A mechanical cardiac assistance device that works in synergy with the heart.

Abstract

The invention relates to a device (10) for implantable mechanical cardiac assistance in a heart (1) and comprising: - a hydraulic module (11) with an outlet (16) having an internal surface area of ​​less than 113.1mm², - a controller (13) configured to automatically and dynamically control the blood flow velocity according to a velocity profile synchronized with the heart rate, and to control the blood flow velocity to a value greater than or equal to vmax for an ejection time Tpulse;vmax being greater than or equal to 0.6 m / s and the ejection time Tpulse being less than or equal to the duration Tsystole of a systole of the cardiac cycle, - a first measuring device to measure the evolution of the ventricle volume during a cardiac cycle, the ejection time Tpulse and the maximum velocity vmax being a function of this evolution of the ventricle volume, and - a second measuring device to measure the viscosity of the blood, the ejection time Tpulse and the maximum velocity vmax being a function of this blood viscosity. Figure for the abstract: Fig. 1;

Description

Title of the invention: Cardiac mechanical assistance device operating in synergy with the heart. technical field

[0001] The present invention relates to a mechanical cardiac assist device implantable in a heart and a method for controlling the speed of blood flow. Prior art

[0002] Since the 1980s, many companies have developed mechanical cardiac support devices to treat patients with severe heart failure for whom pacemaker therapy is ineffective. Indeed, when the patient's problem is mechanical, and not electrical, pacemakers are ineffective, and it is necessary to act mechanically on the patient's circulatory system to restore sufficient aortic flow. These are known as MCS (Mechanical Circulatory Support) devices. The original idea came from cardiac surgeons, who were accustomed to using cardiopulmonary bypass (CPB) during heart transplants or intracavitary surgical procedures. The initial concept was to implant a device capable of restoring aortic flow using a pump mechanism placed in bypass of the heart, such as a bypass or a second external circuit.This pump is generally aspiratory and draws blood from the left ventricle (for example, at the apex) and reinjects it into the aorta, ensuring an average flow rate of around 5 L / min and sufficient pressure (approximately 80-120 mmHg). This is the basic principle of what is still called an LVAD-Bypass (Left Ventricular Assist Device), placed in a bypass of the heart.

[0003] This is generally a continuous flow pump, or at best a pulsatile pump but not synchronous with native cardiac activity.

[0004] It has been observed, however, that when an LVAD-Bypass pump is used to restore cardiac output to a defective natural pump (the heart itself), there is necessarily a "competition" between the two pumps; there is a master and a slave. This situation is amplified when the LVAD-Bypass pump is synchronous, the primary reason why pulsed LVAD-Bypass pumps are asynchronous. In this situation, the LVAD-Bypass must be the master in order to control aortic flow, which leads to the native heart becoming "slaved and dependent" on the device. Furthermore, one of the consequences of the operating principle of LVAD-Bypass pumps is that the pump must perform work to ensure both the flow and the pressure in its entirety in order to compensate for the defective natural pump (the heart). The electrical consumption of the LVAD-Bypass pump must therefore be very high, on the order of 8 to 15 W. (20W for some models due to frequent restarts in this type of suction-propeller device). This electrical consumption includes the pump's efficiency, which is around 5 to 15%, given that the hydraulic power ultimately transmitted to the left cardiovascular system is around 1.3W.

[0005] Ultimately, only 5% of patients with severe heart failure are fitted with LVAD-Bypass. This is due to numerous other side effects which explain why the benefit / risk ratio of LVAD-Bypass is largely unfavorable to widespread deployment and leaves 95% of patients with severe heart failure in a therapeutic dead end.

[0006] Documents FR2955499, FR3040304, and FR3130624 are known to describe an intraventricular cardiac pump, called ICOMS, implanted in a ventricle of the heart. The blood flow from the ICOMS pump is ejected inside the ventricle, unlike LVAD-Bypass devices in which the forced blood flow is ejected into the aorta, outside the heart, downstream of the aortic valve. An intraventricular pump such as ICOMS complements the heart's natural pumping action. To assist the failing natural pump (the heart), the principle of ICOMS is not to use another pressure source placed in parallel downstream of the aorta (high-pressure outlet), but rather another flow source placed in parallel upstream of the aorta.Under these conditions, the two flow rates combine so that the natural pump continues to contribute to the work of ejecting blood towards the aorta and is assisted by an artificial flow pump. This has the benefit of retraining the heart muscle by relieving its workload towards a common outflow pathway, namely the aortic valve. This concept contributes to the rehabilitation of a defective heart muscle. This is precisely what the ICOMS intraventricular pump does by propelling and channeling blood through the aortic valve, acting directly within the left ventricle, synchronously with the heart. The ICOMS intraventricular pump contributes only a portion of the ejection work, which inherently reduces its energy consumption. The order of magnitude of the ICOMS intraventricular pump's contribution to flow is 33% nominally (between 20% and 50%) of the total flow.Thus, its workload (and therefore its power consumption) is reduced by a factor of 3 in nominal terms (between 2 and 8) compared to that of an LVAD-Bypass. This reduction in power consumption allows for a relaxation of certain power supply constraints by a transcutaneous energy transfer (TET) system without risking hyperthermia of the tissues (internal and skin), and thus eliminates the need for a percutaneous electrical power supply (driveline). In principle, since the ICOMS intraventricular cardiac pump does not replace the work of the heart, the constraints in terms of pump and TET reliability are also much less severe than for an LVAD-Bypass. Finally, the reduced power consumption allows the intraventricular cardiac pump... ICOMS to function longer by being powered by an implantable battery, which significantly improves the quality of life of patients.

