Blood flow assistance device and associated control system

Abstract

A control system 202, for a blood assistance device comprising a pump 206, the control system configured to control operation of the pump on a respiration cycle having an inspiration and an expiration

Description

The invention relates to the field of control systems for blood flow assistance pumps, and in particular, although not exclusively, to cardiac pumps that are configured to provide a respiratory sinus arrhythmia (RSA). Cardiac pumps may be used to produce cardiac activity in a subject in which the normal biological systems for stimulating such activity have failed. This is particularly relevant for patients with late-stage heart failure where one or more heart structures have become damaged, atrophied, or weakened, in which case the function of the heart needs to be artificially supplemented or replaced. According to a first aspect there is provided a control system for a blood flow assistance device comprising a pump, the control system configured to control operation of the pump based on a respiration cycle having an inspiration phase and an expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein the first mode of operation is different to the second mode of operation. The control system may be configured to pulse the pumping operation. The control system may be configured to pulse the pumping operation having a respiratory sinus arrhythmia. A frequency of the pulses may be greater during the first mode of operation in the inspiration phase than during the second mode of operation in the expiration phase. For example, a frequency of pumping in the inspiration phase (finsp) may be greater than a frequency of pumping in the expiration phase (fexp). The control system may be configured to synchronize the pulsing of the pumping operation with the onset of the inspiration phase or the expiration phase. The control system may be configured to generate the timing of the pumping pulses as a function of a respiration duty cycle in order to maintain a bias towards synchronization between the respiration period and an integer ratio of the periods between pumping pulses. To this end, the control system may comprise apparatus for determining timing of electrical stimulus signals temporally modulated by a respiration signal, comprising: a first input stage configured to receive a first input signal indicative of respiration; a respiration analysis module configured to determine, from the first input signal, a signal indicative of an instantaneous (or current) respiration duty cycle; and a synchronization module configured to generate the timing of the stimulus signals as a function of the signal indicative of respiration duty cycle in order to maintain a bias towards synchronization between the respiration period and an integer ratio of the periods between stimulus signals. The electrical stimulus signals may be cyclical electrical stimulus signals. The integer multiple of the stimulus signal periods may be a predetermined number. The synchronization module may modulate the timing of the stimulus signals according to a non-linear function. The synchronization module may comprise a neuronal oscillator. The timing of the stimulus signals may comprise modulating a base frequency for the heart. The neuronal oscillator may comprise a single neuronal oscillator. That is, the neuronal oscillator may comprise only one neuron or a system emulating only one neuron. The apparatus may further include a non-linear oscillator. The non-linear oscillator may be configured to receive a second input indicative of the base frequency for the heart, for example, the current stimulation lexp that determines fexp, and the signal indicative of respiration duty cycle. The non-linear oscillator may synchronize to the signal indicative of respiration duty cycle. The current lexp may be set, by a clinician for example, to provide either a fixed heart rate or made to vary with the respiration rate according to a set calibration curve. In the latter, an averaged respiration rate, which may be averaged over the last 5 respiration cycles for example, may be used as a measure of physical activity. The respiration rate may be read by a frequency counter and a digital or analogue amplifier that generates a current proportional to this average frequency. The respiration duty cycle may have an inspiration phase. The respiration duty cycle may have an expiration phase. The apparatus may further include means for setting a parameter (for example, RSA or (finsp-fexp) / fexp) determining a differential stimulus signal rate between the inspiration and expiration phases. Each of the periods between stimulus signals within a single inspiration phase or expiration phase may have the same target duration. Each of the stimulus signal periods within a single inspiration phase or expiration phase are intended to be of the same duration but may depart from this state by small deviation to maintain synchronization to respiration. Maintaining a bias towards synchronization between the respiration period and an integer ratio, or integer multiple, of the periods between stimulus signals may comprise setting a period for the periods between stimulus signals in the inspiration phase or expiration phase. The synchronization module may be configured to maintain a bias towards synchronization between the inspiration phase or expiration phase of the respiration period and an integer ratio, or integer multiple, of the periods between stimulus signals. A different integer number of periods between stimulus signals may be provided in the inspiration phase and the expiration phase. The apparatus may further including means for setting a strength of coupling factor. Said means may determine (i) a speed of synchronization between the non-linear oscillator and the respiration duty cycle signal and / or (ii) a tolerance to frequency mismatch between the respiration period and one or more target heart beat intervals. The apparatus may comprise an analogue electronic signal processing chain. The electronic signal processing chain may provide the respiration analysis module. The electronic signal processing