Cardiac support or backup device

Description

Technical field.

[0001] The present invention relates to a device for assisting or supplementing the heart.

[0002] The term assistance is used when a portion of the blood arriving at the heart is taken away, while the term replacement is used when all, or almost all (>90%), of the blood arriving at the heart is taken away by the heart replacement device.

[0003] It concerns the technical field of devices used by extracorporeal circulatory assistance / supplementation which maintain hemodynamics compatible with life, in the context of cardiac muscle failure, or more generally cardiopulmonary failure, threatening the patient's life prognosis, following acute or chronic heart failure (coronary insufficiency and / or cardiovascular disease, or other associated pathology). State of the art.

[0004] The heart's function is to distribute blood throughout the body, transporting oxygen and nutrients necessary for the functioning of the various organs, and to transport metabolic waste products to the lungs, the organs for CO2 elimination, and urea and creatinine to the kidneys. It is divided into two parts: the right side of the heart, which receives deoxygenated blood (depleted of oxygen) via the vena cava and pumps it through the right ventricle to the pulmonary artery, where it is oxygenated and CO2 released; and the left side of the heart, which receives oxygenated blood via the pulmonary veins and pumps it through the aorta to the various organs. The contractions of the heart muscle (the pump) generate a pulsating blood flow throughout the body. This allows for the continuous regulation of blood flow to meet the physiological needs of the body and each organ.

[0005] In the presence of cardiac muscle failure (for example in the case of a myocardial infarction), it is necessary to temporarily replace the pumping function of said cardiac muscle with a mechanical circulatory support system, the objective being to ensure good blood circulation while the heart recovers.

[0006] Among the circulatory support systems currently in use are implantable circulation devices, such as the one described in US patent 2003 / 032853, but their implantation is particularly complex and risky. Extracorporeal circulation (ECC) devices are also known. These bypass a failing heart using a servo-controlled pump located outside the body, which receives blood from the vena cava and injects it into the aorta (thoracic vein), mechanically generating a continuous blood flow compatible with the body's vital needs, thus enabling the operation to proceed. The primary objective of circulatory support is usually to regulate the pumped flow rate to perfectly match the patient's physiological needs, which vary continuously.

[0007] As a power source, a console or central processing unit incorporates a mathematical model based on the physical laws governing fluid movement in a closed circuit. This circuit typically consists of a pump, a heat exchanger, a flow meter, a blood gas and electrolyte analyzer, a pressure sensor, and biocompatible equipment such as tubing, arterial and venous cannulas, a venous reservoir, an oxygenator, and an arterial filter. Usually, a centrifugal or peristaltic pump is used as the arterial head pump, and four additional peristaltic pumps are used for cardiotomy suction, cardiac chamber circulation, cardioplegia administration, and as a backup pump. If a failure persists, ECMO (Extracorporeal Membrane Oxygenation) or ECLS (Extracorporeal Life Support) is indicated.These systems are simpler than standard cardiopulmonary bypass (CPB) and are portable, with a usage time of several days, unlike conventional CPB. Indeed, unlike conventional CPB, ECMO and ECLS are maintained until the patient's cardiopulmonary recovery or as a bridge before transplantation.

[0008] Although these systems have demonstrated their effectiveness in their conventional circulatory support functions, they are not entirely satisfactory because the regulation of the generated blood flow is not optimal with regard to the patient's physiology. The current survival rate for some of the conditions they address does not exceed 30%. A fatigued or failing heart will become even more fatigued if aortic pressure is high during systole: the aortic valves do not open properly, afterload increases, and so does left ventricular end-diastolic pressure, with a lack of ejection. Mitral valve regurgitation may also occur, potentially leading to irreversible pulmonary edema. This is one of the main limitations of these prior art systems.

[0009] Patent document WO2006 / 016047 (BERTHIER) describes an extracorporeal circulatory support device for the heart equipped with a linear motor configured to move a piston in a reservoir, thereby generating a blood flow characterized by alternating aspiration and ejection phases. The linear motor is servo-controlled to drive the piston's displacement, speed, and acceleration according to the patient's physiological needs. However, it has been shown that the match between the pulsatile flow and the patient's physiological needs is not optimal and does not achieve the desired performance.

[0010] An implantable circulatory support device was also described in US patent 2003 / 032853 A1 (KORAKIANITIS THEODOSIOS). This device comprises a linear actuator designed to induce pulsatile blood flow by moving an organ within a chamber. It is designed to support cardiac activity by ejecting blood into the systemic circulation, based on a control law that takes into account the patient's hemodynamic parameters (such as ECG signals). However, it does not implement a succession of distinct aspiration and ejection phases, nor does it allow for feedback control of the ejection pressure according to physiological needs assessed in real time. Furthermore, the device is implantable, which makes its placement invasive and more complex.

[0011] The present invention aims to overcome all or part of the aforementioned drawbacks. In particular, one objective of the invention is to provide an assistance or supplementary device whose operation maximizes the benefits of the assistance or supplementary device.

