Abstract:
A method is provided of controlling a pump including a electrical motor coupled to a rotor which carries first and second impellers at opposite ends thereof. The method includes: (a) driving the rotor using the motor, so as to circulate fluid from the first impeller through a first fluid circuit, the second impeller, a second fluid circuit, and back to the first impeller; (b) determining a resistance of the first fluid circuit, based on a first motor parameter which is a function of electrical power delivered to the motor; (c) determining a flow rate through the first fluid circuit based on a second motor parameter which is a function of electrical power delivered to the motor; and (d) varying at least one operational parameter of the pump so as to maintain a predetermined relationship between the flow rate and the resistance of the first fluid circuit.

Description:
BACKGROUND OF THE INVENTION 
     This invention relates generally to artificial hearts, and more particularly to methods of control therefor. 
     Heart transplant is a course of action for patients with end stage heart failure, a leading cause of premature death. Because of the unavailability of donor hearts, electromechanical blood pumping systems are being developed and are increasingly coming into use. These devices can provide a bridge to transplant, bridge to recovery, or as a permanent treatment for patients who may not receive a donor heart. Most such patients will be treated with a ventricular assist device (“VAD”), which draws blood from the left or right ventricle, and discharges to the aorta or pulmonary artery, respectively. Some patents require a total artificial heart (TAH) as either a bridge to transplant, or as a permanent therapy. 
     One known type of TAH is a continuous flow total artificial heart (CFTAH) which includes two centrifugal pumps on one rotor supported on a hydrodynamic bearing and driven by a single motor. The CFTAH replaces the ventricles of the heart, and delivers blood flow to both the systemic (left) and pulmonary (right) circulation of the patient. An example of such a CFTAH is described in U.S. Patent Application Publication 2007/0253842. 
     While this type of CFTAH can be operated under external control, there is a need for physiologic control therefor, preferably using the least number of sensors. 
     BRIEF SUMMARY OF THE INVENTION 
     This need is addressed by the present invention, which provides a system and method for physiologic control of an artificial heart. 
     According to one aspect of the invention, a method is provided of controlling a pump including a electrical motor coupled to a rotor which carries first and second impellers at opposite ends thereof, the method including: (a) driving the rotor using the motor, so as to circulate fluid from the first impeller through a first fluid circuit, the second impeller, a second fluid circuit, and back to the first impeller; (b) sensing electrical power delivered to the motor and determining a resistance of the first fluid circuit, based on a first motor parameter which is a function of the electrical power delivered to the motor; (c) determining a flow rate through the first fluid circuit based on a second motor parameter which is a function of the electrical power delivered to the motor; and (d) varying at least one operational parameter of the pump so as to maintain a predetermined relationship between the flow rate and the resistance of the first fluid circuit. 
     According to another aspect of the invention, an artificial heart system includes: (a) an artificial heart including a electrical motor coupled to a rotor which carries first and second impellers at opposite ends thereof, where: (i) the first impeller communicates with a patient&#39;s systemic vasculature; and (ii) the second impeller communicates with the patient&#39;s pulmonary vasculature; (b) a power source; and (c) a controller coupled to the power source and the pump, the controller programmed to: (i) drive the rotor using the motor, so as to circulate blood from the first impeller through the systemic vasculature, the second impeller, the pulmonary vasculature, and back to the first impeller; (ii) sense electrical power delivered to the motor and determine a resistance of the systemic vasculature, based on a first motor parameter which is a function of the electrical power delivered to the motor; (iii) determine a flow rate through the systemic vasculature based on a second motor parameter which is a function of the electrical power delivered to the motor; and (iii) vary at least one operational parameter of the artificial heart so as to maintain a predetermined relationship between the systemic flow and the resistance of the systemic vasculature. 