[0007] However, electricity consumption is considered to remain significant and a key point for improvement.

[0008] The present invention addresses the following functional issues: - Lasting restoration of the heart's contractile and compliance functions, - Control of the restoration of the patient's cardiac output, via improvement of their stroke volume, - Minimizing energy consumption for the use of a TET and increasing the autonomy of an implantable battery, - Reduction in device size to facilitate surgery and reach a maximum number of patients.

[0009] The present invention also addresses the following dysfunctional issues: - Minimizing the risk of infection (related to surgery and the driveline), - Minimizing the risk of pump-induced hemolysis, - Minimizing the risk of pump-induced thrombosis, - Minimizing the risk of pulmonary edema related to right ventricular degradation induced by the physiological incidence of the pump (suction effect on the septal wall), - Minimizing the risk of gastrointestinal bleeding induced by the physiological incidence of the pump, in particular non-synchronous pulsatility and the anticoagulant level. Description of the invention

[0010] At least one of the objectives is achieved with a mechanical cardiac assistance device implantable in a heart and comprising a hydraulic module intended to be disposed entirely inside the left ventricle of the heart, this hydraulic module comprising an inlet for blood flow and an outlet of blood flow into the ventricle and to the aortic valve of the heart, a base intended to be fixed on the ventricle and a controller capable of controlling the hydraulic module.

[0011] According to the invention: - the outlet of the hydraulic module has an internal surface area of ​​less than 113.1 mm², corresponding for example to an equivalent internal diameter of less than 12 millimeters, - The controller is configured to automatically and dynamically control the blood flow velocity according to a velocity profile synchronized with the pulse of the heart; and to control the blood flow velocity to a value greater than or equal to vmax for an ejection time Tpulse, vmax being greater than or equal to 0.6 m / s and the ejection time Tpulse being less than or equal to a systole time Tsystole of a cardiac cycle, - a first measurement device to measure the evolution of ventricular volume during a cardiac cycle, with the ejection time Tpulse and the maximum velocity vmax being functions of this evolution of ventricular volume, and - a second measuring device to measure blood viscosity, the ejection time Tpulse and the maximum speed vmax being a function of this blood viscosity.

[0012] The mechanical cardiac assistance device according to the invention is a heart pump with a hydraulic module constituting the part implanted in the ventricle. Such a heart pump makes it possible to improve the functioning of the ventricle, without disrupting the rest of the cardiac system, with a view to restoring healthy heart activity and improving the patient's quality of life.

[0013]

[0014] To achieve this, a synergistic operation is planned between the cardiac mechanical assistance device according to the invention and the heart.

[0015] Synergy is understood to mean an overall augmented effect of two systems cooperating together, this overall augmented effect being greater than the sum of the effects produced respectively by the two systems in an independent or isolated operation.

[0016]

[0017] With such a device, the expected synergistic result is that the instantaneous aortic flow (defined at the aortic valve) when the hydraulic module is activated is strictly greater than the sum of the instantaneous aortic flow when the hydraulic module is inactive and the instantaneous flow supplied by said hydraulic module. The features of the invention make it possible to achieve such a synergistic effect.

[0018]

[0019] The hydraulic module is capable of converting the mechanical energy of its rotor into hydraulic energy of a fluid (here, blood) in a fluid column. A fluid column is defined as a coaxial cylinder of the module with a diameter corresponding to the module's outlet diameter. This hydraulic energy is essentially hydrodynamic, meaning that there is practically no increase in the static pressure of the fluid, only an increase in its velocity, and thus generates a high-velocity fluid flow in a small-diameter fluid column.

[0020]

[0021] The cardiac mechanical assist device according to the invention is designed as a flow source and not as a pressure source. Its operation relies on synergistic effects that increase the efficiency of the natural pump (the heart) and do not replace it. Its action is primarily during systole and contributes to increasing aortic flow. Under no circumstances does the cardiac mechanical assist device according to the invention contribute to generating the static pressure necessary for opening the aortic valve. On the other hand, the increase in aortic flow helps restore static pressure in the arterial system, which transforms the kinetic energy of the blood into arterial (static) blood pressure.