[0012] Another objective of the invention is to optimize the real-time synchronization of ejection phases with regard to the patient's physiology in order to maintain hemodynamics compatible with the life of said patient.

[0013] Yet another objective of the invention is to provide an assistance or supplementary device whose control is simple to implement and particularly reliable and robust to changes in the patient's physiological parameters.

[0014] The invention also aims to enable better performance and increased accuracy of pumping cycles. Presentation of the invention.

[0015] The solution proposed by the invention is a device for temporary extracorporeal circulatory support or replacement of a patient's heart, comprising: a linear actuator configured to move an organ in translation within a chamber so as to cause a pulsating fluid flow capable of supporting the activity of the patient's heart, which flow is characterized by a succession of aspiration phases and fluid ejection phases, a control unit configured to pilot the actuator and control the movement of the organ, which piloting is carried out by taking into account input data from measurements of the patient's hemodynamic parameters, during each ejection phase the unit moving the organ a distance D.

[0016] The control unit is configured to: Sequence the displacement D into an integer number N of cycles Ci, with N > 1, each cycle having the same duration Δt. On each cycle Ci, calculate a displacement length Li of the organ such that D = ∑ i = 1 N Li , which length L i is calculated based on input data from: ∘ one or more hemodynamic parameters of the patient measured during cycle C i , or during one or more cycles prior to said cycle C i and close to said cycle C i , ∘ the value of a setpoint pressure Π i based on the Franck-Starling law, on each cycle C i , transmit an instruction to the actuator to move the organ by the calculated length L i.

[0017] During each cycle Ci, the control of the length Li allows for optimal adjustment of the organ's displacement velocity according to the patient's hemodynamic parameters and, consequently, the pressure of the ejected fluid during that cycle to the patient's physiological needs. As a result, during each ejection phase, the ejected fluid pressure can be very precisely aligned with the Frank-Starling curve in response to changes in the patient's hemodynamic parameters at each cycle Ci. According to the Frank-Starling law, the greater the stretch of the cardiac muscle fibers (produced by preload), the higher the stroke volume. In other words, the volume of blood ejected per beat will be proportional to the volume of blood contained in the ventricle at the end of diastole.

[0018] According to the invention, the ejection pressure is not adjusted to a constant target pressure value, but rather to the patient's actual needs to achieve the Frank-Starling mechanism. The device according to the invention thus maximizes the benefits of the assistance or replacement therapy, while remaining particularly reliable and robust to changes in the patient's physiological needs. The applicant has found, surprisingly, that the survival rate after 15 days of assistance or replacement therapy with the device according to the invention can reach 80%.

[0019] Furthermore, by sequencing the organ's movement during the ejection phase and adapting its displacement length for each cycle of the sequence, the pumping cycles achieve increased measurement accuracy. Indeed, the ejection pressure value on the Frank-Starling curve is monitored at each cycle; in other words, the patient's responsiveness is checked at each cycle before introducing fluid into the patient's body. This high-frequency control (N times greater than the heart rate) ensures optimal matching between the pulsatile flow and the patient's physiological needs.

[0020] Other advantageous features of the apparatus of the invention are listed below. Each of these additional features may be considered alone or in combination with the notable features defined above. Each of these additional features contributes, where applicable, to solving specific technical problems defined further in the description and in which the notable features defined above do not necessarily participate. These additional features may, where appropriate, be the subject of one or more divisional patent applications.

[0021] According to one embodiment, the control unit is configured to calculate the displacement length L i according to the following formula: L i = L i − 1 + Π i − PR i − Δ P cor × Δ t 2 S × μ

[0022] Or L i-1 is the displacement length over cycle C i-1; Π i is the setpoint pressure over cycle C i; PR i is the actual fluid pressure measured in the chamber over cycle C i; ΔP cor is a correction pressure value, S is the equivalent area of ​​the surface of the component in contact with the fluid, µ is a constant.

[0023] According to one embodiment, the device comprises an ejection cannula implantable in the human body and fluidically connected to the chamber, and in which the control unit is configured to calculate the displacement length L i according to the following formula: L i = L i − 1 + Π i − PR i × Δ t 2 S × μ

[0024] Or L i-1 is the displacement length on cycle C i ; Π i is the setpoint pressure on cycle C i ; PR i is the actual fluid pressure measured at the ejection cannula (16); S is the equivalent area of ​​the surface of the organ in contact with the fluid, µ is a constant.

[0025] According to one embodiment, the control unit is configured to generate, for each ejection phase, a setpoint pressure curve according to the Franck-Starling law, the value of Π i corresponding to the value of the pressure on said curve over the cycle C i.

[0026] According to one embodiment, the setpoint pressure curve according to the Franck-Starling law is identical for each ejection phase.

[0027] According to one embodiment, the setpoint pressure curve according to the Franck-Starling law is redefined at each ejection phase or redefined at regular time intervals as a function of arterial pressure values ​​measured at a pre-parameterized frequency, at the level of an extremity of the patient's body.

