CONTROL FOR NON-OCCLUSIVE BLOOD PUMPS

Description

Technical area

[0001] The present invention relates to control and regulation units for non-occlusive blood pumps of extracorporeal circulatory support as well as systems comprising such a control and regulation unit and corresponding methods. State of the art

[0002] If the heart's pumping capacity or pumping function fails, cardiogenic shock can occur. This can lead to reduced perfusion or blood flow to end organs such as the brain, kidneys, and the vascular system in general due to a reduction in cardiac output or cardiac output. This acute heart failure causes an acute lack of blood supply to the tissues and organs and thus a lack of oxygen, also known as hypoxia, which can lead to end organ damage. In most cases, this type of cardiogenic shock occurs as a complication of an acute myocardial infarction (AMI) or heart attack, making it one of the most common causes of death from myocardial infarctions. In Germany alone, around 350,000 people are struck by a myocardial infarction each year, of which around ten percent develop cardiogenic shock.

[0003] Furthermore, cardiogenic shock can occur as a complication of surgical treatment such as bypass grafting, inadequate or impaired lung function, or as a result of conduction disturbances, structural heart disease, or myocardial inflammatory processes. Although factors such as early revascularization, the administration of inotropic drugs, and mechanical support can improve the patient's physiological status, the mortality rate from cardiogenic shock remains above 50 percent.

[0004] To improve the patient's condition, circulatory support systems have been developed that provide mechanical support and can be rapidly connected to the circulatory system to improve blood flow and perfusion to the organs, including the coronary arteries, and prevent a hypoxic state. For example, systems have been developed in which a balloon is invasively inserted into the aortic region and controlled to rapidly inflate during the diastolic phase of the cardiac cycle, for example, with helium, to improve blood flow to the myocardial tissue during this phase. Such systems are known as "intra-aortic balloon counterpulsation" (IABP). However, the effectiveness of such systems is very limited, especially since they only deliver blood to the coronary arteries that was not sufficiently oxygenated during a cardiogenic shock.Therefore, this procedure does not offer any significant improvement compared to drug treatment with, for example, inotropic drugs or catecholamines.

[0005] Furthermore, extracorporeal circulatory support systems are known. These are based on withdrawing blood from, for example, the right atrium or vena cava through a cannula inserted into the femoral vein and returning it to the patient via a membrane ventilator through a perfusion cannula inserted into the femoral artery. The membrane ventilator decarburizes, i.e., removes carbon dioxide, and oxygenates the blood. Such systems are also known as "extracorporeal membrane oxygenation" (ECMO).

[0006] Although this improves organ perfusion, perfusion occurs continuously during both the systolic and diastolic phases, resulting in retrograde blood flow that counteracts the heart's normal ejection direction and thus increases afterload. As a result, the heart's ejection is hindered and empties to a lesser extent, while at the same time the heart must perform more muscular work and therefore requires an increased oxygen supply. This is particularly disadvantageous for a heart already damaged by a myocardial infarction. Therefore, such a system is primarily used in cases of cardiac arrest or when the patient's own cardiac output is intentionally bypassed, for example, during a surgical procedure or intervention.To reduce afterload in a beating heart, the patient may also be given medication such as vasopressors, which dilate the blood vessels and thus reduce resistance and afterload. In these cases, however, the patient's existing cardiac output is less therapeutically supported, but rather replaced by the system.

[0007] Furthermore, blood flow to the heart's own coronary arteries, which normally supply the heart muscle with sufficient oxygen, generally occurs during the diastole of the cardiac cycle and is therefore dependent on a corresponding emptying of the left ventricle. When the filling pressure in the left ventricle is as low as possible at the end of systole or at the beginning of diastole, the coronary arteries can expand their lumens as much as possible, thus increasing the blood flow rate and oxygen supply. The percutaneous femoro-femoral extracorporeal circulatory system and the resulting retrograde flow reduce the emptying of the left ventricle, making the system also disadvantageous for the perfusion of the coronary arteries. Furthermore, the flow of the blood pump can be set in such a way that the blood flow closes the aortic valves.Because blood from the pulmonary vein continues to be pumped through the right atrium into the left ventricle and cannot escape, the pressure in the left ventricle increases further, which can result in a further reduction in coronary artery perfusion. In other words, the assisted perfusion that is beneficial to the end organs has a detrimental effect on coronary perfusion.

[0008] Another problem with many mechanical assist systems is that they do not provide blood flow that corresponds to normal physiological blood flow, i.e., pulsatile blood flow. Although systems with pulsatile or non-occlusive blood pumps are known and can improve end-organ perfusion, the negative impact on afterload and the corresponding reduction in coronary artery perfusion has not been sufficiently considered. Due to the heart rate, which can also be variable, and the correspondingly short duration of the respective phases, the options for providing adequate end-organ perfusion while simultaneously improving coronary artery perfusion are severely limited.In order to take possible fluctuations into account, a generally lower blood flow or a shorter pulsatile perfusion may be set, but this is fixed and does not allow for perfusion optimized for the physiological state.

[0009] Accordingly, there is a need to optimize the blood flow characteristics provided by a blood pump in such a way that they are adaptable to the patient's physiological condition and provide optimization of both end-organ and coronary artery perfusion. US 2015 / 0141911 A1 discloses a blood pump that provides pulsatile blood flow in an extracorporeal blood circulation system. Description of the invention

[0010] Based on the known prior art, it is an object of the present invention to enable improved perfusion of the coronary arteries during extracorporeal circulatory support.

[0011] The problem is solved by the independent claims. Advantageous further developments emerge from the subclaims, the description, and the figures.

[0012] Accordingly, a control and regulation unit for a non-occlusive blood pump of extracorporeal circulatory support is proposed, which is configured to receive a flow value of the extracorporeal circulatory support, to receive a measurement of an arterial pressure and an ECG signal of a supported patient over a predetermined period of time, to determine a mean arterial pressure of the extracorporeal circulatory support or of the supported patient based on the measurement of the arterial pressure and an energy equivalent pressure based on the flow value and the arterial pressure, and to determine a required value of at least one pulse parameter of the blood pump based on the mean arterial pressure, the energy equivalent pressure and the ECG signal in order to effect a ratio of the energy equivalent pressure to the mean arterial pressure which is greater than 1.0.The control and regulation unit is also designed to adjust the pulse parameter depending on the ECG signal.

[0013] The control and regulation unit can, for example, control or regulate one or more pump drives or pump heads for non-occlusive blood pumps present in a system for extracorporeal circulatory support. The blood pump can be arranged for oxygen enrichment and decarburization, i.e., carbon dioxide (CO2) removal or depletion, of the blood. For example, the blood pump can be fluidically connected to an oxygenator or membrane ventilator provided in the extracorporeal circulatory support system, which is connected, for example, to a venous access via a venous cannula and to an arterial access via an arterial cannula for sucking in or pumping the blood in order to provide blood flow from a side with a low pressure to a side with a higher pressure.In this way, the blood pump can provide extracorporeal oxygenation of the blood and improved oxygen supply to the patient. These features can be further enhanced due to the pulsatile nature of the non-occlusive blood pump and the ratio between the energy-equivalent pressure and the mean arterial pressure, particularly with regard to the perfusion of the coronary arteries and end organs. By specifying the aforementioned ratio, the control and regulation unit is not dependent on fixed or predetermined values ​​of the pulse parameter. In particular, the pulse parameter can be directly adapted and optimized to the physiological requirements of the respective patient.

[0014] The non-occlusive blood pump can, for example, be magnetically coupled to the pump drive of a motor to enable torque transmission. A centrifugal pump such as a rotor pump or impeller pump, for example a centrifugal pump or a semi-axial pump, also known as a diagonal pump, can be provided as the non-occlusive blood pump. While a centrifugal pump can generate high pressure at a low flow rate, diagonal pumps can generate a high flow rate at a low pressure, which meets the general requirements for pulsatile blood pump systems and is particularly advantageous for extracorporeal circulatory support and respiratory support in the event of cardiogenic shock. For example, the rapid provision of a high flow rate with a low pressure level can enable improved and more accurate synchronization based on an ECG signal.

[0015] Furthermore, unlike a roller pump or occlusive blood pump, a rotor pump cannot build up pressure and release it at a predetermined time. However, by changing the speed of the rotor blade or impeller, a small amount of pressure can be built up, which is then transmitted as a flow. This has the advantage of enabling precise pulsation without the need for additional components such as valves. Furthermore, the small dimensions of the rotor blade / impeller also allow for low inertia, thus enabling very precise control and precise delivery of a pulse at any speed.