     According to another aspect of the invention, a method is provided for controlling a pump including a electrical motor coupled to a rotor which carries first and second impellers at opposite ends thereof. The method includes: (a) driving the rotor using the motor, so as to circulate fluid from the first impeller through a first fluid circuit, the second impeller, a second fluid circuit, and back to the first impeller; (b) modulating the speed of the rotor to generate a pulsatile flow; (c) monitoring a motor parameter indicative of suction or rubbing; and (d) in response to an indication of suction or rubbing, reducing the peak motor speed. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       The invention may be best understood by reference to the following description taken in conjunction with the accompanying drawing figures in which: 
         FIG. 1  is a cross-sectional view of an artificial heart and control system constructed according to an aspect of the present invention; 
         FIG. 2  is schematic block diagram of the pump of  FIG. 1  coupled to a circulatory system; 
         FIG. 3  is a graph showing a target operating characteristic; 
         FIG. 4  is a block diagram showing a method of controlling a pump in accordance with an aspect of the present invention; 
         FIG. 5  is a graph showing observed systemic vascular resistance values plotted against a first measured motor power characteristic; 
         FIG. 6  is a graph showing observed systemic flow values plotted against a second measured motor power characteristic; 
         FIG. 7  is a graph showing a control relationship for speed pulsatility; 
         FIG. 8  is a set of graphs showing operation of an artificial heart with modulated speed under normal conditions; 
         FIG. 9  is a set of graphs showing operation of an artificial heart with modulated speed which is experiencing intermittent suction; 
         FIG. 10  is a graph showing normalized current flow versus time in an artificial heart operating at various speeds; and 
         FIG. 11  is a graph of the statistical variation in the normalized current flow shown in  FIG. 10 . 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring to the drawings wherein identical reference numerals denote the same elements throughout the various views,  FIG. 1  depicts a continuous flow total artificial heart  10  used to temporarily or permanently support an human patient. The artificial heart  10  includes a hollow housing  12  with opposed left and right inlets  14  and  16 . An electrical stator  18  comprising a plurality of coil windings is disposed in the housing  12 . While a total artificial heart  10  is used as an illustrative example, the principles of the present invention are equally applicable to other kinds of mechanical configurations and pumps, for example ventricular assist devices. 
     A rotor  20  is disposed inside the stator  18 . The rotor  20  includes a magnet assembly  22  comprising one or more permanent magnets arranged in an annular configuration. A left impeller  24  comprising an annular array of vanes is carried at the left end of the rotor  20  adjacent the left inlet  14 . A right impeller  26  comprising an annular array of vanes is carried at the right end of the rotor  20  adjacent the right inlet  16 . The left and right impellers  24  and  26  discharge into separate right and left peripheral outlets, which are not shown in  FIG. 1 . The left impeller  24  along with the portion of the housing  12  surrounding it may be referred to as a “left pump” while the right impeller  26  along with the portion of the housing  12  surrounding it may be referred to as a “right pump”. 
     All of the portions of the artificial heart  10  which will come into contact with blood or tissue are constructed from known biologically compatible materials such as titanium, medical grade polymers, and the like. 
     The rotor  20  and the stator  18  operate as a brushless DC motor through the application of varying electrical currents to the stator  18 . The artificial heart  10  is coupled by a cable  28  to a controller  32 , which is in turn powered by a power source  30 , for example a battery, both of which are shown schematically in  FIG. 1 . The controller  32  is effective to provide pulsed DC current to the stator  18  in a known manner, and includes a microprocessor or other hardware suited to carry out a preprogrammed control method, as described in more detail below. The controller  32  may include subcomponents such as a CPU or main processor coupled to a known type of motor driver circuit. The degrees of freedom for the controller  32  are mean pump speed (RPM), DC pulse rate, speed pulsatility (i.e. RPM modulation about the mean), and/or duty cycle. The controller  32  is further configured to measure one or more control parameters, in particular electrical power (wattage) delivered to the artificial heart  10 , and to receive a feedback signal from the artificial heart  10  indicative of the pump speed. Speed pulsatility (i.e. RPM modulation) may be used to create a pulse in a patient, and also provide an additional parameter for physiologic control. 