[0022]

[0023] Reducing the internal diameter of the hydraulic module outlet and increasing the maximum exit velocity of the blood flow goes against the accepted assumptions of those skilled in the art. Indeed, it is generally understood in the prior art (ICOMS) that an increase in aortic flow is achieved by adding an artificial flow to the native flow. This artificial flow is proportional to the fluid velocity multiplied by the area of ​​the module outlet. However, increasing the velocity for a given outlet cross-section increases the effects of blood hemolysis. Thus, conventional design principles aimed at minimizing blood velocity seek to increase the outlet diameter of the hydraulic module. The present invention, however, takes the opposite approach by proposing to reduce the flow rate at the hydraulic module outlet while simultaneously increasing the aortic flow.

[0024] One of the consequences is that a small outlet diameter makes it possible to considerably reduce energy consumption.

[0025]

[0026] A reduced cross-section and high-velocity outlet flow according to the invention makes it possible to:

[0027] - reduce macroscopic turbulence (vortex in the flushing chamber which is the volume in front of the aortic valve, which can create a "plug" by literally pushing the fluid towards the aortic valve. The "plug" is greatly reduced, or even eliminated, at the start of ejection, which is properly initiated, and this results in a significant increase in flow throughout the duration of systole, almost independently of the duration of forced ejection.

[0028] - significantly reduce the hydraulic resistance of the flushing chamber by Amplification effect. The fluid velocity cannot experience discontinuities (Bernoulli's principle), but is zero at the walls. Thus, an increase in velocity in the blood column exiting the pump, even over a small diameter, leads to an increase in velocity across the entire cross-section of the pumping chamber. This flow amplification effect explains the synergy; this effect is strongly non- This effect is linear and even more pronounced when blood viscosity increases. The synergistic effect thus increases the cardiac support provided by the device, while reducing the need for blood thinning through excessive use of anticoagulants.

[0029]

[0030] The mechanical cardiac assistance device according to the invention is thus physiological in that it can both accelerate blood flow through the aortic valve upstream while respecting the rhythm imposed by the heart and not generate negative pressure in the right ventricle at the level of the septum since it generates virtually no static pressure difference. By adjusting the rotational speed of the propeller, the increase in additional flow can be precisely controlled and adapted to the severity of the patient's condition, without reducing the heart's natural contribution.

[0031]

[0032] According to an advantageous feature of the invention, the outlet of the hydraulic module has an internal surface area of ​​less than 78.5mm2, 50.24mm2, 28.26mm2 or 19.62mm2, corresponding to an equivalent internal diameter of the outlet of the hydraulic module of less than 10mm, 8mm, 6mm or 5mm. Prior art systems have an internal diameter greater than 12 mm.

[0033] The internal diameter of the hydraulic module outlet can be reduced without reducing its external diameter. The advantage of reducing the internal surface area of ​​the hydraulic module outlet is to significantly increase the velocity of the outlet flow and concentrate it on a shorter fluid column to benefit from the Bernoulli amplification effect. For this purpose, an outlet diameter of less than 12 mm, or even 10 mm, 8 mm, or 6 mm, is used.

[0034] The maximum velocity vmax of the blood flow output of the hydraulic module can be greater than or equal to 0.8m / s, Im / s, l,5m / s, or 2m / s.

[0035]

[0036] According to an advantageous feature of the invention, the hydraulic module may include a helical rotor whose diameter-to-length ratio is greater than or equal to 0.54. This original feature aims to increase the hydrodynamic pressure at the expense of the static pressure required by conventional pumps.

[0037]

[0038] According to an advantageous feature of the invention, the base may include a reversible electromechanical converter capable of converting both electrical energy into mechanical energy according to motor operation, and mechanical energy into electrical energy according to generator operation.

[0039] The reversible electromechanical converter allows electrical power to be converted into mechanical power in order to rotate the rotor at speed Q. Preferably, this reversible electromechanical converter contains a motor connected to the rotor of the hydraulic module and an electrical energy converter that controls the speed Q based on a setpoint, thus controlling the rotational speed profile Q(t). Preferably and uniquely, the synergistic effect allows for the recovery of hydraulic energy; that is, the hydraulic module can operate, at least transiently, as a turbine instead of a pump. To recover this energy, the motor is preferably a reversible motor (capable of operating as a generator), and the electrical energy converter is itself reversible in order to supply power to the grid, such as, for example, a reversible inverter.