[0016] Because human organs and cells inherently detect and process dynamic flow and pressure changes, they also respond differently to different flow profiles. In particular, in the vascular endothelium, which forms the innermost layer of the vessels, mechanotransduction and signaling occur in response to shear stress and pulse pressure, regulating functions such as apoptosis, angiogenesis, atherosclerosis, and systemic blood pressure. In this respect, the non-occlusive blood pump is further advantageous, especially since it provides pulsatile blood flow, thereby reducing the mechanical and / or thermal stress on blood cells and, in particular, preventing the destruction of red blood cells and the activation of platelets.

[0017] In general, pulsatile blood flow also results in improved microcirculation of vital organs, improved vascular compliance, reduced inotropic support, higher cerebral oxygen saturation, increased urine output, and reduced gastrointestinal bleeding rates, resulting in patients generally requiring a shorter hospital stay.

[0018] The blood pump can be connected to a venous access via a venous cannula and to an arterial access via an arterial cannula for aspirating or pumping the blood, respectively, in order to provide blood flow from a side with a low pressure to a side with a higher pressure. The blood pump is preferably designed as a disposable or single-use item and is fluidically separated from the respective pump drive and easily coupled, for example, via a magnetic coupling. The control and regulation unit actuates the motor of the pump drive and is configured to adjust the speed of the rotor blade of the blood pump such that the blood pump generates a wave-like flow corresponding to a pulsatile blood flow.

[0019] Furthermore, by changing the speed, the pulse amplitude can be specified, which can be entered into the control and regulation unit and optionally calculated to a base flow. For example, a base flow can be specified, with the respective pulses causing an increase in the total blood flow. This results in an average flow, which is then calculated accordingly by the control and regulation unit. For example, the blood pump can be set for a speed between approximately 50 RPM and approximately 10,000 RPM to provide a blood flow between approximately 0.1 l / min and approximately 8 l / min, with the pulsatile flow resulting from the change in speed.However, a blood flow between about 0.5 l / min and about 5.0 l / min (for example between 1 l / min and 2.5 l / min) is preferred in order to prevent coagulation on the one hand and an increase in afterload on the other hand, and to achieve a desired flow that provides adequate organ oxygenation, i.e. a sufficient number of oxygenated red blood cells.

[0020] Another advantage of pulsatile blood flow is that it generates pulse energy, which results from the pressure change with respect to the flow gradient over time. Multiplying the respective integrals yields the hemodynamic work, which, when divided by the flow integral, yields the energy equivalent pressure (EEP). This relationship is mathematically expressed in the following formula: EEP mm Hg = ∫ t 1 t 2 fpdt / ∫ t 1 t 2 fdt where the blood flow (f) and pressure (p) are taken into account over a time interval from t1 to t2.

[0021] The energy-equivalent pressure is comparable to the patient's blood pressure and is based on the idea that hemodynamic energy is transferred from the aortic root to the periphery through the axial movement of blood flow and the elastic movement of the aortic wall. Accordingly, the energy-equivalent pressure can be further converted into excess hemodynamic energy, also known as "surplus hemodynamic energy" (SHE), which preferentially reflects the energy for a single pulse and corresponds to the additional energy generated by pulsatile flow relative to non-pulsatile or baseline flow. Thus, both the energy-equivalent pressure and the excess hemodynamic energy provide quantifiable measures of circulatory and / or cardiac support.

[0022] To achieve an improvement over constant blood flow and an increase in the patient's blood pressure, the energy-equivalent pressure should be higher than the mean arterial pressure, also known as "mean arterial pressure" (MAP). In other words, a ratio of the energy-equivalent pressure to the mean arterial pressure that is at least 1.0 or greater than 1.0 (e.g., > 1.0 and < 1.25) ensures that a physiological blood flow is provided. A pulse contribution is calculated that causes an increase in the mean arterial pressure. For example, the pulse contribution can be a minimum of 5 mmHg and preferably between approximately 10 mmHg and 20 mmHg. The pulsing provided by the control and regulation unit is therefore to be understood as physiological or quasi-physiological pulsing.

[0023] For example, the energy-equivalent pressure due to pulsatile blood flow or the pulsation of the blood pump itself can be higher than with constant blood flow at the same mean arterial pressure, thus providing improved blood flow to the distal vessels. The difference in energy-equivalent pressure is an indicator of the higher mean pressure required with non-pulsatile flow to achieve the same blood flow.

[0024] Accordingly, the energy-equivalent pressure can be determined by the control and regulation unit, whereby the flow value provided by the blood pump can be entered manually, for example, via the rotor blade speed, and only a measurement of the arterial pressure is required. From this measurement, the mean arterial pressure can also be determined by averaging the values.

[0025] The arterial pressure is preferably fed into the control and regulation unit via an interface that is communicatively connected to at least one pressure sensor. The pressure measurement preferably takes place before the blood pump, after the blood pump and / or after an oxygenator or membrane ventilator provided in the extracorporeal circulation system and connected to the blood pump. The pressure measurement of the arterial pressure of the assisted patient is preferably carried out invasively in the patient, for example at the tip of an arterial cannula, for example in the patient's aorta, or by another pressure measuring probe that can be positioned at any convenient location in the aorta. Alternatively, the measurement can also be performed at the a. femoralis and / or a. iliaca or or a. radialisIn addition, there is the possibility of pressure measurement using non-invasive pressure sensors (without direct blood contact) or invasive pressure measurement with blood contact on the tube ends of the patient kit.

[0026] Both the flow rate and the arterial pressure can be recorded non-invasively and received by the control and regulation unit, for example, by recording the flow rate via the rotor blade speed and / or the arterial pressure via non-invasive tubing or in the circulatory system outside the patient. This allows for a compact design of the circulatory system and reduces its complexity. For the patient, this also means that inserted cannulas can be smaller and surgical procedures can be performed less invasively.

[0027] From the determined energy-equivalent pressure and the determined mean arterial pressure, the control unit can determine the corresponding ratio and adjust the blood pump's pulse parameter accordingly to achieve the desired ratio of more than 1.0. This can be done iteratively and preferably continuously, so that the effect of changing the blood pump's pulse parameter on the patient's physiological condition is taken into account and the relationship between the energy-equivalent pressure and the mean arterial pressure is not dependent on fixed values ​​of the pulse parameter and can be optimized.

[0028] Preferably, the ratio is greater than 1.1 to further improve therapeutic support for the patient. For example, a determined ratio of 1.0 may require an increase in the energy equivalent pressure and / or a decrease in the mean arterial pressure, and a first change in a pulse parameter may result in a ratio of 1.05 and a second change in the value of the pulse parameter may result in a ratio of 1.15, so that no further change in the pulse parameter is required. The ratio may further be 1.15 to 1.25, for example, approximately 1.2, for example in the event of cardiac arrest during patient resuscitation when the patient's mean arterial pressure is within a lower limit.

[0029] The value of the pulse parameter is further adjusted based on an ECG signal, preferably the heart rate. The ECG signal is preferably a characteristic ECG signal that enables synchronization of the control and regulation unit with the blood pump, so that a regular or periodic output of a signal corresponding to the value of the pulse parameter can occur from the control and regulation unit to the pump drive or to the blood pump. For example, the ECG signal or the respective region in the electrical excitation line can be characteristic of the systolic or diastolic phase of the heart, so that the adjustment of the pulse parameter and the actuation of the motor can occur at a predetermined time and in a predetermined phase, without causing any overlap with other phases.

[0030] By synchronizing the pulse with the ECG signal, a pulsatile blood flow is provided that corresponds to the patient's cardiac cycle. This allows the pulse rate of the blood pump to be adjusted to the patient's heart rate and, if necessary, other physiological parameters. In the case of tachycardia, for example, a pulse parameter can be changed so that the pulse does not overlap or interfere with a phase of the cardiac cycle in which the pulse is not intended. The required ratio between the energy-equivalent pressure and the mean arterial pressure nevertheless ensures sufficient perfusion of the coronary arteries and end organs, within the physiological limits and the pump's properties.

[0031] Compared to known ECMO systems with a continuous blood flow, the invention thus makes it possible to achieve both a physiological blood flow and an adaptation to the patient's cardiac cycle, while keeping the afterload, ie a remaining pressure in the aortic region before or during the systolic phase, as low as possible or not increasing it.

