PERFORMANCE OF THE RHONE-POULENC NON-OCCLUSIVE ROLLER BLOOD PUMP

P.D. Richardson, P.M. Galletti and L.A. Trudell,     Brown University, Providence, R. I. 02912

Publisher Summary

This chapter describes the performance of Rhone-Poulenc nonocclusive roller blood pump. The Rhone-Poulenc blood pump consists of a rotor, motor, support frame, and two pumping tubes, one for venous blood and the other arterial. The variation of volume pumped per revolution of the pump is imitative of Starling’s Law of the heart. According to this law, the volume pumped by the ventricle depends upon the filling pressure available to distend the pump before systole. The outlet pressure has an effect on the performance of the pump. The reason for this is the nonocclusive feature of the rollersystem. The pump outflow can be brought to zero by an outflow pressure that is sufficiently high but that is modest in terms of the risk of rupture of an extracorporeal circuit.

INTRODUCTION

The Rhone-Poulenc blood pump consists of a rotor, motor, support frame, and two pumping tubes, one for venous blood and the other arterial. The motor is totally enclosed, together with reduction gear, so that the rotor is the only exposed mechanically-driven component. The rotor consists of three equi-sized coaxial solid disks with spacers between them, the spacers providing sufficient width for the pump tubes to be stretched flat across the rollers mounted near the periphery of the disks. For each pump tube three 10mm diameter rollers are provided on a pitch circle of 95mm diameter with 120° mutual separation; rollers for the two tubes are staggered at 60° to each other. The disposable pump tubes have an unstretched length of 610mm between the clamping faces of integral collars which fit against the corresponding faces of the yoke that is part of the frame, these yoke faces being 200mm below the rotor axis and holding the tube axes 140mm apart. The pump tubes are produced to have different natural cross-sections when exposed to zero transmural pressure, neither being exactly circular and the venous pump tube more non-circular (i.e., flattened) than the arterial. There are no valves. Both tubes are lined with silicon-free silicone rubber.

PUMP DISPLACEMENT

When the pump tubes have been assembled onto the pump they are somewhat stretched, and also bent where they pass over the rollers. This tends to flatten both tubes further than when they are lying free from the pump. When the rotor rotates the rollers catch a bolus of liquid three times each revolution and carry it round. Even if each roller in rolling against the tube serves as a perfect occluder, the volume displaced in each bolus is limited by the extent to which the tube has filled. This extent depends upon the transmural pressure for the tube, and because the external pressure is atmospheric the size of the bolus depends on the internal (filling) pressure. With the venous pump tube in particular the bolus is small when the filling pressure is sub-atmospheric i.e., below 0 mmHg. As the inlet pressure is raised the effect is to distend the pump tube somewhat so that the volume of the bolus is increased. However, there is a finite range over which the increase of bolus size with increase of inlet pressure is pronounced, because it requires only a finite pressure to distend the pump tube to a circular cross-section. This is achieved first in the region of the pump tube mid-way between the pump rollers. Further gain in bolus volume is achieved by distending the cross-section closer and closer to the rollers which tend to keep the tube flat, and the rate of gain with increase in pressure is relatively small.

The variation of volume pumped per revolution of the pump is imitative of Starling’s Law of the heart. According to this law, the volume pumped by the ventricle depends upon the filling pressure available to distend the pump before systole. This pump characteristic is maintained over the range of pump speeds provided. The manufacturer has limited the maximum rotor speed to approximately 50 r.p.m., and a graph of volume pumped per revolution as a function of filling pressure is relatively invariant with pump speed provided the outlet pressure is low enough, Fig. 1.

EFFECT OF OUTLET PRESSURE

The outlet pressure has an effect on the performance of the pump. The reason for this is the non-occlusive feature of the roller system. The pump tubes simply pass over the rollers with some tension applied because of the tube fixture system. When the outlet pressure is raised sufficiently the tube is held open slightly as the roller passes underneath it. Normally there is a pressure difference between the liquid in the pump tube on one side of a roller and on the other side, with that on the distal side being higher, so that when the tube is held open in this way there is a backflow. The backflow rate depends primarily on the size of the opening and thereby on the outlet pressure. The impact of this on pump performance is more pronounced at lower pump speeds because there is then more time per pump revolution for the backflow to occur. Thus, if the volume pumped per revolution with a constant inlet pressure is plotted as a function of outlet pressure, there is a progressive effect of pump speed, Fig. 2.

The pump outflow can be brought to zero by an outflow pressure which is sufficiently high–– but which is modest in terms of the risk of rupture of an extracorporeal circuit. This outflow pressure rises with the pump speed, but is very much lower than the pressure which can be developed by an occlusive roller pump. This characteristic provides a safety feature in use, but it also limits the maximum pressure against which the pump can deliver a useful blood flow rate.

The pumping characteristics typical of the venous blood pump are repeated by the arterial blood pump in general form, but with some quantitative displacement–– the inlet pressure required to achieve a given degree of filling of the tube is lower for the arterial pump than for the venous pump. This is due to the different cross-sectional shape of the arterial pump tube when there is zero transmural pressure: it is distinctly rounder.

PUMP PERFORMANCE IN AN EXTRACORPOREAL CIRCUIT

In normal use the output of the venous pump is connected to the input of the arterial pump via a blood oxygenator. The inlet pressure of the arterial pump is lower than the output pressure of the venous pump because of the flow resistance of the blood oxygenator. The flow resistance of a blood oxygenator varies depending on its type and on the operating conditions (1,2).

With a bubble oxygenator the input pressure to the arterial pump is usually governed by the level of blood in the arterial reservoir, and this does not vary much. With a membrane oxygenator there is no similar control of the pressure, but it is usually desired to keep the pressure in the blood phase above the gas pressure in the gas phase to avoid the risk of bubbling of gas through pinholes in the membrane. This risk is much greater with a microporous membrane, of course. In steady state operation the flow rates through the venous and arterial pumps must equal each other, and with the capacity of the arterial pump to pump a greater stroke volume for a given inlet pressure one might suspect that the pump-oxygenator-pump system would reach a steady state where the blood pressure at the oxygenator outlet is very low i.e., limiting the overall flow rate by limiting the inlet pressure of the arterial pump. With some membrane oxygenators one finds considerable compliance effects (3), whereby the flow resistance is increased as the oxygenator outlet pressure falls, an effect which would help sustain low inlet pressure to the arterial pump. In fact, the flow limitation arises more from the outflow pressure required from the arterial pump. The pressure which the pump must sustain is not just the subject’s arterial pressure (when on veno-arterial bypass) but also the flow resistance of the blood return line and the cannula. Depending on the cannulation technique and blood flow rate, this may involve an additional 50-300mmHg.

In raising the backpressure on the arterial pump, a “pump-jump” phenomenon can be produced, Fig. 3. For small increases in outlet pressure the blood flow rate decreases slowly, but the backflow leakage increases in the arterial pump and a point can be reached where the arterial pump unloads and actually becomes a resistor. The whole pumping load then falls on the venous pump and the blood flow rate drops. The oxygenator is subjected to the maximum pressure in the system. The condition can be “cured” by increasing the tension on the arterial pump tube at the price of a shorter tube life.