chain may provide the synchronization module. The apparatus may comprise a blanking module. The blanking module may be configured to provide a blanking period in the first input signal indicative of respiration based on i) the timing of the stimulus signals or ii) based on detection of stimulus signal interference in the first input signal. The control system may be further configured to provide counterpulsation. The control system may be configured to provide the counterpulsation by timing the pumping pulses such that the pumping pulses are applied in synchrony and out of phase with a heartbeat cycle. The control system is configured to pulse the pumping operation such that pulses are applied during a diastole phase of the heartbeat cycle. The control system may be configured to modulate the pumping operation to provide opening the aortic valve while at the same time providing increased pulse pressure. The control system may further be configured to modulate the pumping operation to provide ventricular unloading and increased myocardial supply / demand ratio. The pumping operation may be continuous. The pump may have a higher flow rate during the first mode of operation than during the second mode of operation. Controlling operation of the pump may comprise modulating an instantaneous power supplied to the pump by the control system. The instantaneous power may be greater during the first mode of operation than during the second mode of operation. The control system may be configured to receive a signal indicative of a feature of the respiration cycle and to control operation of the pump in accordance with the feature. The feature may be the onset or end of the inspiration phase or of the expiration phase. The signal may be indicative of an instantaneous respiration cycle of a subject. The control system may be further configured to determine, from the respiration cycle, an respiration duty cycle. The control system may have a memory and is configured to retrieve data describing the respiration cycle from the memory. The data may be saved subject data for a subject having a respiratory sinus arrhythmia. The respiration cycle may be a synthetic respiration cycle. The control system may be configured to generate the synthetic respiration cycle having a respiratory sinus arrhythmia. According to a further aspect there is provided a blood flow assistance device controller comprising: a control system for a blood flow assistance device comprising a pump; a power source electrically connected to the control system; a housing, wherein the power source and the controller of the control system are provided within the housing; According to a further aspect there is provided a blood flow assistance device comprising: a pump comprising an inlet and an outlet; a control system for a blood flow assistance device comprising a pump, or the blood flow assistance device controller defined above, electrically connected to the pump. The blood flow assistance device may further comprise a sensor for determining a feature of a respiratory cycle of a patient. The blood flow assistance device may comprise: an LVAD device; an RVAD device; an artificial heart a perfusion pump; or a cardiopulmonary bypass machine According to a further aspect there is provided a control system for a blood flow assistance device comprising a memory, the control system configured to: retrieve data describing the respiration cycle from the memory; and control operation of a pump based on the respiration cycle having an inspiration phase and an expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein the first mode of operation is different to the second mode of operation. According to a further aspect there is provided a non-transitory computer-readable storage medium comprising computer program code configured to cause a processor to execute a method for controlling operation of a pump based on a respiration cycle having an inspiration phase and an expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein the first mode of operation is different to the second mode of operation. Brief Description of Figures Embodiments of the present invention will now be described by way of example and with reference to the accompanying drawings in which: Figure 1 illustrates a schematic block diagram of a typical blood flow assistance device comprising a pump; Figure 2 illustrates a schematic block diagram of a blood flow assistance device comprising a pump; Figure 3a is a profile of calcium ion (Ca2+) concentration observed when monotonic pacing is applied to a heart. Figure 3b is a profile of calcium ion (Ca2+) concentration observed when RSA pacing is applied to a heart. Figure 4 illustrates a schematic block diagram of an example of a control system for a blood flow assistance device comprising a pump; Figure 5 illustrates a schematic block diagram of an alternative example of a control system for a blood flow assistance device comprising a pump; Figure 6 illustrates a schematic block diagram of an example of a blood flow assistance device controller; Figure 7 illustrates a schematic block diagram of an example of a blood flow assistance device comprising a pump, comprising an inlet and an outlet, and a control system for the blood flow assistance device; Figure 8 illustrates a schematic block diagram of an example of a blood flow assistance device comprising a pump, comprising an inlet and an outlet, and a blood flow assistance device controller; Figure 9 illustrates a schematic block diagram of a system comprising a control system for a blood flow assistance device, a cardiac pump, and a cardiac pacemaker; Figure 10 is a graph showing the dependence of frequency of oscillation of the neuron membrane voltage as a function of injected current; Figure 11 illustrates an example implementation of a neuronal oscillator. Description The human heart pumps blood around a body in a pulsatile manner. Notably, the body’s natural regulation of heartbeat is in phase with breathing or respiratory cycle. For example, healthy hearts in humans and other mammals slow down during expiration (breathing out) and speed up during inspiration (breathing in). This is referred to as