[0032] Furthermore, compared to constant or continuous flow, pulsatile flow also provides higher mean arterial pressure and higher energy-equivalent pressure, as well as improved excess hemodynamic energy. This results in higher systemic blood pressure with less inotropic support. This not only improves perfusion of the coronary arteries and end organs, but can also provide improved oxygenation, for example, if an oxygenator or membrane ventilator is included in the extracorporeal circulatory support system. Compared to existing IABP systems, which provide no or insufficient extracorporeal circulatory support, this results in improved unloading and recovery of the myocardial tissue.

[0033] The at least one pulse parameter preferably comprises a pulse amplitude, a pump speed, a pump speed change per unit time, a pulse duration, a systolic pump duration, a diastolic pump duration, a drive wheel deceleration, a drive wheel acceleration, an average pump flow, a base flow and / or a peak flow.

[0034] For example, the pulse amplitude and pump speed can be used to specify the maximum blood flow during the pulse and thus the mean flow. The pump speed is preferably optimized to a maximum value based on the determined ratio between the energy-equivalent pressure and the mean arterial pressure in order to obtain a required pulse amplitude. For example, a pulse amplitude can be provided by changing the pump speed, whereby the change in pump speed can have a value between 50 rpm and 10,000 rpm and is preferably between 1,000 rpm and 5,000 rpm, particularly preferably between 3,000 rpm and 4,000 rpm. In this way, a pulsatile blood flow with a sufficient mean flow can be provided, for example between approximately 2 and 3 l / min.A pump speed change can optionally be added to a base flow, whereby the maximum pump speed achieved is specified from a base speed, the pump speed change per unit of time and the pulse duration.

[0035] The pulse can also be determined by the pulse duration, with a pulse duration preferably being between 20 ms and 400 ms. For example, the pulse duration can be approximately 200 ms to provide a sufficient pulse at a normal heart rate. The pulse comprises a systolic pumping duration and a diastolic pumping duration, during which a drive wheel acceleration and deceleration, respectively, occur. These phases of the pulse duration can be symmetrical, so that the systolic phase comprises approximately 50% of the pulse duration and, for example, comprises 100 ms.

[0036] However, the systolic phase can also be adapted to the pulse duration, for example, if the pulse duration is shortened due to an increased heart rate and the systolic phase is prolonged in order to provide a sufficient pulse even with a shortened pulse duration. In this case, the pulse duration is asymmetrical, whereby the systolic phase can, for example, comprise approximately 60% or approximately 70% of the pulse duration. To prevent increased blood flow from being provided even after the pulse duration, the drive wheel deceleration can also be increased during the diastolic phase. Conversely, the drive wheel acceleration can also be increased during the systolic phase in order to provide a shorter systolic phase with a sufficient change in speed. In this case, the systolic phase can, for example, comprise approximately 30% or approximately 40% of the pulse duration.

[0037] Accordingly, even with a shorter time interval, the blood pump can be regulated to provide sufficient perfusion without causing an undesirable pulse overlap with a specific phase of the cardiac cycle. This can reduce or even prevent unwanted afterload.

[0038] To further improve the pulse duration and pulse energy, the speed reached during the pulse can be passively decelerated (without active braking) or actively braked. In the case of active braking, for example, a four-quadrant control system can be provided, with four transistors arranged in mirror symmetry, forming an H-bridge with one p-channel and one n-channel. Appropriate activation of the transistors enables a reversal of the current direction through a brushless DC motor, so that the pump drive motor can be switched in two opposite directions, and a speed reduction can be actively achieved by reversing the current direction.

[0039] Furthermore, the blood pump can also be controlled to cause negative blood flow at the end of the pulse. This causes aspiration, which can, for example, be synchronized so that it occurs during the systolic phase to further reduce afterload.

[0040] Cardiac output can also be used as a further reference variable for the pulse, which can be determined using an impedance measurement. For example, the impedance can be determined using pressure sensors provided in the extracorporeal circulatory support system, such as integrated pressure sensors, or via an impedance measurement in the patient's thoracic region, either with transcutaneous sensors or an impedance measurement probe placed in the esophagus.

[0041] The required value of the pulse parameter can further be determined based on a given mean arterial pressure and / or a given energy-equivalent pressure. Alternatively, or additionally, this can be determined based on a measured cardiac output volume, stroke volume, and / or ejection fraction.

[0042] For example, predefined threshold values ​​can be stored in the control and regulation unit, which correspond to key values ​​of vital parameters or the patient's physiological state. In other words, physiological limits for the energy-equivalent pressure and the mean arterial pressure can be predefined, which can be absolute and / or patient-specific values. For example, the mean arterial pressure can be predetermined by an administered catecholamine dose and / or the body volume or weight of the supported patient. Furthermore, a high heart rate of the supported patient, for example, may require a reduction in the mean arterial pressure, so that the pulse parameter is limited to a corresponding maximum value of the energy-equivalent pressure.These values ​​can still be actively specified through current measurements of vital parameters, for example via an interface with a communicatively connected ECG device.

[0043] In order to receive the flow value required for the energy-equivalent pressure, the control and regulation unit can be further configured to determine the flow value based on an input speed of the blood pump or automatically based on a received flow measurement.

[0044] The pump flow value can, for example, correspond to a specific speed, so that entering the speed via a user interface of the control and regulation unit specifies the pump flow value. However, the pump flow value is preferably measured by a flow sensor, so that a currently measured flow velocity is received by the control and regulation unit to determine the energy-equivalent pressure. Since the flow value can also vary for each patient at a given speed, the flow sensor can be arranged, for example, at a cannula tip of the arterial cannula, the venous cannula, and / or in the line between the blood pump and the respective tip, thus enabling an accurate measurement of the flow value.

[0045] Preferably, the flow value comprises a patient-specific mean flow, wherein the control and regulation unit can be further configured to adjust the mean flow based on a received flow measurement and the mean arterial pressure by adjusting the pulse parameter.

[0046] As described above, the median flow can be provided, for example, by increasing the speed and the corresponding pulse amplitude. However, since the physiological and pathophysiological condition of each patient varies, a patient-specific median flow may be required for each patient. In addition to flow measurement, which can be provided by a flow sensor arranged in the extracorporeal circulatory support system as described above, the optimal median flow for the patient can be determined based on the mean arterial pressure and thus adjusted by adjusting the pulse parameter, for example, via the speed of the blood pump.

[0047] This enables automated adjustment of the blood pump and extracorporeal circulatory support. In addition to mean arterial pressure and flow measurement, other vital parameters can be taken into account when determining the optimal flow, such as heart rate, age, clinical picture, and other physiological parameters of the patient.

[0048] The flow, pressure, and pulse characteristic parameters can also be controlled by specifying an EEP value. The current EEP value is shown on the console display, which can then be adjusted to a target EEP. The display can also be shown as an EEP percentage, which refers to the maximum possible EEP value for the individual patient. This maximum EEP value can be determined for each patient using a short empirical algorithm.

[0049] To enable optimal synchronization of the pulse with the respective phases of the cardiac cycle, the control and regulation unit is preferably configured to determine a heart rate from the ECG signal and adjust the pulse parameter based on the heart rate. For example, the pulse duration can be adjusted to the heart rate, allowing the motor or rotor blade to be actuated at a time when it does not collide with a phase of the cardiac cycle.

[0050] Alternatively, the heart rate can also be determined from a pressure measurement, a pressure curve, a pressure change, cardiac ejection volume, an ejection fraction, and / or an impedance measurement. The pulse parameter or the setting of the pulse parameter can be adjusted accordingly to the patient's heart rate, for example, by changing the pulse duration and adjusting the speed change.

[0051] The pulse parameter can optionally also be set based on a corresponding pulse rate. For example, the pulse parameter can alternatively be set for every second or third heartbeat, so that the pulse can be set at a ratio of 1:2 or 1:3 to the heart rate instead of a 1:1 ratio if the heart rate exceeds a threshold frequency. The heart rate can be monitored or evaluated for a predefined period of time. For example, a short-term increase in heart rate does not result in any change in the pulse rate and / or an initially reduced pulse rate can be reset or increased again after the heart rate has normalized.

[0052] In particular, the control and regulation unit can be configured to determine an amplitude change from the ECG signal and to adjust the pulse parameter at a predetermined time after the amplitude change.