[0040] With reversible operation, the hydraulic module can feed energy back into an implantable energy storage device, such as a battery. This transfer can occur transiently, particularly during the ejection phase when assistance is no longer required. At a minimum, this involves recovering the kinetic energy of the rotor during its deceleration phase. Furthermore, during a transient phase, the hydraulic module can operate as a turbine. The electromechanical converter then functions as a generator, and the entire system is capable of recovering energy supplied by the cardiovascular system. The advantage of such a configuration is clear, as it significantly improves the autonomy of the mechanical cardiac assistance device.

[0041]

[0042] According to an advantageous feature of the invention, the hydraulic module comprises a box having a straight or flared external profile going towards the outlet.

[0043] This chamber is a stator, and its external shape plays a role in the dynamic behavior of the mechanical cardiac assistance device. Specifically, the external shape of the chamber contributes to generating, or even amplifying, a Venturi effect that occurs at the outlet of the hydraulic module during systole. In particular, this shape must not be convergent, and the head of the hydraulic module must not be narrower than its body, as this would counteract the Venturi effect.

[0044] Furthermore, the box may have an external surface traversed by a helical groove. Its internal surface may also be traversed by a helical groove.

[0045]

[0046] The mere presence of the hydraulic module in the ventricle has an impact on the flow regime in that ventricle during systole. It is known that hydraulic resistance increases sharply when the flow is turbulent. Here, we only consider macroscopic turbulence, on a scale of at least millimeters and on lasting on the order of milliseconds, these turbulences contribute both to increased blood heat loss and, more importantly, can create a "plug" effect. A macroscopic vortex is characterized by a volume of blood rotating on its axis and behaves literally like an obstructive object in the outflow tract. The presence of macroscopic turbulence is directly related to the Reynolds number, a dimensionless quantity well known to those skilled in hydrodynamics. To reduce the Reynolds number, and thus reduce macroscopic turbulence, the ventricular volume must be reduced. The perverse effect and vicious cycle in which most patients with severe heart failure find themselves is clear: their ventricular volume increases, which reduces contractility and ventricular compliance, increases internal hydraulic resistance, and consequently reduces the ejection fraction.To compensate in the short term, the muscle exerts more effort and the heart dilates further, which further degrades its performance.

[0047]

[0048] The helical shape of the stator's external surface allows for a passive reduction of the ventricle's internal hydraulic resistance. Indeed, at the stator surface, the flow is imposed by the boundary layers that follow it. Thus, when the external stator has a shape of at least one blade, the boundary layer induces a helical flow of blood passing outside the hydraulic module. As a result, a certain volume of the blood column follows this helical flow. Ideally, this helical flow conforms to a helical flow of the additional flow ejected by the hydraulic module at high speed, for example, initiated by a helical shape of the stator's internal surface. This has the effect of further reducing the heart's internal hydraulic resistance, but this time actively.This can also contribute to creating a vacuum at the outlet of the hydraulic module, which causes a suction effect similar to the Venturi effect mentioned earlier.

[0049]

[0050] According to an advantageous feature of the invention, the hydraulic module can be of axial type and comprise a helical rotor of length less than or equal to 25 millimeters.

[0051]

[0052] A reduced rotor diameter dimension allows for the design of a hydraulic module with a reduced volume, which helps to facilitate the implantation of the device in smaller hearts and reduce the internal hydraulic resistance of the ventricle compared to larger prior art systems.

[0053] The mechanical cardiac assistance device according to the invention can be optimized compared to a conventional pump due to the low requirement in terms of static pressure difference between the inlet and outlet of the hydraulic module. This maximum static pressure can be 30 mmHg, or even 20 mmHg, or even 10 mmHg, which is much lower than all known solutions in the prior art, which range between 50 mmHg and 120 mmHg (LVAD-Bypass and previous ICOMS).

[0054] One solution, not limiting the present invention, consists of using a propeller-type rotor characterized by the shape of its blades, in particular the diameter of the blades can be very large compared to the diameter of the hub. Preferably, the ratio between the diameter of the blades, including the hub, and the diameter of the hub is greater than 2.

[0055] Those skilled in the art know that to optimize the efficiency of an axial pump, it must be constructed with several "stages," which consequently increases the length of its rotor. Generally, a first stage is introduced to create suction at the inlet, followed by other stages to increase the static pressure as one approaches the outlet.

[0056] The present invention goes against this principle and proposes a helical rotor of reduced size, less than 25mm, or even 20mm, or even 10mm.

[0057] This aspect of the design is made possible by the synergy effect with the heart, the latter being able to be used as a "feed pump" for the hydraulic module when the positioning in the ventricle is well adjusted.

[0058] Moreover, since it is not sought to create static pressure, the use of a smaller rotor is not a disadvantage.

[0059]

[0060] Furthermore, with a reduced-size hydraulic module, the cardiac mechanical assistance device according to the invention can be implanted in a majority of patients.

[0061]

[0062] The first measuring device may include at least one cardioimpedance measuring probe for measuring the volume of the left ventricle.