[0028] According to one embodiment, the control unit is configured to calculate, at each cycle Ci, the setpoint pressure Pi according to the following formula: Π i = PM s − K s ⋅ VE i − 1 − R s ⋅ Q i − 1

[0029] Or : PM S is the maximum systolic pressure whose value is pre-parameterized and / or set; KS is a coefficient between 0.9 and 1.1; VE i-1 is the volume of fluid set in motion in the chamber during the ejection phase in the previous cycle C i-1; RS is a coefficient between 0.4 and 0.6; Q i-1 is the flow rate of fluid ejected from the chamber in the previous cycle C i-1.

[0030] According to one embodiment, in which the number N of cycles Ci is between 45 and 55, and the duration Δt of each cycle Ci is between 0.004 s and 0.006 s.

[0031] According to one embodiment, the control unit is configured to separate the aspiration and ejection phases of the fluid so that during an artificial cardiac cycle, the volume aspirated during the aspiration phase is different from the volume ejected during the ejection phases.

[0032] According to one embodiment: The first C1 cycle coincides with natural systole, and starts at the instant a QRS complex of an ECG signal measured on the patient is detected, or the first C1 cycle is offset from natural systole and starts after an offset of between 0.2 and 0.3 seconds from the instant a QRS complex of an ECG signal measured on the patient is detected. Brief description of the figures.

[0033] Other advantages and features of the invention will become clearer upon reading the description of a preferred embodiment which follows, with reference to the attached drawings, which are provided as illustrative and non-limiting examples and on which: [ Fig. 1 [ ] is a schematic view of the temporary circulatory support or replacement device for a patient's heart according to the invention. Fig. 2 [ ] is a schematic view of a pumping system that can be used in a device according to the invention. Fig. 3] illustrates the sequencing of organ movement during an ejection phase. Fig. 4 ] illustrates a setpoint pressure curve according to the Frank-Starling law. Description of the implementation methods.

[0034] The device of the invention is intended for use in a degraded hemodynamic situation that directly threatens the life of a patient 12 (for example, with a tissue perfusion pressure - PF - of less than 50 mm Hg). It allows for the assistance or replacement of the heart 13 of patient 12.

[0035] By referring to the figure 1 The device of the invention comprises a pumping system that allows the heart muscle to be partially or totally supplemented by aspirating a sufficient quantity of blood during an artificial diastole phase and reinjecting it during an artificial systole phase.

[0036] Table 1 below illustrates an example of the sequencing of these aspiration and ejection phases: [Table 1] Natural heart Device that is the subject of the invention natural systole Aspiration Onset of natural diastole Ejection Natural diastole Aspiration

[0037] The sequencing of these aspiration and ejection phases is described in more detail later in the description.

[0038] According to one embodiment, the operator introduces an inlet cannula 15 (e.g., 21 or 23 French or "Fr," the FRENCH representing 1 / 3 of a millimeter) adapted to draw blood from the venous system of patient 12 and an ejection cannula 16 (e.g., 17 or 19 French) adapted to inject the blood into the arterial system of said patient 12. The inlet cannulas 15 and ejection cannulas 16 commonly used in the extracorporeal circulation (ECC, ECMO, ECLS) market are compatible with the invention. Reinforced cannulas are preferably used to prevent suction collapse and / or kinking of said cannulas, which would lead to a reduction in flow.

[0039] The operator can perform: Preferably, a right heart / left heart setup, of the veno-arterial ECMO type, is used in cases of right heart and left heart failure, by positioning the inlet cannula 15 at the level of a vena cava and the outlet cannula 16 at the level of the aorta. In this latter case, the lungs are also bypassed and an oxygenation system 17 (preferably including a heat exchanger) is placed in the bypass circuit to remove CO2 from the blood and charge it with O2 before reinjecting it into the body. This oxygenation system 17 is advantageously located after the reservoir 11, i.e. on the outlet portion 18 of the bypass circuit; possibly a right heart / right heart setup for partial assistance in case of right heart failure, by positioning the inlet cannula 15 at the level of a vena cava and the outlet cannula 16 at the level of the pulmonary artery.

[0040] Cannulas 15 and 16 can be inserted percutaneously in a cardiac and / or vascular catheterization lab, in an intensive care unit, or by a SAMU (Mobile Circulatory Assistance Unit for UMAC and Emergency Medical Assistance Service for SAMU) team. They are introduced through a peripheral blood vessel and guided close to the heart 13, at the level of the targeted veins or arteries. They can also be inserted surgically in a surgical suite, using a combination of percutaneous puncture and surgical opening of the vessels.

[0041] These cannulas 15, 16 are connected to the pumping system by catheter-type tubes, also compatible with those usually used for CEC (for example 3 / 8 gauge) to form on the one hand the intake portion 19 and the ejection portion 18 of the bypass circuit.