[0053] For example, a P-wave, one or more characteristic features of the P-wave, a point or segment, or characteristic features of the QRS complex can be determined as the trigger signal, and the pulse parameter can be set at a predetermined time after the P-wave, the QRS complex, or the R-wave. However, other amplitude changes can also be determined, for example over a predetermined segment of the ECG signal or from a prominent point in the ECG signal. An R-wave is preferably determined from the ECG signal. The pulse parameter can be set at a predetermined time after the detection of the R-wave, for example the detection of the maximum amplitude, typically with a delay. Furthermore, the amplitude change can be averaged over a predetermined period of time so that any deviations or variations can be taken into account when setting the pulse parameter.

[0054] The R-wave can be determined, for example, from a surface ECG of the patient, but can alternatively also be determined invasively, for example via a derivation at the cannula tip of a venous drainage cannula in the right atrium or the superior vena cava, at the cannula tip of the arterial cannula in the aorta or the a. femoralis or the a. iliaca. Furthermore, the ECG signal can be derived via an invasive esophageal probe or via non-invasive pressure sensors on tube legs, for example pressure sensors integrated into the tube system (IPS).

[0055] The R-wave of the QRS complex read from the ECG signal can then be used as a trigger signal, with the pump signal or the setting of the pulse parameter occurring after a predefined delay. For example, the delay can be selected so that the pump signal occurs at the time when a brief increase in pressure occurs in the aorta, which marks the moment of aortic valve closure and thus the beginning of diastole (Dicrot's point). By activating the motor or pump drive and thus the rotor blade, and triggering the pulse at this time, a pulsatile blood flow can be delivered via an arterial cannula into the area (directly) downstream of the left ventricle and into the aorta.The positioning and countercurrent orientation of the cannula tip thus results in increased pulsatile blood flow in the coronary arteries and provides a correspondingly improved perfusion of the myocardial tissue, without impairing the systemic circulation.

[0056] However, the pressure increase caused by the pulse must be completed by the beginning of systole to prevent overlap. Accordingly, the pulse duration is limited, although this is not a problem with a normal heart rhythm. For example, there may be approximately 300 ms between the detected R-wave and the closure of the aortic valves, and a waiting time between the detection of the R-wave and the delivery of the pump signal may be approximately 200 ms, for example, between 150 ms and 250 ms, so that actuation by outputting the pump signal or pulse parameter may take 100 ms. Although the pump signal is generated before the aortic valves close, the pump signal is preferably output such that the pulse or pressure level is at the cannula tip when the aortic valves are closed. This allows a latency for the fluid transport time to be taken into account with regard to the pump and lines used.

[0057] The pulse duration can be approximately 200 ms, for example, 100 ms to 300 ms, with the systolic phase of the pulse being approximately 100 ms if the pulse is delivered symmetrically. To ensure sufficient perfusion, a pulse amplitude can be selected by adjusting the speed between 2500 rpm and 4500 rpm, preferably approximately 3500 rpm.

[0058] However, at an elevated heart rate, for example, between 90 and 130 beats per minute, flow collision can occur between the pump runout and the onset of systole or cardiac output. To prevent this and optimize afterload, a pulse amplitude can be selected by adjusting the speed between 1500 rpm and 5000 rpm, and the pulse duration can range, for example, between 120 ms and 300 ms. The pulse can be asymmetrical, with the systolic phase comprising approximately 70% to approximately 90%. Thus, even at a higher heart rate, afterload reduction and sufficient perfusion of the coronary arteries can be provided.

[0059] Accordingly, pulsatile blood flow and the adjustment of the pulse parameter based on the R wave of the QRS signal can provide improved blood flow to both the heart muscle and the end organs. The blood pump and pulse parameters can be fully adapted to the patient's intrinsic heart rhythm through the appropriate configuration of the control unit. In particular, pulse application can achieve diastolic augmentation without excessively increasing afterload, keeping it as low as possible.

[0060] A diagonal pump is preferably used as the blood pump. This has the advantage that the flow rate or mean flow can be low at a higher pressure, at least in diastole, and both the pressure and the flow rate can be increased in the diastolic phase. Thus, the remaining pressure after the pulse and before the systolic phase is as low as possible, preferably even negative with appropriate control. Furthermore, a diagonal pump has the advantage over a centrifugal pump that the impeller has a smaller circumference and a smaller mass to accelerate, allowing faster acceleration and deceleration, and the pulse can be adjusted more precisely because of the lower mass inertia.

[0061] In addition to the received ECG signal, the control and regulation unit can be further configured to receive a measurement of an aortic pressure of the assisted patient over a predetermined period of time and to adjust the pulse parameter at a predetermined aortic pressure and change in the aortic pressure.

[0062] Thus, the aortic pressure can serve as verification of the specified pulse delay and / or pulse duration, taking into account the latency for the fluid transport time. This allows a user of extracorporeal circulatory support to manually set the exact timing of the trigger signal and adjust it to the measured aortic pressure or ejection rate. However, the measured aortic pressure or ejection rate can optionally also provide feedback for the control and regulation unit, with the measured aortic pressure or ejection rate providing a feedback signal that is automatically considered when setting the pulse parameter. Thus, the setting of the pulse parameter can be further improved and performed iteratively.

[0063] Alternatively, the measured aortic pressure can also replace the ECG signal, with the aortic pressure being evaluated by the control unit, and a change in aortic pressure in a pressure curve being characteristic of a specific cardiac phase of the cardiac cycle. For example, a brief increase in aortic pressure following a drop in aortic pressure (Dicrot's point) can indicate the closure of the aortic valves and thus the onset of diastole, or the filling phase of the heart. Such a pressure change can be detected, for example, in a pressure curve as a Dicrot's point.

[0064] Preferably, the control and regulation unit is configured to determine a diastolic phase and systolic phase of the heart of the assisted patient from the ECG signal and / or the aortic pressure and to adjust at least the pulse duration such that it ends before the systolic phase.

[0065] While the R-wave of the ECG signal indicates the onset of systole or cardiac ejection, a brief rise in aortic pressure following a drop in aortic pressure indicates the closure of the aortic valves, the onset of diastole or heart filling, as described above. Accordingly, the accuracy of the signal output for regulating the blood pump can be further increased by combining the ECG signal and the measured aortic pressure, so that any ECG disturbances do not compromise patient safety or the functionality of the control and regulation unit.

[0066] Terminating the pulse before the systolic phase is, as described above, beneficial for afterload reduction and still allows for diastolic augmentation with sufficient end-organ perfusion. Such diastolic augmentation can be achieved, for example, by retrograde blood flow with appropriate cannula positioning, for example, when positioned in the aortic region or in the a. femoralis. However, depending on the patient, central cannulation may also be required to provide antegrade blood flow. In this case, the pulse parameters can alternatively be adjusted based on the systolic phase, resulting in pulses with the systolic phase.

[0067] The above-mentioned object is further achieved by a system for extracorporeal circulatory support of a patient. Accordingly, the system comprises a venous patient access and an arterial patient access, a non-occlusive blood pump fluidically connected to the venous patient access and the arterial patient access and designed to provide blood flow from the venous patient access to the arterial patient access, and an interface for receiving a measurement of an arterial pressure and an ECG signal of the patient. According to the invention, the system further comprises a control and regulation unit as described above.

[0068] In order to fluidically connect the non-occlusive blood pump to the venous patient access and the arterial patient access, a suction line and a delivery line can be provided in the system, each having an inner lumen with an inlet and outlet opening, whereby blood can flow from the venous patient access to the arterial patient access via the blood pump when the blood pump is in operation or activated. The suction line can also be connected to a fluid container, for example a blood reservoir, via a specific branch so that additional fluids can be administered to the patient during operation of the system, for example blood for a blood transfusion or to maintain a sufficient blood volume. The lines can be connected to the blood pump via simple connections such as Luer connections and are preferably designed, like the blood pump, as replaceable disposable items.The blood pump can be coupled to a pump drive via a pump head, for example, using a magnetic coupling. This allows the system to be used successively for multiple patients while still maintaining hygiene regulations.

[0069] The suction line and delivery line can each have a cannula, so that the system comprises a venous cannula and an arterial cannula, which are configured to be inserted into the patient's heart or cardiac region, for example, the right atrium or aortic region, respectively. This allows blood withdrawal and delivery to be precisely tailored to the anatomy of the heart and provides pulsatile blood flow precisely in the area of ​​the coronary arteries. This ensures that ejection in the systolic phase is not impaired, while perfusion of the coronary arteries is optimized in the diastolic phase.