[0063] The controller and probe make it possible to determine the shape of the ventricle and, in particular, the evolution of its volume during a cardiac cycle. This measurement can be carried out using means external to the device, such as an external echocardiography system to measure the evolution of the ventricle's volume, or automatically with an embedded system as described in document WO2024 / 052032.

[0064] Measuring the left ventricle volume allows both verification of the device's proper functioning and optimization of its performance. In particular, it allows adjustment of parameters such as the module's flow rates. hydraulics, ejection time, positions of the hydraulic module relative to the flushing chamber, etc.

[0065]

[0066] The mechanical cardiac assistance device according to the invention may include a fastening device in which the hydraulic module is inserted in such a way:

[0067] - adjustable for an inclination of the axis of rotation of the hydraulic module by relative to the axis of the flushing chamber, and

[0068] - removable for removal or movement along the axis of rotation of the module hydraulic.

[0069]

[0070] The fluidic interface between the hydraulic module and the core characterizes the interaction between the core and the hydraulic module and is influenced by the external shape of the module's stator as well as by the position of the hydraulic module relative to the flushing chamber.

[0071] The heart behaves with respect to the hydraulic module like an external pump that generates fluid displacement in the left ventricle. This displacement results in the conversion of static energy stored by the blood during isovolumic contraction (carried by static pressure) into kinetic energy of the blood (carried by dynamic pressure). By the principle of conservation of energy, the local increase in dynamic pressure is accompanied by a local decrease in static pressure; this is the Venturi effect. This phenomenon produces a static pressure field in the heart that is not uniform. A difference in static pressure appears between the outlet and inlet of the hydraulic module, which directly impacts its operating point depending on the module's position in the left ventricle.

[0072] The position of the hydraulic module within the ventricle and the shape of the stator are optimized to reduce the differential pressure, or even make it negative, at least transiently. Thus, the hydraulic module benefits from a natural suction effect due to the synergistic interaction between the heart and the hydraulic module. When the differential pressure is negative, this means that the external or native pump (the heart) is supplying energy to the hydraulic module.

[0073] It is therefore advantageous for the present invention to place the outlet of the hydraulic module in the negative pressure zone of the flushing chamber (or just below it) in order to benefit both from reduced consumption and, more importantly, from a feeding effect from the heart. The reference point used in terms of distance is the aortic valve. Indeed, it is known that the efficiency of a flow pump depends on its ability to generate a negative pressure at the outlet in order to "feed" the pump, i.e., to draw in the fluid. Without this feeding, the pump does not function. This aspect is This is particularly important when the pump has to withstand significant loads, that is, when it must generate a pressure at least greater than that of its load to produce a flow rate. Without this priming, the pump does not function at all, or only with very poor efficiency. In the present invention, we advantageously utilize the Venturi effect of the priming chamber to prime the hydraulic module and thus reduce its length. To achieve this, the outlet can be positioned near the priming chamber, but not too close to the aortic valve, as this could cause obstruction.

[0074] Moreover, the benefit of this Venturi effect, as well as the efficiency of the device, is all the better when the axis of the hydraulic module is aligned with the axis of the flushing chamber.

[0075] For these two reasons, the device thus includes an adjustable fixing system for the hydraulic module, capable of adjusting both the distance and the angle of the latter relative to the flushing chamber.

[0076] The optimal distance between the hydraulic module and the aortic valve, typically between 20 mm and 40 mm, depends on the shape of the left ventricle and the characteristics of each patient and can be adjusted during implantation by modifying these two parameters. In some cases of restoration of satisfactory native cardiac activity, a significant reduction in cardiac hypertrophy is observed. This rehabilitation can lead to misalignment or a shift in the hydraulic module's position relative to the aortic valve, thus reducing the device's effectiveness. Preferably, the fixation system can be automatically controlled by a remotely operated electromechanical actuator. However, it can also be adjusted manually, particularly during implantation. In some cases of permanent restoration of cardiac activity, the hydraulic module can also be removed.

[0077]

[0078] The fastening system can also be used to modify the positioning of the hydraulic module according to a measurement of the left ventricle volume.

[0079]

[0080] The device according to the invention may include a third measuring probe to measure the electrical activity of the heart and thus synchronize the speed of the hydraulic module with the cardiac activity.

[0081] With the device according to the invention, the speed profile is synchronized with the native cardiac activity, for example, by means of this probe for measuring the heart's electrical activity. Conventionally, the flow rate of a pump is almost proportional to the rotor's rotational speed, which allows for precise control of the hydrodynamic pressure through the measurement and regulation of the motor's speed Q. Consumption The energy consumption of the device increases with the rotational speed Q, following a classical law based on a third-order polynomial. Therefore, for energy efficiency reasons, it is preferable to minimize the duration and intensity of the rotational speed. Furthermore, there is a risk of hindering proper left ventricular filling during diastolic flow if the rotational speed is too high during this phase. This is why the rotational speed is designed to be regulated synchronously with cardiac activity.