[0042] The device comprises a chamber 11 forming a reservoir (equivalent to an artificial external ventricle) and allowing the temporary storage of a volume of fluid (e.g., blood, blood substitute, blood + blood substitute). A linear actuator 10 is configured to move an organ 70 in translation within the chamber 11 so as to induce a pulsatile fluid flow capable of supporting the activity of the heart 13. This flow is characterized by a succession of aspiration and ejection phases of the fluid.

[0043] As an example, and to give an idea of ​​the scale, for a heart beating at 70 bpm (beats per minute), the average flow rate is approximately 5 L / min during both ejection and intake. Ejection is brief (high flow rate), while aspiration is slower (approximately half the flow rate). The typical blood volume is about 70 cm³ or 70 ml per beat. In 1 minute (70 beats), the pumped volume is: 70 x 70 ml = 4900 ml, or 4.9 L. The 70 ml are ejected between 0.2 s and 0.35 s (depending on the patient), so the maximum instantaneous flow rate during ejection is 270 ml / s or 16 L / min max. This flow rate is roughly half that during aspiration. Thanks to the device according to the invention, it is possible to increase (or possibly reduce) the quantities of blood aspirated / ejected to correspond as closely as possible to the actual functioning of a heart and the needs of the body.

[0044] In one embodiment, the actuator 10 is a linear motor or any other equivalent means (for example, a cylinder) for moving the member 70. In another embodiment, the member 70 is a diaphragm fixed, on its periphery, to the inner wall of the chamber 11. It is advantageously made of a flexible and elastic material, such as an elastomer. The diaphragm 70 houses a central insert, not visible in the accompanying figures, equipped with an actuating arm forming a piston 31, this piston being fixed to the actuator 10 by means of a mechanical engagement, fastening, or connection. In another embodiment, the member 70 is in the form of a piston of the type described in the aforementioned patent application WO2006 / 016047.

[0045] In one embodiment, the actuator 10 is configured to execute high-velocity displacement commands in order to adjust the actual fluid motion to these commands. This can be achieved using a linear motor with a short time constant (advantageously less than or equal to 10 ms) and capable of generating a significant force (advantageously greater than or equal to 500 N), to which a lever arm system can be added for further acceleration. The fluid can then be abruptly set in motion and stopped at the desired moment. The physical flow regime of the fluid can thus closely follow the control signals.

[0046] A control unit 40 is provided to automatically control the actuator 10 and control the movement of the component 70. This unit 40 may be in the form of a processor, microprocessors, CPU (for Central Processing Unit) integrated into a computer, a calculator or similar means.

[0047] Each step of the actuator 10's movement corresponds to an internal volume of the chamber 11. This correspondence between the step of the actuator 10 and the internal volume of the chamber 11 is stored or recorded in a memory area of ​​the unit 40. By controlling the movement of the actuator 10, it is therefore possible to control very precisely the volume of fluid aspirated and ejected by the device during the successive phases of aspiration and ejection.

[0048] In one embodiment, one or more sensors 2 are used to acquire an electrocardiographic signal (ECG signal) corresponding to the electrical activity of the heart 13. The control unit 40 is configured to control the movement of the actuator 10 based on input data from measurements of the patient's hemodynamic parameters, including this ECG signal. The input data may come from other measurements of hemodynamic parameters, such as the pressure at the cannula 15 and / or cannula 16, blood oxygen saturation (measured using a pulse oximeter), pressure or volume variations in an organ or limb segment (measured using a plethysmograph), etc. In a preferred embodiment, this data is acquired in real time at a high frequency of 200 Hz. Artificial systole

[0049] With reference to the figure 2, during each phase of fluid ejection out of chamber 11 (each artificial systole), unit 40 moves organ 70 by a distance D. This distance D is not necessarily constant, but can be variable at each artificial systole.

[0050] According to one embodiment, the displacement D is sequenced into an integer number N of cycles Ci, with N > 1, each cycle having the same duration Δt. figure 3This illustrates, by way of explanation, a sequencing of displacement D in 5 cycles Ci (N=5). The number N is advantageously between 2 and 100, preferably between 10 and 80, and most preferably between 20 and 60. According to a preferred embodiment allowing an excellent compromise in terms of calculation accuracy and the computing resources required for said calculations, N is between 45 and 55, preferably equal to 50. The duration Δt is advantageously between 0.1 s (second) and 0.003 s, preferably between 0.004 s and 0.006 s, and most preferably equal to 0.005 s. Thus, the unit 40 defines at each cycle Ci the patient's need 12 and precisely controls the supply of the patient's body according to this same need.

[0051] The maximum beat duration (systole) is approximately 0.3 s. In one embodiment, to accurately discretize this dynamic range, it is divided into a minimum of 20 steps, resulting in a maximum time step of 0.3 / 20 = 0.015 s, corresponding to a minimum frequency of 60 Hz. A good compromise is achieved with a precision margin of 200 Hz, allowing the movements of organ 70 to be controlled with an accuracy of 0.005 s. This notably results in optimal and precise tempo adjustment to the hundredth of a second.