[0070] The blood pump is further preferably a diagonal pump, as described above. This allows a low flow rate to be provided at a higher pressure or with higher pressure peaks, which is advantageous for precise synchronization with the ECG signal, in particular for synchronization with the diastolic phase. Furthermore, the impeller of such a diagonal pump has a smaller circumference and a lower mass to accelerate, so that it can be accelerated and decelerated more quickly and the pulse can be adjusted more finely. The pump output causes a blood flow to be provided from a side with a low pressure to a side with a higher pressure, wherein the control and regulating unit actuates the motor or pump drive of the blood pump and is configured to adjust the speed of the rotor blade such that the blood pump accordingly generates a wave-like flow corresponding to a pulsatile blood flow.

[0071] By providing synchronized pulsatile blood flow, the system can be used particularly for the treatment of oxygen deficiency in the myocardial tissue, for example, in cardiogenic shock. This improves coronary artery perfusion even in a pathophysiological state and with reduced oxygen supply to the myocardium, without impairing end-organ perfusion or simultaneously improving it. The system thus primarily provides therapeutic support for the patient's own cardiac output. Furthermore, the system can be used as an adjunct during catheterization procedures or as an artificial replacement and support for cardiac function in the event of a heart attack or cardiac arrest, or in cases of variable or absent cardiac output.

[0072] One or more components of the system can also be integrated into a housing. For example, the control and regulation unit can be housed in a console, which has a user interface for entering and reading system settings, in particular blood pump parameters. For example, the console can include a touchscreen and / or a display with a keyboard that can be operated by a user. The control and regulation unit operates, activates, controls, regulates, and monitors the blood pump and enables synchronization of the blood pump with the patient's cardiac cycle.

[0073] For example, the control and regulation unit can record the received ECG signal and the heart rate, with the display showing the current ECG signal graphically and the current or averaged heart rate numerically. Furthermore, characteristic properties of the ECG signal can be emphasized or highlighted in the graphic representation, so that, for example, an R-wave detected in a QRS signal for outputting the pulse parameter or pump signal can be marked in the ECG signal. Furthermore, other settings such as the time of the output pulse parameter and the pulse duration can be displayed in the ECG signal, allowing a user to monitor the control and regulation of the blood pump in relation to the patient's physiological condition.

[0074] The interface can, for example, be designed as a sensor box, which can be connected via connectors to various sensors such as pressure sensors integrated in the tube system and an ECG device.

[0075] The system preferably further comprises an ECG device that is communicatively connected to the interface. This allows the system to be used independently of the presence of other components. The ECG device can also be attached to the control and regulation unit, thus enabling a compact system. The ECG device is preferably integrated into a single housing of the system, for example, in the sensor box in the form of an ECG card or an ECG module. Alternatively, the control and regulation unit can also be configured to receive an external ECG signal from the supported patient, for example, from a cardiac monitor arranged outside the system. This allows the system to be designed even more compactly.

[0076] Furthermore, the system for extracorporeal circulatory support can be configured as a portable system, so that patient treatment does not have to take place at a fixed location and patient mobility is improved. For this purpose, the various components of the system, such as the blood pump, the control unit, and, if applicable, the ECG device, can be powered by an integrated battery. Preferably, the non-occlusive blood pump and / or the control unit, more preferably both, can be designed for extracorporeal installation. The ECG device can also preferably be designed for extracorporeal use.

[0077] To enable more precise blood sampling and delivery of pulsatile blood flow, the system preferably comprises at least one cannula for introducing the arterial patient access into the patient and one cannula for introducing the venous patient access into the patient, as described above. The required ECG signal can also be derived through the cannulas, for example, via a lead at the cannula tip or at a location on the venous cannula, a venous drainage cannula in the right atrium, or the superior vena cava, at the cannula tip or at a point of the arterial cannula in the aorta or the a. femoralis or the a. iliaca. This means that, for example, an R-wave of a QRS signal can be accurately detected and interference that can occur with a surface ECG is largely avoided.

[0078] The cannula for introducing the arterial patient access is preferably designed for insertion into the aortic region, for example into the ascending or descending aorta, or into the a. femoralis, and / or the cannula for introducing the venous patient access is designed for insertion into the right atrium or the vena cava This allows different cannula sizes and areas of the patient's circulation to be selected, depending on the patient's physiological condition.

[0079] Alternatively, the cannulas can also be designed as a single cannula or a double-lumen cannula, which can be inserted intercostally into the apex of the heart, for example. Such a cannula can be used, for example, if insertion into the aortic region is difficult or disadvantageous for the patient. Thus, depending on patient-specific circumstances, intravascular or intracardiac cannulas can be provided in the system.

[0080] The system can further comprise an oxygenator or membrane oxygenator, which enables decarburization, i.e., carbon dioxide removal or depletion, and oxygenation of the blood drawn from the patient and results in a significantly better oxygen supply to the heart muscle tissue compared to known IABP methods. For example, the system can include a membrane ventilator such as an iLA membrane ventilator, which can also be operated without a pump and solely by cardiac output because of its low resistance. Preferably, the system comprises an oxygenator that is relatively flat and angular compared to the housing and has laid mats. Alternatively, however, any gas exchanger can be provided, including a round and / or cylindrical oxygenator, for example, with wound mats. The membrane oxygenator is preferably designed to be arranged extracorporeally.

[0081] The above-mentioned problem is further solved by a method for controlling / regulating a non-occlusive blood pump of extracorporeal circulatory support. The method comprises at least the following steps: Receiving a flow value of the extracorporeal circulatory support; receiving a measurement of an arterial pressure and an ECG signal of a supported patient over a predetermined period of time; determining a mean arterial pressure of the extracorporeal circulatory support or the supported patient based on the measurement of the arterial pressure and an energy equivalent pressure based on the flow value and the arterial pressure; determining a required value of at least one pulse parameter of the blood pump based on the mean arterial pressure, the energy equivalent pressure, and the ECG signal to effect a ratio of the energy equivalent pressure to the mean arterial pressure that is greater than 1.0; and adjusting the pulse parameter depending on the ECG signal.

[0082] Preferably, an R-wave is determined from the ECG signal and the pulse parameter is adjusted at a predetermined time point after the R-wave, as described above. This enables an improvement in coronary artery perfusion in the diastolic phase and significantly reduces afterload.

[0083] Pulse parameters that can also be set include a pulse amplitude, a pump speed, a pump speed change per unit of time, a pulse duration, a systolic pump duration, a diastolic pump duration, a drive wheel deceleration, a drive wheel acceleration and / or an average pump flow.

[0084] Furthermore, a diastolic and systolic phase of the assisted patient's heart can be determined from the ECG signal, and at least the pulse duration can be adjusted so that it ends before the systolic phase. Thus, the pulse and the delivered energy-equivalent pressure are adapted to the patient's heart rate, and the potential perfusion of the coronary arteries is optimized accordingly.

[0085] Furthermore, it can be provided that a value for the ejection performance is received, which is either calculated or determined, for example, by an impedance measurement or pressure measurement.

[0086] As described above, depending on the patient, central cannulation may also be required to provide antegrade blood flow. In this case, the pulse parameters can alternatively be adjusted based on the systolic phase, resulting in pulses with the systolic phase.

[0087] The required value of the pulse parameter can further be determined depending on a given target value or a given target percentage of the energy equivalent pressure.

[0088] The flow, pressure, and / or pulse characteristic parameters can also be controlled by specifying a value of the energy equivalent pressure. For example, the current energy equivalent pressure value can be displayed on a console and then optionally adjusted to a target value. The display can also be shown as a percentage of the energy equivalent pressure, which refers to the maximum possible value of the energy equivalent pressure for the individual patient. This maximum value can, for example, be determined individually for each patient using a short empirical algorithm.

[0089] The method according to the invention preferably provides that it is carried out with a control and regulating unit according to the invention.

[0090] Further preferably, in a method according to the invention, the controlled / regulated non-occlusive blood pump (40) is fluidically connected to a membrane oxygenator, typically arranged extracorporeally. This allows oxygen-enriched and / or carbon dioxide-depleted (autologous) blood to be supplied to the patient after passing through the membrane oxygenator. The components by which the method steps are carried out are typically arranged extracorporeally. This allows individual, several, or preferably all method steps of the method according to the invention, as defined in claim 17, to be carried out non-invasively but extracorporeally.