respiratory sinus arrhythmia (RSA). The loss of RSA is a predictor of cardiovascular risk and a prognostic indicator for multiple diseases including heart failure. Left-sided heart failure is diagnosed when the left ventricle cannot effectively pump blood to the body, and right sided heart failure is diagnosed when the right ventricle cannot effectively pump blood to the lungs. Heart failure may be systolic, where the cardiac muscle cannot contract forcefully enough to pump the blood out of the ventricles, or diastolic, where the cardiac muscles are too stiff and cannot relax enough to allow adequate blood to enter the ventricles. Patients with systolic heart failure will typically require systems or devices for artificially supplementing cardiac function to restore adequate blood flow until a transplant can take place. Examples of devices typically used for treatment of systolic heart failure include Ventricular Assist Devices (VADs), such as a Left Ventricular Assist Device (LVAD). Such devices use cardiac pumps to assist the heart in pumping blood around the body. Another application of blood flow assistance pumps is the use of perfusion pumps for maintaining consistent blood flow through an organ being transported or held for organ transplantation. The hemodynamic effects of RSA are associated with better patient outcomes. However, heart failure patients typically exhibit weaker RSA. Furthermore, the pumps employed in assistance devices, such as LVADs, typically provide a linear or continuous flow. In various examples in the present disclosure, a control system is provided in order to artificially restore blood flow using a cardiac pump in a way that is hemodynamically similar to the blood flow produced by a heart with RSA. The control system may be used to slow down disease progression and improve cardiomyocyte structure and function whilst an organ is being assisted by the cardiac pump in order to improve quality of life outcomes for the subject of the cardiac device. For profusion device, the provision of the disclosed methods aims to improve organ function post-transplant and therefore better outcomes for patients, reduce rejection rates or reduce the need for rejection medication. It is contemplated that this applies to cardiac transplants and other types of organ transplants. The vasculature of all organs would likely benefit from a more physiological perfusion. Figure 1 illustrates a blood flow assistance device 100, such as a Left Ventricular Assist Device (LVAD). The device 100 comprises a control system 102 and a pump 106. The control system 102 is configured to control the power supplied to the pump 106. The power supplied to the pump 106 may be adjusted, for example, to account for changes in blood pressure, or to ensure that the pump is operating efficiently. The continuous, linear blood flow provided by continuous flow cardiac pumps may contribute some of the complications that have restricted the use of these devices, including: aortic valve insufficiency, gastrointestinal bleeding, stroke, pump thrombosis, and hemolysis. Cells lining the cardiovascular system, such as endothelial cells lining the blood vessels, are known to be sensitive to flow dynamics of the blood. The morphology of endothelial cells changes depending on whether they are cultured in static or flowing media. The shear stress caused by the liquid causes the cells to align with the direction of flow. This can contribute to endothelial cell damage or dysfunction. Turbulent flow and eddies are known to contribute to atherosclerotic plaque development around bifurcations in the arterial system. it has been found that the effect of modulating the pump speed is different depending on the timing of the Continuous Flow LVAD (CFLVAD)’s speed modulation relative to native ventricular contraction. For example, LVADs may employ a counterpulsation or copulsation mode, speed modulation systems, pulse shape modulation, or pulse width modulation. Compared with constant speed operation, counterpulsation operation increases left ventricular unloading and the myocardial supply / demand ratio; however, it decreases pulse pressure. Copulsation operation, on the other hand, increases pulse pressure but decreases left ventricular unloading and the myocardial supply / demand ratio. In addition, the aortic valve is difficult to open for copulsation because arterial pressure increases during systole when the aortic valve starts to open. Other types of speed modulation may allow the aortic valve to open in an effort to prevent fusion and insufficiency of the aortic valve. None of the above mentioned speed modulation methodologies can provide everything — ventricular unloading, myocardial supply / demand ratio, and opening of the aortic valve while at the same time providing increased pulse pressure. Healthy physiological cardiac function provides not only pulsatility to blood flow but aligns this variability with the respiratory cycle in a mechanism known as respiratory sinus arrythmia (RSA). The inventors have identified that restoring this respiratory modulation to the blood flow may benefit the cardiorespiratory system including positive changes in cardiomyocyte morphology and protein expression and structure, as well as improving endothelial function. In this way, artificial cardiac pump function can be improved through modulation of its pumping function by the respiratory cycle to restore an RSA-like rhythm to the blood flow. The benefits of restoring RSA to the blood flow include improved function of remaining cardiomyocytes, improved endothelial function and therefore reduced incidence of thrombosis and atherosclerotic plaque development, and improved blood pressure control. These benefits may therefore reduce incidence of stroke, myocardial infarction (Ml) or aneurysm in patients with heart failure. Figure 2 illustrates a schematic block diagram of a blood flow assistance device 200 comprising a control system 202 and a pump 206. The control system (or controller) 200 is configured to control operation of the pump 206 of the blood flow assistance device 200 based on a respiration cycle. The controller 202 may be implemented using various combinations