[0052] In one embodiment, the first C1 cycle coincides with natural systole and begins, for example, at the instant the QRS complex of the ECG signal is detected. In another embodiment, the first C1 cycle is offset from natural systole by a delay of approximately 0.25 seconds (i.e., between 0.2 and 0.3 seconds) from the instant the QRS complex of the ECG signal is detected. This time delay, made possible by the high calculation rate, ensures that the aortic valve is fully closed when artificial systole is initiated, thus protecting the ventricle from any overload.

[0053] On each cycle C i , unit 40 calculates a displacement length L i of the organ 70 such that D = ∑ i = 1 N Li For example, the figure 3, D=L 1 +L 2 +L 3 +L 4 +L 5 , with L 1 ≠L 2 ≠L 3 ≠ 4 ≠L 5 . Each length L i is calculated based on input data from one or more hemodynamic parameters of patient 12 measured during cycle C i and a setpoint pressure value on said cycle C i . As soon as the length L i is calculated, the unit 40 transmits an instruction to the actuator 10 to move the organ 7) by said length.

[0054] The hemodynamic parameters measured during cycle Ci may be all or some of those mentioned previously (ECG signal, pressure at cannulas 15 and 16, oxygen saturation, pressure or volume variations, etc.). These parameters are therefore acquired in each time interval Δt, allowing for real-time monitoring of patient 12's condition and corresponding control of organ 70, which is extremely precise over distance D.

[0055] In addition, on each cycle C i, unit 40 determines a setpoint pressure based on the Frank-Starling law.

[0056] According to a first embodiment, unit 40 generates, for each ejection phase, a setpoint pressure curve initially governed by the Frank-Starling law. figure 4 illustrates a pressure variation curve according to the Frank-Starling law, which typically has a bell-shaped curve passing through a maximum pressure PM S. This setpoint curve can be identical for each ejection phase or, preferably, be redefined at each ejection phase or at regular time intervals (for example, every 5 minutes or every 10 minutes), depending on the patient's condition, notably defined by measurements of hemodynamic parameters and / or a calculation algorithm based on an artificial intelligence model, while remaining based on the Frank-Starling law as claimed.

[0057] Advantageously, the setpoint curve is redefined at regular time intervals based on blood pressure values ​​measured at a pre-set frequency at a specific extremity of the patient's body, for example, at the tip of a finger. This pressure measurement provides information on the patient's physiological state12 and allows for recalibration of the setpoint curve. Indeed, the systemic resistances and compliances of the patient's body change throughout the day and / or depending on any medication the patient may be taking. The artificial pulsatile flow can thus be easily adapted to these variations. If the pressure measured at the patient's extremity is below a threshold value (for example, below PMS x 10%), then the PMS value (i.e., the peak of the bell curve) is increased.Conversely, if the pressure measured at the patient's extremities exceeds a threshold value (e.g., greater than PM S x 10%), then the PM S value is reduced. This ensures that fluid penetrates deeply into the body and adequately irrigates the patient's organs throughout the duration of circulatory support or replacement therapy.

[0058] The setpoint curve is defined over the period NxΔt, where N is the number of cycles Ci and Δt is the time interval of each cycle. Each cycle Ci corresponds to a setpoint pressure value Pi, which is used by unit 40 as input data for calculating the length Li. The pressure of the ejected fluid thus closely follows this setpoint curve.

[0059] According to a second embodiment, unit 40 calculates, at each cycle Ci, the setpoint pressure Pi according to the formula [Math 1] based on the Frank-Starling law: Π i = PM s − K s ⋅ VE i − 1 − R s ⋅ Q i − 1

[0060] Or : PM S is the maximum systolic pressure, the value of which is pre-parameterized and / or set by the practitioner or automatically, for example equal to 140 mmHg; as explained previously, the value of PM S can be redefined at regular time intervals based on blood pressure values ​​measured at a pre-parameterized frequency, at one extremity of the patient's body; KS is a systolic coefficient between 0.9 and 1.1; VE i-1 is the volume of fluid set in motion in chamber 11, during the ejection phase, in the cycle preceding C i-1; RS is a coefficient expressing the circulatory resistance in the arteries, between 0.4 and 0.6; Q i-1 is the flow rate of fluid ejected from chamber 11 in the cycle preceding C i-1.

[0061] According to a third embodiment, the setpoint pressure value Π i is determined by means of a calculation algorithm based on an artificial intelligence model.

[0062] According to one embodiment, the length L i is determined according to the following formula [Math 2]: L i = L i − 1 + Π i − PR i − Δ P cor × Δ t 2 S × μ

[0063] Or L i-1 is the displacement length on the previous cycle C i-1; Π i is the setpoint pressure on cycle C i; PR i is the actual fluid pressure measured in chamber 11 on cycle C i; ΔP cor is a correction pressure value, S is the equivalent area of ​​the surface of organ 70 in contact with the fluid, µ is a constant (corresponding to blood inertness).