       FIG. 2  shows a simplified schematic of the artificial heart  10  coupled to a patient&#39;s circulatory system. In operation, the left impeller  24  pushes blood through the body&#39;s systemic vasculature, which defines a fluid circuit “S” and is represented from a hydraulic standpoint by a systemic vascular resistance labeled “SVR”. Blood then flows back to the right atrium (right impeller inlet). The right impeller  26  pushes the blood through the body&#39;s pulmonary vasculature, which defines another fluid circuit “P” and is represented from a hydraulic standpoint by a pulmonary vascular resistance labeled “PVR”. Blood flows from the PVR back to the left atrium (left impeller inlet). 
     If the systemic (i.e. left) flow is lower than the pulmonary (i.e. right) flow, then the left atrial pressure increases, and the right atrial pressure decreases. If the left output is greater than the right, then the atrial pressures reverse. Thus, an unbalance in flows is automatically accompanied by an unbalance in atrial (pump inlet) pressures. 
     The magnet assembly  22  in the rotor  20  is axially shorter than the stator  18 , allowing a degree of free axial movement of the rotor  20 , in response to any unbalance of pump inlet (i.e. atrial) pressures. This axial movement changes the distances “D 1 ” and “D 2 ” (see  FIG. 1 ) which represent axial operating clearances of the left impeller  24  and right impeller  26 , respectively. This change in pump geometry affects the relative left/right performance in a direction to correct the atrial pressure imbalance. Thus, the artificial heart  10  is self-balancing, acting as an inlet pressure balancing regulator while at the same time pumping both systemic and pulmonary circulation. 
     The artificial heart  10  is controlled as follows. First, a desired or targeted characteristic is determined by a physician. The characteristic describes the relationship between the volumetric flow rate in the systemic vasculature S and the SVR. In the example shown in  FIG. 3 , the characteristic is a linear plot between endpoints selected by the physician. 
     Referring to  FIG. 4 , the controller  32  delivers power to the artificial heart  10  to spin the left and right impellers  24  and  26 . The speed of the rotor pulse may be modulated in order to create a pulse in the patient. As used herein, “modulation” refers generally to any change in a cyclic property of the rotor speed, whether this change is made directly or indirectly, and may be accomplished by various means. For example, direct closed loop control of the rotor speed may be implemented. Alternatively, the electrical current supplied to the motor from the controller  32  may be modulated, and the resulting changes in rotor speed accepted. The rotor modulation signal (i.e. the speed or current wave form) may be sinusoidal, or a sine wave with a duty cycle transformation, or other wave forms such as ramp, triangular, or square. At block  100 , the controller  32  senses the average electrical power (i.e. wattage) delivered to the motor, which can be measured by the controller  32  in a known manner, and the rotor speed, based on a feedback signal from the motor. 
     Next, at block  110 , the controller  32  computes two parameters: PSnorm, which is defined as average Watts divided by kRPM 3 , and PQnorm, defined as average Watts divided by kRPM 2 . Because these parameters are derived from or related to the power delivered to the motor using mathematical expressions, they may be described as being “functions of” power delivered to the motor. 
     Next, at block  120 , the SVR and the systemic flow rate are determined based on the computed parameters.  FIG. 5  shows a sample of a suitable correlation for SVR (in dyn·s/cm 5 ) vs. PSnorm. The correlation is derived from empirical test data. It is depicted as a graph in  FIG. 5 , but it may be implemented or stored by the controller  32  in any equivalent fashion, for example as a graph, as a lookup table or matrix, or as a mathematical expression (e.g. a linear or polynomial curve fit). 
       FIG. 6  shows a sample of a suitable correlation for systemic flow rate (in 1 pm) vs. PQnorm. The correlation is derived from empirical test data. It is depicted as a graph in  FIG. 6 , but it may be implemented or stored by the controller  32  in any equivalent fashion, for example as a graph, as a lookup chart or matrix, or as a mathematical expression (e.g. a linear or polynomial curve fit). 
     Once the systemic flow rate and SVR have been determined, their relationship can be computed to determine if the current operating point lies on the prescribed characteristic shown in  FIG. 3 . If it does not, then one or more operational variables are increased or decreased until it does, at block  130 . Examples of such operation variables include mean pump speed, pulse rate, speed pulsatility, and/or duty cycle. The process repeats at block  100  so long as pump operation continues. 