of hardware and software. The respiration cycle has an inspiration phase and an expiration phase, and the controller 202 is configured to change a mode of operation of the pump 206 between a first mode that is operated during the inspiration phase and a second, different, mode that is operated during the expiration phase. The pump 206 may be implantable into or disposed external to a patient or subject. The controller 202 may be electrically connected to the pump 206 and may communicate with or control the pump 206 using one or more electrical signals over a wired or wireless communications link. For example, the pump may be implanted into a chest cavity of a patient for fitting on or around a heart, and the controller 202 may communicate with the implanted pump via a driveline electrically connected to the controller 202 and the pump 206. The control system 202 may restore the effects of RSA on blood flow provided by the pump 206 of the blood flow assistance device by operating the pump in a first mode of operation during an inspiration phase of the respiration cycle and a second mode of operation during an expiration phase of the respiration cycle. In other words, the control system 202 may restore the effects of RSA on blood flow provided by the pump 206 of the blood flow assistance device by synchronizing the blood flow provided by the pump with the phase of the respiration cycle. In some embodiments, the pump 206 may be a pulsatile pump, and the control system 202 may be configured to pulse the pumping operation of the pump 206 with respiratory sinus arrythmia such that the blood flow provided by the pump 206 exhibits RSA-like pulsatility. The effects of RSA may be restored by having a greater frequency of pumping pulses during the first mode of operation than during the second mode of operation. The control system 202 may be configured to synchronize the pulsing of the pumping operation with the onset of the inspiration phase or the expiration phase. The control system 202 may be configured to generate the timing of the pumping pulses as a function of a respiration duty cycle in order to maintain a bias towards synchronization between the respiration period and an integer ratio of the periods between pumping pulses. In some examples, the timing of the pulsing is generated according to a non-linear function which may be provided by, for example, a neuronal oscillator. The neuronal oscillator may be provided by analogue circuitry implementing, for example, a Central Pattern Generator, or by a digital simulation. A neuronal oscillator is only an example of a non-linear oscillator. Any nonlinear oscillator would synchronize to biological rhythms. The neuronal oscillator is a model of the oscillator living systems use. A neuronal oscillator generates "spiky" pulses whose width and period may be tuned independently to meet the specifications of pulse width (typically 1ms for stimulating the sinoatrial node) and heart pacing frequency. Spike-based neuromorphic models communicate using the same type of communication as biological systems. A digitally equivalent pacing device may digitally generate the neuronal membrane voltage by solving the Hodgkin Huxley equations (or the mathematical model of the hardware which predicts nearly identical membrane voltage oscillations) to replicate the nonlinear properties to achieve synchronization. This model is given by the following equations: The state variables of this model are (V,m,h,n) respectively the membrane voltage, Na activation gate variable, Na inactivation gate variable, K activation variable. Iinj is the injected current. An exemplar set of parameters for this model is as follows: Ion channel Parameter IQ Value C (pF cm T 1 ISAM ip x ir Fast and transient Sodium current (NaTj 69 .................MM................ 41 .................MM................ -302 ...................................................... 10 .................................. 0.143 ..................MM.................. 1 099 ......................... 23.39 MM -65.37 -17.63 0.701 ...................&.(M.................. 12.9 27.22 Transient depolarisation activated current (K) ...................... 5.9 MM -100 ....................................MM................................... -34.53 .......................... 2339 22 17 MM 1.291 MM 4,314 Leak current (Q 0.465 -65 A neuronal oscillator has a frequency that increases as a function of current injection as shown in Figure 10. Therefore, the strength of coupling, or RSA = (finsp-fexp) / fexp, is set 5 by varying the amount of current injected in the neuron during the inspiratory part of the cycle (linsp) relative to the current applied in the expiratory part of the cycle (lexp). The current difference (linsp-lexp) may be set, by a user for example, to control the strength of coupling. 10 Figure 10 shows the dependence of frequency of oscillation of the neuron membrane voltage as a function of injected current: f(l) from a numerical simulation of pacemaker oscillations. RSA is generated by injecting in the neuron a rectangular current signal with two levels linsp and lexp which set frequencies finsp and fexp. Both linsp and lexp are greater than the threshold Ith. Because the neuron oscillation frequency increases with 15 current injection, the strength of RSA = (finsp-fexp) / fexp is increased or decreased by increasing Lisp relative to Ith. The main source of nonlinearity in the neuron response is the frequency-current dependence illustrated in Figure 10. In biological implementations, this nonlinear dependence is underpinned by the sigmoidal activation and inactivation curves of sodium and potassium ion channels in the neuron membrane, each of which has an activation threshold. The dynamics of sodium and potassium ionic currents may be modelled by the neuron electronics. Figure 11 illustrates an example implementation of a neuronal oscillator 1500. The neuronal oscillator 1500 is provided by analogue electronics configured to mimic the behaviour of a neuron. The neuronal oscillator 1500 has a potassium ion channel section 1502, a sodium ion channel section 1504 and a membrane section 1506 to model the corresponding ion channels of