[0064] The pressure PR i is advantageously measured by means of a pressure sensor installed in chamber 11. To provide more security to the device, this pressure measurement in chamber 11 can be correlated with one or more other pressure measurements by sensor installed at the cannula 15 and / or cannula 16. If one of these different sensors malfunctions or becomes defective, it is then possible to stop the pressure regulation method according to the invention and / or to replace said method with another regulation method (for example, volume regulation).

[0065] The corrective pressure value ΔPcor can be fixed or variable at each cycle Ci. For example, it ranges from 10 mmHg to 50 mmHg (1 mmHg = 133.322 Pa). Taking into account factors such as the length and elasticity of the tubes connecting chamber 11 to cannula 16, any devices 17 installed in the bypass circuit, the viscosity of the fluid, and the elasticity of the aorta, the pressure PRi measured in chamber 11 may differ from the actual fluid pressure in the aorta. The corrective pressure ΔPcor provides this correction.

[0066] The value of µ represents the inertia of the displaced fluid column, which value is for example between 10 6< kg / m 4< and 5.10 6< kg / m 4< .

[0067] According to another embodiment, the pressure PR i corresponds to the actual fluid pressure measured at the ejection cannula 16 during cycle C i. In this case, good results are obtained even without applying a pressure correction (ΔP cor = 0), so that the length L i can be determined according to the following formula [Math 3]: L i = L i − 1 + Π i − PR i × Δ t 2 S × μ

[0068] Or L i-1 is the displacement length on the previous cycle C i-1; Π i is the setpoint pressure on cycle C i; PR i is the actual fluid pressure measured at the cannula 16 on cycle C i; S is the equivalent area of ​​the surface of the organ 70 in contact with the fluid, µ is a constant (identical to that of formula [Math 2]).

[0069] Thanks to the invention, the fluid expelled during artificial systole penetrates deep into the patient's body12, according to their physiological needs, in order to irrigate the various organs. The flow / pressure profile of the ejected fluid combines with arterial compliances to ensure that the systolic wave propagates correctly throughout the body.

[0070] The best results are obtained when the acquisition of hemodynamic parameters and actual pressures, the determination of the setpoint pressure, the generation of the displacement command Li, and its execution by unit 40 are carried out during cycle Ci. However, according to another embodiment that still gives acceptable results, the acquisition of hemodynamic parameters and actual pressures, the determination of the setpoint pressure, the generation of the displacement command Li, and its execution by unit 40 can be carried out during one or more cycles preceding cycle Ci and close to said cycle Ci. By close, we mean three or four cycles preceding cycle Ci, during which the displacement command is executed.For example, the acquisition of hemodynamic parameters can be performed in cycle C i-3, the actual fluid pressure also measured in cycle C i-3, the calculation of the setpoint pressure Π i-2 performed during cycle C i-2, and the development of the displacement command Li performed in cycle C i-1 to be executed in cycle C i. Those skilled in the art will understand that other time sequences are possible. Artificial diastole

[0071] The execution of artificial diastole is independent of that of artificial systole. According to one embodiment, the admission of blood into chamber 11 (artificial diastole) is carried out throughout the natural cardiac cycle, except for the time dedicated to artificial systole.

[0072] According to one embodiment, the artificial diastole time is about 0.5 seconds (about 1 / 3 of the artificial systole), with organ 70 moving about 1 cm.

[0073] According to one embodiment, the fluid flow rate aspirated into chamber 11 during artificial diastole is significant, even very significant, at least 2 to 3, or even 5 liters per minute. Given this high aspiration rate, the risk of collapse of the patient's vena cava 12 is significant, even very significant. It is therefore particularly important to manage the aspiration as precisely as possible, that is, to be able to modulate or modify the aspiration of fluid drawn from the patient's vena cava in order to avoid any risk of collapse.

[0074] To achieve this objective, a pressure sensor is advantageously installed in the inlet portion 19 (at the inlet cannula 15 or between said cannula and chamber 11). This pressure sensor has a measurement time constant of less than 20 milliseconds, advantageously less than 10 milliseconds, in order to detect the pressure and any changes therein very quickly and regularly over time. The unit 40 can thus modulate or stop the suction when the pressure detected by said sensor reaches a threshold pressure value and / or when the pressure increases / decreases above or below a threshold pressure acceleration / deceleration slope. In one embodiment, the threshold pressure value is between 6 mmHg and 15 mmHg, preferably equal to or substantially equal to 12 mmHg.We avoid any risk of collapse at the level of the vena cava, thanks to the threshold values ​​(absolute and / or acceleration slope) and the continuous and rapid detection (high frequency) of the pressure in the intake portion 19.