[0091] Furthermore, the ratio of the energy-equivalent pressure to the mean arterial pressure can also be shown on a display or monitor. The energy-equivalent pressure is preferably displayed in mmHg and / or as a percentage of the maximum achievable value of the energy-equivalent pressure. Short description of the characters

[0092] Preferred further embodiments of the invention are explained in more detail in the following description of the figures. In the figures: Figure 1 is a schematic representation of a control logic of a control and regulation unit according to the invention; Figure 2 is a schematic representation of the arterial pressure curve and the corresponding flow curve over a given period of time; Figure 3 is a schematic representation of a pulse output according to the invention; Figure 4 is a schematic representation of an extracorporeal circulatory support system according to the invention; Figure 5is a schematic representation of a pulse setting according to the invention based on a received ECG signal at a normofrequency heart rate; Figure 6 is a schematic representation of an overlap of a patient pulse and the support pulse at an elevated heart rate; Figures 7A and 7B show examples of corresponding pulse parameter settings according to the invention according to Figure 5 or Figure 6 ; and Figure 8 is a schematic representation of a pulse setting according to the invention for improved afterload reduction. Detailed description of preferred embodiments

[0093] Preferred embodiments are described below with reference to the figures. Identical, similar, or equivalent elements in the different figures are provided with identical reference numerals, and a repeated description of these elements is partially omitted to avoid redundancies.

[0094] In Figure 1a control logic for a control and regulation unit is shown schematically, wherein corresponding values ​​are input from an interface 24 into a control and regulation unit level 26 and, after processing by the control logic, cause a setting on a blood pump level 28.

[0095] Accordingly, the control and regulation unit is configured to receive, via the interface 24, a flow value 10 of the extracorporeal circulatory support as well as a measurement of an arterial pressure 12 and an ECG signal 14 of a supported patient over a predetermined period of time. Based on the measurement of the arterial pressure 12, a mean arterial pressure 16 is determined. The mean arterial pressure can be determined both from an arterial pressure measurement 12 from the extracorporeal circulatory support and from the patient, for example via a tube leg of an arterial access or via an invasive pressure sensor. According to the present embodiment, the mean arterial pressure 16 of the supported patient is determined, for example either via the a. radialis or directly in the aorta.

[0096] Furthermore, a flow value 10 can be provided via a corresponding tubing leg of the arterial access, providing direct feedback regarding the blood flow in the extracorporeal circulatory support for the control and regulation unit. For example, the blood flow provided by a non-occlusive blood pump can be input via a speed and adjusted to the feedback value. In other words, a flow value 10 can first be set via a corresponding speed and then continuously or periodically corrected based on a current measurement. This allows for any different patient parameters or tubing parameters that could cause a change in the flow value 10 at the same blood pump speed to be taken into account.

[0097] The control and regulation unit further determines an energy-equivalent pressure 18 based on the received flow value 10 and the arterial pressure 16. In order to effect a pulse contribution to the mean arterial pressure, a current ratio of the energy-equivalent pressure 18 to the mean arterial pressure 16 is also determined in the control and regulation unit. For example, this ratio can initially be approximately 1.0, so that no pulse contribution or only a limited pulse contribution is achieved from the extracorporeal circulatory support. To increase the ratio and thus the pulse contribution, a required value of at least one pulse parameter 20 of the blood pump is determined in the control and regulation unit. For this purpose, not only is the ratio between the energy-equivalent pressure 18 and the mean arterial pressure 16 determined iteratively, but the received ECG signal 14 is also taken into account.

[0098] For example, a possible pulse duration can be determined from the ECG signal, which prevents an overlap of two cardiac phases of the cardiac cycle during the setting 22 of the pulse parameter 20. As a result, the pulse can be optimized over the determined possible pulse duration in order to achieve the desired or required pulse contribution and thus increase the ratio between the energy-equivalent pressure 18 and the mean arterial pressure 16, for example to 1.1 or higher. As a further pulse parameter, for example, a speed change per unit of time can be selected, which increases the pulse amplitude during the determined and thus predetermined pulse duration in such a way that a pulse with a predetermined volume or flow per unit of time is output, which causes a corresponding contribution to the mean arterial pressure.

[0099] The pulse parameters 20 are adjusted accordingly to the determined value 22, still based on the ECG signal 14, so that the pulse output is precisely timed. This allows, for example, diastolic augmentation to occur without increasing the afterload or increasing it as little as possible, especially since the pulse is output in such a way that it ends before the systolic phase of the cardiac cycle, thus keeping the remaining pressure very low.

[0100] In Figure 2An example of an arterial pressure curve (P) and the corresponding flow curve (Q) over a given period of time (t) is shown. For a time interval t1 to t2 (t2 > t1, i.e., t1 is in the past compared to t2), both an integral of the pressure change, for example, in mmHg, and an integral of the flow change, for example, in l / min, can be calculated. These serve as the basis for calculating the energy-equivalent pressure, for example, according to the following formula: EEP mm Hg = ∫ t 1 t 2 fpdt / ∫ t 1 t 2 fdt where the blood flow (f) and pressure (p) are taken into account over a time interval from t1 to t2.

[0101] Out of Figure 2 It is also evident that the increase in pressure and flow may be temporally delayed. In other words, the flow change may occur delayed after the pressure change, thus causing a latency period in extracorporeal circulatory support.

[0102] To provide pulsatile blood flow, at least one pulse parameter is set depending on an ECG signal. This allows regular pulsing and the respective pulse can be adapted to the patient's physiological condition. An example of such pulse parameters is shown in Figure 3 shown schematically.

[0103] Accordingly, two pulses are shown at regular intervals over time (t), with the pulses being calculated to an optional base flow. The base flow can be achieved by adjusting the speed to provide a blood flow between, for example, approximately 2 l / min and 3 l / min.

[0104] In the present example, the respective pulse is achieved by a change in speed, which is brought about by operating the motor and the pump drive in the blood pump. This creates a pulse amplitude 32, which causes a corresponding change in pressure (P) and thus in flow (Q). The pulse is output for a predetermined pulse duration 34, which corresponds to the determined possible pulse duration, so that the pulse duration 34 and the change in speed per unit of time result in a maximum blood flow or a maximum pulse amplitude 32. After the pulse amplitude 32 is reached, the pump drive is no longer operated, so that the speed of the blood pump is reduced to the base flow after the predetermined or set pulse duration 34. The pulses and the base flow thus result in an average flow 30, which is shown by the dashed line.The mean flow 30 is thus greater than the base flow and provides an additional contribution to cardiac output, which increases the patient's mean arterial pressure and enables sufficient perfusion or oxygenation of the end organs.

[0105] A system for extracorporeal circulatory support 36 is shown schematically in Figure 4shown. The system 36 is connected to the patient 38 and enables extracorporeal blood flow by means of a non-occlusive blood pump 40. Accordingly, the blood pump 40 is connected to the circulation of the patient 38 via a venous patient access 42 and an arterial patient access 44, so that blood is withdrawn from the patient 38 via a suction line and returned to the patient 38 via a delivery line with higher pressure. The patient accesses 42, 44 are optionally designed as cannulas and are inserted in such a way that a venous blood sampling point (cannula tip and cannula shaft) is in the area of ​​the right atrium, e.g., in the v. cava superior and / or v. cava inferior, and the arterial return is carried out via a femoral patient access 44, which is inserted either into the a. femoralisor in the aortic region to provide retrograde blood flow. Furthermore, an optional flow restrictor 46 may be provided downstream of the blood pump 40. Although this is not Figure 4 As shown, the system 36 may further comprise an oxygenator for oxygenating and / or reducing CO 2 in the venous blood, which may, for example, be arranged downstream of the blood pump 40 in the flow direction.

[0106] The blood pump 40 is operated and regulated by a control and regulation unit 48. The control and regulation unit 48 can be designed as a console comprising a user interface 50 and a monitor 52, via which settings can be made and information can be output. Thus, parameters of the system 36 can be changed and / or monitored by a user.

[0107] The control and regulation unit 48 is further coupled to an ECG device so that an ECG signal 54 can be received. The ECG device can, for example, be integrated into an interface, such as a sensor box of the system, as an ECG card or ECG module, but can alternatively also be coupled to the system as an external device via an interface. The ECG signal 54 is provided by sensors 56 attached to the patient 38, for example, as a 4-pole surface ECG, and can be displayed in single-channel or multi-channel representation with lead selection, for example, a Goldberger lead and / or extremity lead.

[0108] Although the sensors 56, shown only schematically here, may also comprise a pressure sensor, which may be arranged at a different anatomical region of the patient, in the present embodiment, a pressure sensor is provided which measures an arterial pressure of the patient 38. The pressure sensor may, for example, be coupled to an arterial cannula or be inserted as a standalone cannula, for example, into the a. radialis The pressure sensor thus outputs a pressure measurement 55, which is received by the control and regulation unit 48.