a neuron cell. The neuronal oscillator 1500 provides a membrane voltage as an output on an output rail 1501 The potassium ion channel section 1502 comprises a plurality of field effect transistors (FETs), a plurality of variable voltage sources, a starter circuit 1507 and a timing capacitor 1508 determining the recovery rate of the potassium activation gate. The plurality of FETs comprises first, second, third, fourth, fifth, sixth, seventh, eighth, ninth, tenth, and eleventh and twelfth FETs 1510 -1532. Each FET 1510-1532 has a gate, a source and a drain. The variable voltage sources, which may be used to set the conditions of the neuron, may be provided by a potential divider formed by a potentiometer. These voltage sources are used to set the voltage thresholds Vt Vat, Vt, and conductances mentioned in the Table above. The gate of the first FET 1510 is coupled to a voltage source 1511. The source of the first (n-type) FET 1510 is coupled to ground, the drain of the first FET 1510 is coupled to the source of the second (n-type) FET 1512, to the source of the third (n-type) FET 1514 and the starter circuit 1507. The starter circuit 1507 comprises an RC timing circuit provided by a resistor and a capacitor in series between the drain of the first FET 1510 and ground. The gate of the second FET 1512 is coupled to the rail 1501 (the membrane voltage V.) The source of the fourth (p-type) FET 1516 and the source of the fifth (p-type) FET 1518 are coupled to the positive voltage source. The gate of the fourth FET 1516 is coupled to the gate of the fifth FET 1518. The drain of the fourth FET 1516 is coupled to the gate of the fourth FET 1516 and the drain of the second FET 1512. The drain of the fifth FET 1518 is coupled to the drain of the third FET 1514. The gate of the third FET 1514 is coupled to the drain of the third FET 1514 and the gate of the sixth (n-type) FET 1522. The timing capacitor 1508 is coupled between the gate of the third FET 1514 and ground. The source of the twelfth (n-type) FET 1520 is coupled to ground. The gate of the twelfth FET 1520 is coupled to second voltage source 1534. The drain of the twelfth FET 1520 is coupled to the source of the sixth (n-type) FET 1522 and the source of the seventh (n-type) FET 1524. The gate of the seventh FET 1524 is coupled to a third voltage source 1536 which sets the Potassium activation threshold. The drain of the seventh FET 1524 is coupled to the positive voltage source. The source of the eighth (p-type) FET 1526 is coupled to the positive voltage source. The gate of the eighth FET 1526 is coupled to the drain of the eighth FET 1526, the drain of the sixth FET 1522 and the gate of the ninth (p-type) FET 1530. The source of the ninth FET 1530 is coupled to the positive voltage source. The source of the tenth (n-type) FET 1528 is coupled to ground. The gate of the tenth FET 1528 is coupled to the drain of the drain tenth FET 1528, the drain of the ninth FET 1530 and the gate of eleventh (n-type) FET 1532. The source of the eleventh FET 1532 is coupled to ground. The drain of the eleventh FET 1532 is coupled to the rail 1501 and the sodium ion channel section 1504. The sodium ion channel section 1504 also comprises a plurality of field effect transistors (FETs) 1540-1568, a plurality of variable voltage sources and a timing capacitor 1571. The plurality of FETs comprises first, second, third, fourth, fifth, sixth, seventh, eighth, ninth, tenth, eleventh, twelfth, thirteenth, fourteenth and fifteenth FETs 1540-1568. Each FET 1540-1568 has a gate, a source and a drain. The first (n-type) FET 1540 has a source coupled to ground. The first FET 1540 has a gate coupled to the first variable voltage source 1538. The first FET 1540 has a drain coupled to the source of the second (n-type) FET 1542 and the source of the fifteenth (n-type) FET 1544. The gate of the second FET 1542 is coupled to the rail 1501 and the potassium ion channel section 1502. The gate of the second FET 1542 is also coupled to the drain of the fourth (p-type) FET 1548. The source of the fourth FET 1548 is coupled to the positive voltage source. The gate of the fourth FET 1548 is coupled to the gate of the third (p-type) FET 1546, the drain of the third FET 1546 and the drain of the fifth (p-type) FET 1550. The source of the third FET 1546 is coupled to the positive voltage source. The source of the fifth FET 1550 is coupled to the positive voltage source. The drain of the fifteenth (n-type) FET 1544 is coupled to the rail 1501. The gate of the fifteenth FET 1544 is coupled to the second variable voltage source 1570 that sets the activation threshold of the sodium channel The gate of the sixth (p-type) FET 1552 is coupled to the drain of the sixth FET 1552 and the gate of the fifth FET 1550. The source of the sixth FET 1552 is coupled to the positive voltage source. The source of the seventh FET is coupled to ground. The gate of the seventh (n-type) FET 1554 is coupled to the third variable voltage source 1576. The drain of the seventh FET 1554 is coupled to the source of the eighth (n-type) FET 1556 and the source of the ninth (n-type) FET 1558. The drain of the ninth FET 1558 is coupled to the drain and gate of the sixth FET 1552. The gate of the eighth FET 1556 is coupled to the fourth variable voltage source 1574, which sets the sodium inactivation threshold. The source of the eighth FET 1556 is coupled to the positive voltage source. The timing capacitor 1571 is coupled between the gate of the ninth FET 1558 and ground. The gate of the ninth (n-type) FET 1558 is also connected to the gate of the tenth (n-type) FET 1560, the drain of the tenth FET 1560 and the drain of the thirteenth (p-type) FET 1566. The source of the thirteenth FET 1566 is coupled to the positive voltage source. The source of the eleventh (n-type) FET 1562 is coupled to ground. The gate of the eleventh FET 1562 is coupled to the fifth variable voltage source 1578. The drain of the eleventh FET 1562 is coupled to the source of the tenth (n-type) FET 1562 and the source of the twelfth (n-type) FET 1564. The drain of the twelfth FET 1564 is coupled to the drain of the fourteenth (p-type) FET 1568, the gate of the fourteen FET 