     Independent of the control process, the self-balancing process described above is also taking place during operation of the artificial heart  10 . The relative left/right performance of the two pumps can be further affected by the relative impedance seen by the pump outputs. In this pump configuration, speed modulation at high SVR can decrease the left pump output while increase the right pump output. This effect can be moderated by using the controller  32  to reduce or eliminate any speed pulsatility at high SVR values. For example, the controller  32  may be programmed to follow a characteristic of enforced speed pulsatility vs. SVR similar to the example shown in  FIG. 7 . This allows the artificial heart  10  to self-balance over a wider range of physiologic conditions, allowing an additional degree of freedom in balancing left/right performance. 
     Operating the artificial heart  10  in a modulated mode can cause an intermittent suction of tissue around one of the left or right inlets  14  or  16  at the cyclic peak speed with physiologic decreases in blood volume returning to the artificial heart  10 . This intermittent occlusion can cause erratic and amplified oscillation in axial movement of the rotor  20 , and touching of the left or right impellers  24  or  26  against the pump housing  12 , which is reflected in the speed and current signals processed in the controller  32 . Intermittent suction and the associated effects are undesirable and can cause excessive wear or damage to the artificial heart  10 . In addition to control of the artificial heart  10 , the present invention provides a method for detecting this intermittent suction and responding to it through peak speed reduction. 
     Normally, when the speed is a sinusoidal speed wave form, this will yield a current of a similar wave form, and vice-versa. An example of normal system response is shown in  FIG. 8 , showing the sinusoidal speed and a near sinusoidal current. In this example, the pump speed was modulated at frequency of 1.33 Hz (i.e. 80 beats per minute), and an amplitude of +/−25% of the mean speed of 2800 RPM. In  FIG. 9 , the same modulation at a mean speed of 3000 RPM resulted in intermittent suction, with a non-sinusoidal speed/current relationship (indicated at “A” and “B”), and spikes in the rate of current change (di/dt), clearly indicating the point of rubbing (high positive di/dt) at “C” and period of suction (high negative di/dt) at “D”. Based on these observations, a high absolute value of (di/dt) can serve as a predetermined limit or trigger for a response. The limit value for a particular application may be derived from empirical test data. Alternatively, the controller  32  may be programmed to evaluate the speed or current wave forms and trigger a response based on the existence of particular characteristics in the wave forms, such as the brief RPM drops indicated at A in  FIG. 9 , or the peak discontinuities shown at B in  FIG. 9 . 
     Another suitable test for triggering peak speed reduction is analysis of normalized current.  FIG. 9  shows normalized current (where the current signal is normalized by dividing by the cube of the speed signal) for an artificial heart  10  operating at 2600 RPM, 2800 RPM, and 3000 RPM. Normally, the resulting signal of current divided by speed cubed has small variation (see  FIG. 11 ). However, when the speed is high enough to cause an abnormal suction condition, the variation in normalized current is dramatically higher. This is clearly seen in the example of operation at 3000 RPM. Rubbing is indicated at “C′” and suction at “D′” in  FIG. 10 , with a correspondingly high variation in normalized current in  FIG. 11 . Based on these observations, a high value of variation in normalized motor current can serve as a predetermined limit or trigger for a response. The limit value for a particular application may be derived from empirical test data. 
     Regardless of which specific trigger or limit value is used to determine the presence of suction and/or rubbing, the controller  32  may be programmed to evaluate the parameter and look for the specified trigger and/or compare the parameter to the predetermined limit. If suction or rubbing is indicated, the controller  32  responds by reducing the peak speed. This can be done by lowering the mean speed, reducing the speed modulation amplitude, changing the duty cycle (portion of time at high speed), or any combination thereof. 
     The foregoing has described a method of operating a total artificial heart. While specific embodiments of the present invention have been described, it will be apparent to those skilled in the art that various modifications thereto can be made without departing from the spirit and scope of the invention. Accordingly, the foregoing description of the preferred embodiment of the invention and the best mode for practicing the invention are provided for the purpose of illustration only and not for the purpose of limitation.