[0075] Unit 40 is equipped with computer systems that have programmed two types of alerts related to the pressure detected in the inlet port 19. First, there is a pressure threshold value which, if reached, triggers a modification of the blood pumping, typically by decreasing the pumping rate. Second, the alert consists of pressure acceleration / deceleration between two pressure measurements over time: again, if this threshold slope (acceleration / deceleration) is reached, the blood suction is modified, typically by decreasing or stopping it. Regarding the threshold slope, the predictive program for pressure evolution uses the Reynolds number, also known as the Reynolds wave. The Reynolds number is a dimensionless number used in fluid mechanics. This quantity characterizes a flow, in particular the nature of its regime.It is therefore possible to determine whether a flow is laminar, transient, or turbulent.

[0076] With each heartbeat, the volume of fluid aspirated is approximately 50 ml (milliliters), but this can reach 100 ml, depending on the patient. A natural cardiac output is between 2.5 and 4.5 l / min / m² (liters per minute per square meter). In other words, the larger the surface area of ​​a human body (and therefore its weight), the greater the blood circulation. In practice, to adjust the weight or volume of fluid to be aspirated with each heartbeat, the physician considers the patient's weight and deduces the volume at each heartbeat.

[0077] For a heart beating at 70 bpm (beats per minute), the aspiration time is approximately 0.56 seconds, so a normal flow rate is about 90 ml / s (milliliters per second per systole), or 5 L / min (liters per minute). However, in practice, physicians aim for an average flow rate of 3 or 4 L / min. Thanks to the device according to the invention, the amount of blood aspirated can be increased (or possibly decreased) to more closely match the actual functioning of a heart, without risking collapse of the vena cava.

[0078] For example, the pressure threshold value (first alert) and the pressure slope or acceleration threshold value (second alert): If the measured pressure slope or acceleration is 50% lower than expected during a suction cycle, this indicates an imminent risk of collapse, and Unit 40 immediately reduces the suction flow rate. If the absolute pressure drops below the threshold pressure of 12 mmHg, Unit 40 immediately stops pumping for that suction cycle because the risk of collapse is very high. Pumping restarts in the next suction cycle.

[0079] In one embodiment, during an artificial cardiac cycle, the volume aspirated during artificial diastole corresponds to the volume ejected during artificial systole. However, depending on the patient's condition (e.g., hypertension, venous system status, arterial obstruction, blood viscosity, etc.), the volume aspirated during artificial diastole may differ from the volume ejected during artificial systole. For example, unit 40 can control the displacement of organ 70 so that during artificial diastole, a sufficient volume of fluid is aspirated to ensure good circulation in the patient's body. This aspirate can be slow (e.g., 70% of the cardiac cycle) and with low dynamics to avoid the risk of collapse. For a patient with hypertension, it may be advantageous not to reinject all of this aspirated volume to avoid contributing to increased blood pressure in the patient's body.For example, for a patient 12 suffering from arterial obstruction, unit 40 can be configured to perform two artificial diastoles on two successive cardiac cycles (without artificial systole), to command the artificial systole on the following cycle, with an optimal flow / pressure profile so that the systolic wave propagates deep into the body of patient 12 despite the arterial obstruction.

[0080] More generally, the configuration of unit 40 allows for the dissociation of artificial diastoles and artificial systoles: the aspiration of the fluid can be independent of the ejection regime in the aorta, which corresponds to the reality of the patient's body 12. The chamber 11 combined with the controlled displacement of organ 70, ensures a diastolic balance assignment (AED) function: the average volume of chamber 11 allows for the regulation of the fluid volume in the patient's body 12, and therefore the average blood pressure in said body.

[0081] The best results in achieving optimal synchronization between artificial circulation and the patient's body's needs 12 are obtained in particular by mastering the following two elements: Rapid processing of measurements (preferably with a fast acquisition frequency of 200 Hz or of this order of frequency) to send commands in real time to actuator 10. High-velocity execution of commands to adjust the actual fluid movement to the commands.

[0082] The arrangement of the various elements and / or means and / or steps of the invention, in the embodiments described above, should not be understood as requiring such an arrangement in all implementations. In any event, it will be understood that various modifications may be made to these elements and / or means and / or steps without departing from the spirit and scope of the invention.

[0083] Furthermore, one or more features described only in one embodiment can be combined with one or more other features described only in another embodiment. Similarly, one or more features described only in one embodiment can be generalized to other embodiments, even if that feature or those features are described only in combination with other features.

[0084] In any event, in claims, any reference sign in parentheses cannot be interpreted as a limitation of the claim.

[0085] Another aspect not covered by the invention and claims concerns a device for temporary circulatory assistance or replacement of a patient's heart, which is not extracorporeal but implantable and comprises: a linear actuator configured to move an organ in translation within a chamber so as to cause a pulsating fluid flow capable of supporting the activity of the patient's heart, which flow is characterized by a succession of aspiration phases and fluid ejection phases, a control unit configured to pilot the actuator and control the movement of the organ, which piloting is carried out by taking into account input data from measurements of the patient's hemodynamic parameters, during each ejection phase the unit moving the organ a distance D.