1568 and the gate of the thirteenth FET 1566. The gate of the twelfth FET 1564 is connected to the membrane section 1506. The membrane section 1506 comprises a membrane capacitor 1580 and a leakage resistance 1582 of the neuron membrane. The membrane capacitor 1580 is coupled between the gate of the twelfth FET 1564 of the sodium ion channel section 1504 and ground. The gate of the twelfth FET 1564 is also coupled to the rail 1501 and the leakage resistance of the neuron membrane 1582 within the membrane section 1506. In general, there may be provided a computer program, which when run on a computer, causes the computer to configure any apparatus, including a circuit, unit, controller, device or system disclosed herein to perform any method disclosed herein. The computer program may be a software implementation. The computer may comprise appropriate hardware, including one or more processors and memory that are configured to perform the method defined by the computer program. In non-pulsatile pumping systems, the first mode may implement RSA-like flow by increasing the pump rotor speed when the respiratory cycle is in an inspiration phase, thus increasing the rate of blood flow through the pump, and the second mode may implement RSA-like flow by decreasing the pump rotor speed when the respiratory cycle is in an expiration phase, thus decreasing the rate of blood flow through the pump. In pulsatile pumping systems, the first mode may implement RSA-like flow by increasing the speed of the pump when the respiratory cycle is in an inspiration phase thus providing a greater number of pulses in a given time frame, which may otherwise be described as decreasing the R-R interval, and the second mode may implement RSA-like flow by decreasing the speed of the pump when the respiratory cycle is in an expiration phase thus providing a smaller number of pulses in a given time frame, which may otherwise be described as increasing the R-R interval. Returning to Figure 2, in both non-pulsatile and pulsatile pumping systems, the change in pump speed may be achieved by modulating an instantaneous power supplied to the pump 206 by the control system 202. That is, the instantaneous power may be greater during the first mode of operation than during the second mode of operation. The control system 202 may be further configured to provide counterpulsation. The control system 202 may be configured to provide the counterpulsation by timing the pumping pulses such that the pumping pulses are applied in synchrony and out of phase with a heartbeat cycle. The control system 202 is configured to pulse the pumping operation such that pulses are applied during a diastole phase of the heartbeat cycle. Advantages associated with application of counterpulsation include increasing ventricular unloading and increasing the myocardial supply / demand ratio. The control system 202 according to the present example may provide the additional advantages of opening the aortic valve while at the same time providing increased pulse pressure. Figure 3a is a profile of calcium ion (Ca2+) concentration observed when monotonic pacing is applied to a heart. The profile shows the stabilisation of the calcium ion concentration over time. Cardiomyocytes rely on Ca2+ signaling for their core task as contractile units in the heart, and Ca2+ dysregulation is one of the main contributors to failure of the heart as pump. Yet, Ca2+ signaling in cardiomyocytes is equally pivotal as a mechanism for regulating cardiomyocyte growth and physiological remodeling. Further, Ca2+-handling mechanisms are remodeled in disease, generating a unique link with contraction and arrhythmias in heart disease. Calcium transient amplitude is the difference between the peak value and the diastolic value of the cytosolic Ca2+ concentration. It is a major factor in determining the contractile force. In heart failure, the amplitude of the calcium transient decreases, and the calcium transient duration increases. Figure 3b is a profile of calcium ion (Ca2+) concentration observed when RSA pacing is applied to a heart. The profile shows the variation of the calcium ion concentration over time. Rhythmic increases in intracellular calcium ion (Ca2+) concentration underlie the contractile function of the heart. The present inventors, using a computational model of the heart, have demonstrated that an RSA like contraction of the heart is associated with changes in calcium ion (Ca2+) concentrations in the myocytes vs. a monotonic rhythm of contraction. These likely influence cardiac contractility and morphology - influencing the contractile proteins in particular. In general, corresponding reference numerals are used between the different figures to refer to corresponding components of the system. Figure 4 illustrates a schematic block diagram of an example blood flow assistance device 400 generally corresponding to the embodiment described in Figure 2. In this example, the controller 402 comprising one or more processors 410 and optionally memory (not shown). The controller 402 receives a signal indicative of a respiration cycle as an input and determines whether the respiration cycle is in an inspiration phase or an expiration phase. The signal may be indicative of an instantaneous respiration cycle. The signal may be indicative of a feature of the respiration cycle, such as the onset or end of the inspiration phase or of the expiration phase, and the control system may control the operation of the pump in accordance with that feature. In a preferred example, a respiratory cycle of a patient is measured using one or more sensors 408 connected to or implanted into a patient. The sensors 408 may be dEMG (diaphragmatic electromyography) sensors for the detection of the contraction of muscles. The sensors 408 may convert the measured respiratory cycle into a signal indicative of the instantaneous respiratory cycle, which is received by the control system. One or more intermediary steps may convert the signal indicative of the respiration cycle into any one of: a signal indicative of a feature of the respiration cycle, or a signal indicative of a respiration duty cycle. In some embodiments, the controller