump outputs than the serial circuit because the reciprocating actuator pumps out twice per one beat operation. As a result, the parallel A circuit achieves higher pump output and generates acceptable pulsatility, as with the natural heart. Considering such factors, we suggest that the optimal circuit configuration is parallel A. However, when the T-PLS circuit is configured as parallel A type, the type of MO must be considered. To make flow into the T-PLS sufficient without a high pressure gradient for drain, a low-resistance MO is needed.
In a particular circuit setup, the percent changes of EEP are varied with afterloads and pump rates. They are decreased as afterload is increased and pump rate is decreased. However, the SHE levels were kept nearly constant regardless of afterloads in a same pump rate. The reason for this is thought to be that the energies generated by the T-PLS are similar with various afterload conditions unless the pump rate is changed. Because the SHE shows less dependency on afterloads, it would be more desirable than EEP in defining the pulsatility performance to various afterloads, although it is not intuitive.
The EEP and SHE are useful tools in the quantification of a blood pump’s pulsatility, and are expected to be used as additional performance indexes to define the performance of blood pumps (e.g., pump output, pump rate). Moreover, these in vitro tests are suitable to define pulsatility because the mock circulation system offers stable hemodynamic conditions to a blood pump. Further in vivo studies are planned to verify these in vitro results.
This study was supported by a grant from the Korea Health 21 R&D Project, Ministry of Health & Welfare, Republic of Korea (02-PJ3-PG6-EV09-0001), and by the Brain Korea 21 Project of the Ministry of Education and Human Resources Development, Republic of Korea.
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