[0086] The control unit is configured to: Sequence the displacement D into an integer number N of cycles Ci, with N > 1, each cycle having the same duration Δt. On each cycle Ci, calculate a displacement length Li of the organ such that D = ∑ i = 1 N Li , which length L i is calculated based on input data from: ∘ one or more hemodynamic parameters of the patient measured during cycle C i , or during one or more cycles prior to said cycle C i and close to said cycle C i , ∘ the value of a setpoint pressure Π i based on the Franck-Starling law, on each cycle C i , transmit an instruction to the actuator to move the organ by the calculated length L i.

[0087] The other features claimed and described above apply to this embodiment not covered by the invention.

Claims

1. Device for the temporary extracorporeal circulatory assistance or replacement of the heart (13) of a patient (12), comprising: - a linear actuator (10) configured to displace a member (70) in translation in a chamber (11) so as to cause a pulsatile fluid flow capable of supporting the activity of the heart of said patient, which flow is characterised by a succession of phases of aspiration and phases of ejection of the fluid, - a control unit (40) configured to drive the actuator (10) and control the displacement of the member (70), which driving is carried out while taking into account input data originating from measurements of haemodynamic parameters of the patient (12), said unit displacing the member (70) over a distance D during each ejection phase, characterised in that the control unit (40) is configured: - to sequence the displacement through the distance D in an integer number N of cycles Ci with N>1, each cycle having the same duration Δt, - in each cycle Ci, to calculate a length Li of the displacement of the member (70) such that D = ∑ i = 1 N Li , which length Li is calculated as a function of the input data derived from: ∘ one or more haemodynamic parameter(s) of the patient (12), which are measured during the cycle Ci or during one or more cycle(s) that precede(s) said cycle Ci and are close to said cycle Ci, ∘ the value of a setpoint pressure Πi based on the Frank-Starling law, - in each cycle Ci, to transmit an instruction to the actuator (10) in order to displace the member (70) through the calculated length Li.

2. Device according to claim 1, wherein the control unit (40) is configured to calculate the displacement length Li according to the following formula: L i = L i − 1 + Π i − PR i − Δ P cor × Δ t 2 S × μ where - Li-1 is the displacement length in the cycle Ci-1; - Πi is the setpoint pressure in the cycle Ci; - PRi is the actual pressure of the fluid as measured in the chamber (11) in the cycle Ci; - ΔPcor is a corrective pressure value, - S is the equivalent cross section of the surface of the member (70) in contact with the fluid, - µ is a constant.

3. Device according to claim 1, further comprising an ejection cannula (16) which can be implanted in the human body and is fluidically connected to the chamber (11), and wherein the control unit (40) is configured to calculate the displacement length Li according to the following formula: L i = L i − 1 + Π i − PR i × Δ t 2 S × μ where - Li-1 is the displacement length in the cycle Ci; - Πi is the setpoint pressure in the cycle Ci; - PRi is the actual pressure of the fluid as measured at the ejection cannula (16); - S is the equivalent cross section of the surface of the member (70) in contact with the fluid, - µ is a constant.

4. Device according to one of the preceding claims, wherein the control unit (40) is configured to generate a setpoint pressure curve according to the Frank-Starling law for each ejection phase, the value of Πi corresponding to the value of the pressure on said curve in the cycle Ci.

5. Device according to claim 4, wherein the setpoint pressure curve according to the Frank-Starling law is identical for each ejection phase.

6. Device according to claim 4, wherein the setpoint pressure curve according to the Frank-Starling law is redefined for each ejection phase or redefined at regular time intervals as a function of arterial pressure values measured with a preconfigured frequency, at one extremity of the body of the patient.

7. Device according to one of claims 1 to 3, wherein the control unit (40) is configured to calculate the setpoint pressure Πi for each cycle Ci according to the following formula: Π i = PM s − K s ⋅ VE i − 1 − R s ⋅ Q i − 1 where: - PMS is the maximum systolic pressure whose value is preconfigured and / or adjusted; - KS is a coefficient between 0.9 and 1.1; - VEi-1 is the volume of fluid set in movement in the chamber (11) during the ejection phase in the preceding cycle Ci-1; - RS is a coefficient between 0.4 and 0.6; - Qi-1 is the fluid flow rate ejected from the chamber (11) in the preceding cycle Ci-1.

8. Device according to one of the preceding claims, wherein the number N of cycles Ci is between 45 and 55 and the duration Δt of each cycle Ci is between 0.004 s and 0.006 s.

9. Device according to one of the preceding claims, wherein the control unit (40) is configured to dissociate the aspiration phases and the ejection phases of the fluid in such a way that, during an artificial cardiac cycle, the volume aspirated during the aspiration phase is different from the volume ejected during ejection phases.

10. Device according to one of the preceding claims, wherein: - the first cycle C1 coincides with the natural systole and begins at the instant when a QRS complex of an electrocardiography ECG signal measured on the patient (12) is detected, or - the first cycle C1 is offset with respect to the natural systole and begins after an offset of between 0.2 and 0.3 seconds from the instant when a QRS complex of an electrocardiography ECG signal measured on the patient (12) is detected.