may further comprise a respiration analysis module configured to determine a signal indicative of a feature of the respiration cycle or a signal indicative of the respiration duty cycle. The respiration analysis module may comprise a threshold comparator configured to determine whether the respiration signal is associated with an inspiration phase or an expiration phase. The embodiment of the control system described above in relation to Figure 4 may be used with an in vivo blood flow assistance device comprising a pump 406, such as an LVAD. In such an example, the LVAD is implanted into a patient such that pumping function of a heart may be restored, and dEMG sensors may be placed on or implanted into the patient’s muscles involved in diaphragm contraction in order to generate the respiration cycle. Alternatively to the in vivo example described above, the control system may be used with an ex vivo blood flow assistance device. Various pumps that are used with the cardiovascular system such as perfusion pumps for transport of organs during transplant could be improved by an RSA like rhythm in the supplied blood flow. The benefits may include improved cardiomyocyte structure and function resulting in improved organ function post-transplant and therefore better outcomes for patients and reduced rejection rates, thus reducing rejection medication requirements. The technology is applicable to cardiac transplants. However, the vasculature of other organs also would likely benefit from a more physiological-like perfusion. Figure 5 illustrates a schematic block diagram of an example device 500 generally corresponding to the embodiment described in Figure 2. The device 500 is suitable for use with an ex vivo blood flow assistance pump 506. Such uses may arise when, for example, blood flow needs to be maintained through an organ being prepared, held, or transported for transplantation into a patient. In those situations, it is not possible to receive, sense, or detect a respiration cycle from a patient because a patient is not present. A controller 502 of the device 500 comprises one or more processors 510 and memory 512. The controller502 is configured to retrieve data describing the respiration cycle from the memory 512. In some embodiments, the data is saved subject data for a subject having a respiratory sinus arrhythmia. In other embodiments, the respiration cycle is a synthetic respiration cycle having respiratory sinus arrythmia which may be generated by the controller 502. Figure 6 illustrates a schematic block diagram of an example of a blood flow assistance device 600 generally corresponding to any one of the embodiments described above, and further comprising a power source 614 electrically connected to a controller 602 of the device. The controller 602 also comprises a housing 620 in which the controller 602 and power source 614 are disposed. The controller 602 or power source 614 may comprise driver circuitry configured to provide a signal to power a pump (not shown) in accordance with a pumping signal determined by the controller 602. The housing may comprise a driveline connector or an opening for a driveline to enter the housing. A driveline may be electrically connected to one or more of the controller 602 and power source 614. The driveline may be configured to connect the controller 602 to a pump (not shown) of a blood flow assistance device. More generally, the driveline may be defined as a wired electrical link, or cable. The wired electrical link may be configured to supply power from the control system 500 to the pump of a blood flow assistance device. Figure 7 illustrates a schematic block diagram of an example of a blood flow assistance device 700 comprising a pump 706. The pump 706 comprises an inlet and an outlet. The blood flow assistance device 700 also comprises a control system 702 in accordance with any of the previously disclosed embodiments electrically connected to the pump 706. The inlet and the outlet may be in fluid communication with an organ 716 which is subject to blood flow assistance. Figure 8 illustrates a schematic block diagram of an example of a blood flow assistance device 800 comprising the device discussed above in relation to Figure 6 and a cardiac pump 806. The inlet and the outlet of the pump 806 may be in fluid communication with an organ 816 which is subject to blood flow assistance. The housing 820 containing the controller 802 and power source may be electrically connected to the pump 806 via a wired electrical link 818, such as a driveline. In relation to the embodiments described above, the blood flow assistance device may comprise any one of the following: an LVAD device; an RVAD device; an artificial heart; a perfusion pump; or a cardiopulmonary bypass machine. Figure 9 illustrates a schematic block diagram of a system 901 comprising a blood flow assistance device 900 and a cardiac pacemaker 922. The blood flow assistance device 900 may be in accordance with any of the embodiments discussed above. The device is in fluid communication with a heart 916 of a patient, and the cardiac pacemaker 922 is in electrical communication with the heart 916 of the patient. In some cases of heart failure, the natural biological systems governing the electrical signals which cause RSA in a native heart may be weakened. Lack of or weakened RSA may also be a precursor of heart failure. A cardiac pacemaker according to the disclosure of UK Patent GB2586987 may be fitted alongside a blood flow assistance device according to the present disclosure, such that the native heart is paced by the cardiac pacemaker according to electrical stimulus signals exhibiting RSA, and the cardiac pump supporting the weakened heart muscles is operated by the control system disclosed in the present application such that the supportive pumping operation also exhibits RSA or RSA-like blood flow. This combination may ensure synchrony between the native heart RSA and the cardiac pump RSA. A single controller may be provided for the device 900 and pacemaker 922. Alternative, the device 900 and pacemaker 922 may have respective controllers that are configured to communicate a synchronization signal.