[0082] Advantageously, in particular for certain pathologies which lead to hypertrophy of the heart (the most frequent), the electrical consumption of the device is substantially reduced, without degrading the hemodynamic efficiency of the device by limiting the duration of hydrodynamic assistance to a fraction of the ejection time, i.e. of the duration of opening of the aortic valve.

[0083] Advantageously, the assistance is maximal at the moment the valve opens and for an ejection time of less than 300 ms, or even 200 ms, or even 100 ms. This assistance time is one of the parameters of the electromechanical controller's algorithm and can be modified manually via a device programmer or automatically, in particular based on the measurement of ventricular volume and / or heart rate.

[0084] The first measuring device can be configured to measure a final systolic volume (VolVGmin) and a final diastolic volume (VolVGMax) of the ventricle.

[0085] The second measuring device may include at least one fluid viscosity measuring probe.

[0086] The second measuring device may include an external viscometer system for measuring blood viscosity.

[0087] According to the invention, the hydraulic module may include a stator in which is placed at least one first magnet intended to cooperate magnetically with at least one second magnet placed in the helical rotor of the hydraulic module so as to maintain the rotor in the axis of rotation.

[0088] The controller is capable of controlling the intensity of the dynamic pressure at the output of the hydraulic module according to a time profile determined by a configurable algorithm. This controller is associated with a motor, a converter capable of controlling the speed of this motor, a speed sensor, and a torque sensor. The coupling between the motor and the hydraulic module can thus be achieved with a magnetic coupler composed of at least two magnets, one of which is fixed to the motor and the other to the rotor. This magnetic coupler provides a hermetic seal between the hydraulic module and the motor, thus preventing the ingress of fluid that could damage the motor. This magnetic coupler can advantageously be used to guide rotation and translation of the rotor without any contact, thus minimizing friction and the risk of wear.

[0089]

[0090] In order to avoid introducing a suction effect on the septal wall located near the blood flow inlet to the hydraulic module, the effective inlet cross-section is preferably as large as possible. This can advantageously be achieved with a stator of the type perforated by vents, known to those skilled in the art.

[0091] Increasing the effective inlet cross-section reduces the flow velocity at this point, thus preventing the generation of a pressure drop. Furthermore, the low blood velocity in the hydraulic module reduces the risk of hemolysis. Therefore, as a general rule, the rotational speed of the rotor Q is minimized. To achieve this, the surface area of ​​the rotor blades is advantageously increased, while minimizing shear stresses.

[0092]

[0093] According to an advantageous feature of the invention, the controller can be configured to control a dynamic pressure at the output of the hydraulic module according to a predetermined time profile, the dynamic pressure being a function predetermined by the design of the hydraulic module, this function depending on the rotational speed Q of the rotor and the viscosity of the blood.

[0094]

[0095] According to an advantageous feature of the invention, the controller can be configured to activate the hydraulic module only every "n" cardiac cycles, "n" being a function of a diastolic volume of the ventricle.

[0096] The implementation of such a process can lead to remodeling, that is to say a process by which the diseased myocardium returns to a healthier state and recovers its compliance and / or contractility properties.

[0097] For long-term implantation, it can be observed that over time, the heart's contractility and compliance recover as the heart returns to a less dilated state. During the course of therapy, the cardiac mechanical support method according to the invention can preferentially reduce the level of artificial support as the native heart recovers its capacity. For example, therapeutic algorithms can be developed to support recovery by providing support only at every "n" cycles determined by a clinician, or by reducing the Qmax and Tpulse parameters.

[0098] According to another aspect of the invention, a method is proposed for automatically and dynamically controlling the speed of blood flow in a mechanical cardiac assist device implantable in a heart, this device comprising a hydraulic module with a hydraulic module outlet having an internal surface area of ​​less than 113.1 mm², the process comprises the following steps: - measure blood viscosity, ejection time Tpulse and maximum speed vmax being a function of this blood viscosity. - to measure the change in ventricular volume during a cardiac cycle, with the ejection time Tpulse and the maximum velocity vmax being functions of this change in ventricular volume, - to automatically and dynamically control the blood flow velocity according to a velocity profile synchronized with the heart rate, and to control the blood flow velocity to a value greater than or equal to vmax for an ejection time Tpulse; vmax being greater than or equal to 0.6m / s and determined according to the evolution of the ventricle volume and according to the viscosity of the blood, the ejection time Tpulse being less than or equal to a duration Tsystole of a systole of the cardiac cycle and determined according to the evolution of the ventricle volume and according to the viscosity of the blood. Description of the figures and methods of implementation.