[0109] As described above, however, the pressure measurement 55 can also be determined mathematically, for example, or can be performed via a hose line or a hose leg of an arterial cannula or delivery line in order to detect a corresponding pressure value of the extracorporeal circulatory support. Optionally, several pressure sensors built into the hose set can also be provided, with one pressure sensor preferably arranged upstream of the blood pump 40, one pressure sensor downstream of the blood pump 40, and optionally another pressure sensor downstream of an optionally provided oxygenator.

[0110] The control and regulation unit 48 is configured to determine a mean arterial pressure of the patient 38 based on the received pressure measurement 55, preferably continuously and in real time. The mean arterial pressure can serve as a guideline for the blood flow provided by the blood pump 40 in order to ensure sufficient perfusion of the end organs. However, excessive blood flow can lead to retrograde blood flow closing the aortic valve and thus increasing the afterload for the systolic phase or preventing the left ventricle from ejecting blood if it is released at the wrong time of the cardiac cycle. This increases the pressure in the left ventricle, which, due to an increase in the wall tension of the left ventricle, results in a restriction of coronary blood flow. At the same time, a backflow occurs via the v. pulmonalis towards the lungs, which can lead to pulmonary edema and pulmonary dysfunction.

[0111] In order to further adapt the blood flow to the patient 38, the control and regulation unit 48 is configured to receive a flow value of the extracorporeal circulatory support from a flow sensor 58, wherein the flow sensor 58 is attached to a tubing leg of the arterial patient access 44. The flow sensor 58 thus provides feedback for the set speed on the blood pump 40. Furthermore, the flow value enables the control and regulation unit 48 to determine the energy equivalent pressure in addition to the mean arterial pressure, as described above. This allows the control and regulation unit 48 to determine the ratio between the current energy equivalent pressure and the current mean arterial pressure and iteratively determine the value of a pulse parameter required to increase the ratio above 1.0.

[0112] In this case, a heart rate determined from the ECG signal 54 and a possible resulting pulse duration are also taken into account, so that an overlap of the pulse with a cardiac systolic phase can be prevented. The control and regulation unit 48 adjusts the thus determined, at least one pulse parameter accordingly, also depending on the ECG signal 54, for example, based on the R waves determined from the ECG signal 54.

[0113] Such a pulse setting is shown schematically in Figure 5at a normofrequent heart rate of the supported patient. An ECG signal 14 comprising a QRS complex is received continuously from the patient over time (t). Although in principle any amplitude change in the ECG signal can serve as a trigger signal for the blood pump, the present embodiment uses the R wave, which characterizes cardiac output or the systolic phase of the cardiac cycle. This is also shown in the third line, which depicts the arterial pressure curve in the patient over time (t). Accordingly, immediately after the R wave, a change in arterial pressure occurs as a result of cardiac output; this change is indicated by the first two dashed lines and can, for example, be 300 ms.After an initial increase in pressure, the pressure decreases again until the aortic valves close, causing a slight increase in pressure, which is marked by Dicrot's point (closure of the aortic valve). At this point, the diastolic phase, or the filling phase of the heart, begins.

[0114] As in Figure 5 As shown, the pressure continues to decrease at a non-augmented pressure, indicated by the arrow with reference number 60. In other words, this heartbeat is not receiving any support or diastolic augmentation, nor is it pulsed by extracorporeal circulatory support, as also shown by the absence of a pulse in the extracorporeal circulatory support flow (Q) in the second line. Thus, the filling of the heart is not supported, so the arterial pressure in the patient decreases rapidly.

[0115] After the heart rate and the respective delay between the R-wave and the closure of the aortic valves have been determined, the control and regulation unit sends a signal to the pump drive or the blood pump motor to adjust a pulse parameter of the blood pump accordingly. The R-wave is used as a trigger signal and to determine the pulse parameter 20, and the respective pulse parameter 20 is adjusted after a corresponding delay 22, as schematically shown in the fourth line. Accordingly, pulses are provided in the extracorporeal circulatory support, which are timed to one another in such a way that they cause a corresponding pulsatile blood flow during the diastolic phase, as shown in the second line.This results in a pressure increase during diastole, which is included in the calculation of mean arterial pressure by a factor of 2 compared to the pressure during diastole by a factor of 1, thus contributing twice as much to the increase in mean arterial pressure. The resulting augmented pressure 62 ideally starts after the Dicrot's point and thus causes diastolic augmentation without overlapping with the systolic phase.

[0116] However, at an increased heart rate, the support pulse may overlap with the systolic phase of the patient’s pulse or the decay phase of the augmentation may interfere with the rising phase of the left ventricular ejection, the systole, as shown in Figure 6is shown schematically. At an increased heart rate, the distance between the respective R-waves and thus also between the Dicrot's point and the subsequent R-wave is shortened, so that the diastolic phase is also shortened. With a constant pulse duration, the augmented pressure 62 provided by the respective pulse can overlap or interfere with the subsequent systolic phase, which is represented by the pulse overlay 64. This leads to an increase in afterload and requires either an adjustment of the support frequency or an adjustment of the respective pulse parameters, taking into account the maintenance of a sufficient or physiological EEP.

[0117] Although an adjustment of the pulse rate may be necessary, for example in the case of tachycardia, the pulse parameters are preferably adjusted to the respective physiological state of the patient in the case of an increased heart rate, as described in the Figures 7A and 7B is shown schematically.

[0118] Thus, the pulse can be measured at a normofrequency heart rate, as in Figure 7A shown, distributed over a longer diastolic phase, so that a correspondingly longer pulse duration 34 of, for example, approximately 200 ms can be set. At the same time, the speed change per unit of time can be within a corresponding range, for example, between 2500 rpm and 4500 rpm, so that the pulse amplitude 32 also lies within this corresponding range. The pulse duration 34 can further be divided symmetrically into a systolic pulse phase 66 and a diastolic pulse phase 68 of equal duration, each of approximately 100 ms, so that a uniform pulse is output.

[0119] However, if the heart rate is increased, the pulse parameters can be modified, as described in Figure 7Bshown. The pulse duration 34 can thus be shortened and, for example, be approximately 120 ms. To still achieve a sufficient contribution to the mean arterial pressure, provided the EEP is within the physiological range, a speed change between 3000 rpm and 5000 rpm can also be selected to increase the pulse amplitude 32 and the blood flow during the pulse accordingly.

[0120] This prevents flow collisions between pump runout and the onset of systole or cardiac output, even at a heart rate between, for example, 90 and 130 beats per minute, and optimizes afterload accordingly. Furthermore, the pulse can be divided asymmetrically, with the systolic phase comprising approximately 50% to approximately 90% (not shown).

[0121] Further adaptation to higher heart rates to maintain an appropriate fluid flow can be achieved by varying the support ratio to 1:2 and 1:3 depending on the heart rate.

[0122] Adapting the pulse parameters to the patient's physiological condition thus has the advantage that, even at a higher heart rate, afterload is not increased, allowing for adequate mean arterial pressure and sufficient perfusion of the coronary arteries. Furthermore, the pulse parameters are selected to achieve a pulse contribution to the mean arterial pressure and an EEP of physiological pulse quality, so that the pulsatile blood flow results in improved perfusion and improved end-organ protection.

[0123] A further improvement in the afterload reduction can also be achieved by an appropriate pulse setting, as shown schematically in Figure 8In this illustration, the trigger signal and the control signals are not shown for clarity. However, the control unit can adjust the pulse parameter similarly to Figure 5 shown so that even in Figure 8 the R-wave from the QRS complex of the ECG signal 14 serves as a trigger signal and the pulse is delivered to the Dicrot's point, thereby causing a diastolic augmentation.

[0124] The arterial pressure (P), for example measured at an arterial limb of a tubing set, is shown schematically in the second line in Figure 8both without augmentation 76 and with augmentation 78, whereby the third line schematically shows a flow curve or the change in flow (Q) in the extracorporeal circulatory support during augmentation 78. Accordingly, without augmentation 76, a rapid reduction in arterial pressure occurs, similar to the non-augmented pressure 60 in Figure 5 . When the pulse is given, there is an increase in pressure and a slower decrease in pressure at augmentation 78, similar to the augmented pressure 62 in Figure 5 However, the pressure drop can also appear differently, so the shape of the pressure curves is only to be understood as a schematic example and can look different in each patient, for example, depending on patient-related factors such as vascular elasticity, degree of calcification, cardiac strength, viscosity, etc.