Claims

1. A control system for a blood flow assistance device comprising a pump, the control system configured to control operation of the pump based on a respiration cycle having an inspiration phase and an expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein the first mode of operation is different to the second mode of operation.

2. The control system of claim 1, wherein the control system is configured to pulse the pumping operation.

3. The control system of claim 1 or 2, wherein the control system is configured to pulse the pumping operation having a respiratory sinus arrhythmia.

4. The control system of any preceding claim, wherein a frequency of the pulses is greater during the first mode of operation than during the second mode of operation.

5. The control system of any preceding claim, wherein the control system is configured to synchronize the pulsing of the pumping operation with the onset of the inspiration phase or the expiration phase.

6. The control system of any preceding claim, wherein the control system is configured to generate the timing of the pumping pulses as a function of a respiration duty cycle in order to maintain a bias towards synchronization between the respiration period and an integer ratio of the periods between pumping pulses.

7. The control system of any preceding claim, wherein the control system is further configured to provide counterpulsation.

8. The control system of any preceding claim wherein the control system is configured to modulate the pumping operation to provide opening the aortic valve while at the same time providing increased pulse pressure.

9. The control system of any preceding claim, wherein the pumping operation is continuous.

10. The control system of any preceding claim, wherein the pump has a higher flow rate during the first mode of operation than during the second mode of operation.

11. The control system of any preceding claim, wherein controlling operation of the pump comprises modulating an instantaneous power supplied to the pump by the control system.

12. The control system of any preceding claim, wherein the instantaneous power is greater during the first mode of operation than during the second mode of operation.

13. The control system of any preceding claim, where the control system is configured to receive a signal indicative of a feature of the respiration cycle and to control operation of the pump in accordance with the feature.

14. The control system of claim 13, wherein the feature may be the onset or end of the inspiration phase or of the expiration phase.

15. The control system of claim 13 or 14 wherein the signal is indicative of an instantaneous respiration cycle of a subject.

16. The control system of claim 15 wherein the control system is further configured to determine, from the instantaneous respiration cycle, an instantaneous respiration duty cycle.

17. The control system of any preceding claim, wherein the control system has a memory and is configured to retrieve data describing the respiration cycle from the memory.

18. The control system of claim 17 wherein the data is saved subject data for a subject having a respiratory sinus arrhythmia.

19. The control system of claim 17 or 18 wherein the respiration cycle is a synthetic respiration cycle.

20. The control system of claim 19 wherein the control system is configured to generate the synthetic respiration cycle having a respiratory sinus arrhythmia.

21. The control system or any preceding claim further comprising driver circuitry for operating the pump.

22. The control system of any preceding claim, wherein the system is further configured to provide signals for pacing the heart of a patient, wherein the operation of the pump and the pacing is synchronized by the control system.

23. The control system of claim 22, comprising a synchronization module configured to generate the timing of the signals for pacing the heart of the patient as a function of a signal indicative of respiration duty cycle in order to maintain a bias towards synchronization between the respiration period and an integer ratio of the periods between stimulus signals.

24. A blood flow assistance device controller comprising:the control system of any preceding claim;a power source electrically connected to the control system; anda housing, wherein the power source and the control system are provided within the housing.

25. A blood flow assistance device comprising:a pump comprising an inlet and an outlet;the control system of any preceding claim, or the blood flow assistance device controller of claim 24, electrically connected to the pump.

26. The blood flow assistance device of claim 25, further comprising a sensor for determining a feature of a respiratory cycle of a patient.

27. The blood flow assistance device of claim 25 or 26, wherein the blood flow assistance device comprises:an LVAD device;an RVAD device;an artificial heart;a perfusion pump; ora cardiopulmonary bypass machine28. A control system for a blood flow assistance device comprising a memory, the control system configured to:retrieve data describing the respiration cycle from the memory; andcontrol operation of a pump based on the respiration cycle having an inspiration phase and an expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein 5 the first mode of operation is different to the second mode of operation.

29. A non-transitory computer-readable storage medium comprising computer program code configured to cause a processor to execute a method for controlling operation of a pump based on a respiration cycle having an inspiration phase and an 10 expiration phase such that the pump is operated in a first mode of operation in the inspiration phase, and a second mode of operation in the expiration phase, wherein the first mode of operation is different to the second mode of operation.

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