[0099] Other advantages and features of the invention will become apparent upon examination of the detailed description of a non-limiting embodiment and the accompanying drawings, in which:

[0100] [Fig-1] is a schematic view of a cardiac assistance device according to the invention implanted in a heart,

[0101] [Fig.2] is a schematic view illustrating a helical groove made on a surface external to the stator of the hydraulic module of the [Fig.l],

[0102] Figure [Fig. 3] is a comparative graphical representation of the static characteristics flow rate / pressure of a cardiac assist device according to the invention compared to a pump according to the prior art,

[0103] Fig. 4 is a graphical representation illustrating the pressure difference static pressure between the outlet and inlet of the hydraulic module implanted in the left ventricle, during a cardiac cycle,

[0104] Fig. 5 is a graphical representation illustrating an evolution over time of the instantaneous flow rate and the average velocity of the forced blood column, at the outlet of the hydraulic module,

[0105] Figure 6 is a graphical representation illustrating the hydrodynamic pressure of the hydraulic module and comparison with static pressure,

[0106] Figure 7 is a graphical representation illustrating the effect of the synergy of the point from a flow rate perspective,

[0107] Figure 8 is a graphical representation illustrating the temporal evolution of hydraulic power,

[0108] Figure 9 is a graphical representation illustrating the temporal evolution of the work of the device transmitted to the arterial system via the aorta,

[0109] The [Fig. 10] is a graphical representation illustrating the temporal evolution of the hydraulic power transmitted to the cardiovascular system via the aorta.

[0110] The embodiments described below are not in any way limiting; in particular, variants of the invention may be implemented comprising only a selection of the features described below, isolated from the other features described, if this selection of features is sufficient to confer a technical advantage or to differentiate the invention from the prior art. This selection includes at least one preferably functional feature without structural details, or with only a portion of the structural details if this portion alone is sufficient to confer a technical advantage or to differentiate the invention from the prior art.

[0111] The present invention aims to improve the functioning of the left ventricle, without disturbing the rest of the cardiac system, with a view to restoring healthy heart activity and improving the patient's quality of life.

[0112] The invention is particularly concerned with the left ventricle into which a cardiac mechanical assistance device is intended to be inserted.

[0113] In [Fig. 1], a patient's heart 1 is shown, consisting of a left atrium 2, a left ventricle 3, and an aorta 4. Oxygenated blood arrives from the lungs into the left atrium via the left pulmonary vein (not shown). During diastole, blood is ejected from the left atrium 2 into the left ventricle 3 through the mitral valve 5. During systole, the myocardium 6 contracts, and blood is ejected from the left ventricle 3 into the aorta 4 through the aortic valve 7. The wall separating the left ventricle 3 from the right ventricle is called the interventricular septum 8. There is a volume enclosed by the aortic valve, the mitral valve, and the septum, which is called the outflow tract 9.

[0114]

[0115] The cardiac mechanical assistance device 10 according to the invention comprises: - a hydraulic module 11 which is the entire part intended to be placed inside the left ventricle 3, - a base 12 intended to house an electromechanical converter (a motor) 12a integral with the hydraulic module and to serve as a support for a fixing system, for example through the apical wall of the ventricle, and - a controller 13 equipped with electronic means to control the electromechanical converter 12a attached to the hydraulic module 11.

[0116] A battery (not shown) is provided inside the controller 13 to power all the electronic components as well as the electromechanical converter 12a located inside the base and / or in the hydraulic module 11.

[0117] The controller 13 is connected to the base by one or more connecting cables 14.

[0118] The cardiac mechanical assistance device 10 is fixed to the wall of the left ventricle by means of a fixation system 15 which allows axial displacement to move the outlet 16 of the hydraulic module 11 closer to or further from the aortic valve 7. The fixation system also allows orientation of the axis of the cardiac mechanical assistance device so as to place the outlet 16 opposite the aortic valve 7.

[0119]

[0120] During the installation of the mechanical cardiac assist device, the hydraulic module 11 is placed in the left ventricle so that the outlet 16 is at least 20 mm from the aortic valve. To achieve this, the hydraulic module can be moved axially before being locked in place.

[0121]

[0122] Ideally, the outlet 16 has an internal diameter of less than 12 millimeters, for example 6 mm. It is through this opening 16 that the artificial blood flow passes to the aortic valve 7 in the form of a jet, a jet of small diameter but high velocity. The controller 13 is configured to generate an outlet flow at a maximum velocity greater than or equal to Im / s.

[0123]

[0124] The hydraulic module 11 comprises a short rotor 19, for example 25 mm long, equipped with at least one helical blade. Advantageously, the ratio between the internal diameter of the outlet 16 of the hydraulic module and the length of the rotor 19 is greater than or equal to 0.54.