[0125] The corresponding flow (Q) in the extracorporeal circulatory support is calculated to a base flow 72 and results in a maximum flow value due to the pulse, which is predetermined by the speed change per unit of time and the pulse duration and thereby causes a diastolic augmentation 70 compared to the base flow 72.

[0126] To reduce afterload, the speed after the diastolic pumping phase (systolic cardiac phase) is set further below the corresponding value of the base flow 72, resulting in a minimum flow 74 below the base flow 72 and correspondingly reducing the pressure in the systolic phase of the heart, as also results from the pressure reduction during the systolic phase in the augmentation curve 78. This continues to provide pulsatile blood flow with a required mean flow 30 and a base flow 72, while simultaneously allowing the afterload to be further reduced. This can be achieved, as shown in the illustration, with reference to the next cardiac output, but also to the one after that with a 1:2 support ratio.

[0127] Although this is not Figure 8As shown, the speed can alternatively be changed such that a negative blood flow occurs. For example, such a negative blood flow can be provided by a four-quadrant control, as described above. Such a four-quadrant control not only enables a reduction (braking) of the speed or the impeller, but also enables a reversal of the direction of rotation by reversing the current direction, so that the impeller can rotate in the opposite direction. Such a negative blood flow at the end of the pulse and during the systolic phase thus causes a suction effect, which further facilitates cardiac ejection and thus further reduces the afterload for the stressed and supported heart. List of reference symbols

[0128] 10Flow value 12Arterial pressure measurement 14ECG signal 16Determining the mean arterial pressure 18Determining the energy-equivalent pressure 20Determining the blood pump pulse parameter 22Setting the pulse parameter 24Interface 26Control unit level 28Blood pump level 30Mean flow 32Pulse amplitude 34Pulse duration 36Extracorporeal support system 38Patient 40Blood pump 42Venous patient access 44Arterial patient access 46Optional flow limiter 48Control unit 50User interface 52Monitor 54ECG signal 55Pressure measurement 56Sensors 58Flow sensor 60Non-augmented pressure 62Augmented pressure 64Pulse overlap 66Systolic pulse phase or pump duration 68Diastolic pulse phase or pump duration 70diastolic augmentation 72base flow 74minimum flow 76pressure without augmentation 78pressure with augmentation

Claims

1. Control and regulation unit (48) for a non-occlusive blood pump (40) of an extracorporeal circulatory support, configured for - receiving a flow value (10) of the extracorporeal circulatory support; - receiving a measurement of an arterial pressure (12) and an ECG signal (14) of a supported patient (38) over a predetermined period of time; - determining a mean arterial pressure (16) of the extracorporeal circulatory support or of the supported patient from the measurement of the arterial pressure (12) and an energy equivalent pressure (18) from the flow value (10) and the arterial pressure (16); and - determining a required value of at least one pulse parameter (20) of the blood pump (40) from the mean arterial pressure (16), the energy equivalent pressure (18), and the ECG signal (14) so as to effect a ratio of the energy equivalent pressure (18) to the mean arterial pressure (16) of greater than 1.0, wherein the control and regulation unit (48) is further configured to adjust (22) the pulse parameter (20) in dependence on the ECG signal (14).

2. Control and regulation unit (48) according to claim 1, wherein the at least one pulse parameter (20) is selected from a pulse amplitude (32), a pump speed, a pump speed change per unit time, a pulse duration (34), a systolic pump duration (66), a diastolic pump duration (68), a drive wheel deceleration, a drive wheel acceleration and / or an averaged pump flow.

3. Control and regulation unit (48) according to claim 1 or 2, further configured to determine the required value of the pulse parameter (20) from a given or predetermined mean arterial pressure (16) and / or a given or predetermined energy equivalent pressure (18), and / or further configured to determine the flow value (10) based on an inputted speed of the blood pump (40) or automatically based on a received flow measurement.

4. Control and regulation unit (48) according to any of the above claims, wherein the flow value (10) comprises a patient-specific mean flow (30), wherein the control and regulation unit (48) is further configured to adjust (22) the mean flow (30) based on a received flow measurement and the mean arterial pressure (16) by adjusting (22) the pulse parameter (20), and / or which is further configured to determine a heart rate from the ECG signal (14) and to adjust (22) the pulse parameter (20) in dependence of the heart rate.

5. Control and regulation unit (48) according to any of the preceding claims, which is further configured to determine an amplitude change from the ECG signal (14) and to adjust (22) the pulse parameter (20) at a predetermined time point after the amplitude change, preferably configured to determine an R-wave from the ECG signal (14) and to adjust the pulse parameter (20) at a predetermined time point after the R-wave (22).

6. Control and regulation unit (48) according to any of the preceding claims, which is further configured to receive a measurement of an aortic pressure of the supported patient (38) over a predetermined period of time and to adjust (22) the pulse parameter (20) at a predetermined aortic pressure and change of aortic pressure.

7. Control and regulation unit (48) according to any of preceding claims 5 to 6, which is further configured to determine a diastolic phase and a systolic phase of the heart of the supported patient (38) from the ECG signal (14) and / or the aortic pressure and to adjust (22) at least the pulse duration (34) such that it ends before the systolic phase.

8. System (36) for extracorporeal circulatory support of a patient (38) comprising: - a venous patient access (42) and an arterial patient access (44), - a non-occlusive blood pump (40) fluidically connected to the venous patient access (42) and the arterial patient access (44) and configured to provide blood flow from the venous patient access (42) to the arterial patient access (44), - an interface (24) for receiving a measurement of an arterial pressure (12) and an ECG signal (14; 54) of the patient (38), and - a control and regulation unit (48) according to any of the preceding claims 1 to 7.

9. System (36) according to claim 8, further comprising a membrane oxygenator and / or a membrane fan, preferably an extracorporeally arranged membrane oxygenator.

10. System (36) according to claim 8 or 9, further comprising an ECG apparatus communicatively connected to the interface (24), and / or wherein the non-occlusive blood pump (40) and / or control and regulation unit (48), preferably both, are configured to be arranged extracorporeally.

11. System (36) according to any of preceding claims 8 to 10 further comprising at least one cannula for inserting the arterial patient access (44) into the patient (38) and one cannula for inserting the venous patient access (42) into the patient (38), wherein the cannula for inserting the arterial patient access (44) is preferably formed for insertion into the aortic region or into the a. femoralis and / or the cannula for inserting the venous patient access (42) is formed for insertion into the right atrium or into the vena cava.

12. Method for controlling / regulating a non-occlusive blood pump of an extracorporeal circulatory support, comprising the steps of: - receiving a flow value (10) of the extracorporeal circulatory support; - receiving a measurement of an arterial pressure (12) and an ECG signal (14) of a supported patient (38) over a predetermined period of time; - determining a mean arterial pressure (16) of the extracorporeal circulatory support or of the supported patient from the measurement of the arterial pressure (12) and an energy equivalent pressure (18) from the flow value (10) and the arterial pressure (12); - determining a required value of at least one pulse parameter (20) of the blood pump from said mean arterial pressure (16), the energy equivalent pressure (18), and the ECG signal (14) so as to effect a ratio of the energy equivalent pressure (18) to the mean arterial pressure (16) of greater than 1.0; and - adjusting (22) the pulse parameter (20) in dependence on the ECG signal (14).

13. Method according to claim 12, wherein an R-wave is determined from the ECG signal (14) and the pulse parameter (20) is adjusted (22) at a predetermined time point after the R-wave.

14. Method according to claim 12 or 13, wherein the at least one pulse parameter (20) is selected from a pulse amplitude (32), a pump speed, a pump speed change per unit time, a pulse duration (34), a systolic pump duration (66), a diastolic pump duration (68), a drive wheel deceleration, a drive wheel acceleration, and / or an averaged pump flow, and / or wherein a diastolic phase and systolic phase of the heart of the supported patient (38) are determined from the ECG signal (14) and at least the pulse duration (32) is adjusted (22) such that it ends before the systolic phase, and / or wherein the required value of the pulse parameter (20) is further determined in dependence of a predetermined target value or a predetermined target percentage of the energy equivalent pressure (18).

15. Method according to any of preceding claims 12 to 14, wherein the non-occlusive blood pump (40) is fluidically connected to a membrane oxygenator.

16. Method according to any of preceding claims 12 to 15, wherein the method steps, as defined by claim 12, are carried out non-invasively.

17. Method according to any of preceding claims 12 to 16, wherein the method is carried out with a control and regulation unit (48) according to any of claims 1 to 7.