[0125] The casing of the hydraulic module 11 is fixed and constitutes the stator of the mechanical cardiac assistance device. Its external profile is straight in [Fig. 1] and can even be flared towards the outlet.

[0126] A probe 21 for measuring the heart's electrical activity and / or cardioimpedance is placed on the outer wall of the left ventricle 3. This probe 21 allows for the measurement of both the natural action potential that controls the heart and thus synchronizes it, and the left ventricle volume by capacitance, using the hydraulic module chamber as a second electrode. Measuring the volume allows for monitoring the patient's pathological progression and adapting the control of the mechanical cardiac assist device accordingly.

[0127]

[0128] Figure 2 shows a cross-sectional view of the cardiac mechanical assistance device according to the invention with a stator having an external surface traversed by a helical groove 20. A helical groove made on the internal surface of the stator can also be considered.

[0129]

[0130] In [Fig.2], a magnet 22 is also shown arranged inside the head of the hydraulic module 11, around the outlet 16. This magnet 22 is intended to cooperate with another magnet 23 arranged in the rotor 19, near the magnet 22. Such an arrangement constitutes a magnetic bearing which allows the rotor to be kept in its axis of rotation without contact.

[0131] The magnets and the internal or external groove can be made together or independently in the device of [Fig.1].

[0132]

[0133] The heart 1 is a natural pump whose main function is to circulate blood throughout the human body. The cardiovascular system is characterized by the following parameters illustrated in [Fig. 1]:

[0134] - the aortic pressure, denoted Pao(t), which is the relative static pressure in the aorta 4. Expressed in Pascals or millimeters of mercury (mmHg = 133 Pa). This pressure fluctuates over time, typically between 80 and 120 mmHg. The average pressure for a healthy patient is approximately PaoMoy = 100 mmHg.

[0135] - the aortic flow Dao(t) is the instantaneous flow measured across the aortic valve 7 Expressed in L / min, this flow rate is pulsatile, varying from zero when the aortic valve is closed (diastolic phase) to a maximum of approximately 15 L / min for a healthy patient when the aortic valve is open (systolic phase). In a diseased patient, the instantaneous flow rate can be considerably reduced to 1 L / min or even much less.

[0136] - the volume of the left ventricle, denoted VolVG in m3, is the volume enclosed by the The aortic and mitral valves close simultaneously. There is a lower value of left ventricular volume (LVV) at the end of systole called the end-systolic volume (LVVM). There is a higher value of LVVM at the end of diastole called the end-diastolic volume (LVVM). The stroke volume (or stroke volume) is the difference between LVVMmax and LVVMmin and corresponds to the volume of blood ejected by the heart with each cycle. It is approximately 60 mL for a healthy patient and decreases to 30 mL or even less for a diseased patient.

[0137] - Intraventricular pressure, denoted PVG(t), is the pressure of the left ventricle Measured inside the left ventricle and expressed in mmHg, this is a relative static pressure whose value varies considerably during a cardiac cycle. It oscillates between a low pressure value (PVGmin) and a high pressure value (PVGmin). PVGmax. PVGmax is essentially the same as aortic pressure when the aortic valve is open. Thus, PVGmax < Pao. PVGmin, on the other hand, is largely determined by cardiac compliance, i.e., the elastic capacity of the myocardium to naturally regain sufficient left ventricular volume (VolVL) through a "suction" effect. PVGmin is on the order of 5-10 mmHg and, in any case, lower than the static left atrial pressure (Pog). Note that the static pressure of the left ventricle is not actually uniform throughout the entire left ventricle. Indeed, there are dynamic turbulence phenomena, i.e., blood vortices, particularly during the filling phases of diastole (during mitral valve opening) and during the ejection phases of systole (during aortic valve opening). These phenomena tend to locally reduce static pressure.Thus, PVG(t) is defined as the spatial average of the relative static pressure inside the left ventricle.

[0138] To understand part of the dynamics of cardiac function in the presence of a mechanical cardiac assist device, we also consider the static pressure difference at the aortic valve, denoted PvalveAO(t), as a state variable. This static pressure difference is given by the difference between the aortic pressure Pao(t) and the left ventricular pressure PVG(t). Thus, PvalveAO(t) = Pao(t) - PVG(t). It is understood that PvalveAO(t) is necessarily positive and changes significantly during a cardiac cycle between a minimal value of approximately zero (when the aortic valve is open) and a maximum value, denoted PvalveAOmax, reached during the filling phase. The order of magnitude of PvalveAOmax = Pao - PVGmin is approximately 100 mHg, reached during the filling phase.

[0139] The patient's heart, as a natural pump, is a hydrodynamic system capable of generating a certain native aortic flow rate, called Dnat(t), expressed in L / min. This is the cardiac output at the level of the aorta before activation of any mechanical cardiac assist device. Therefore, we have the following relationship:

[0140] Dnat(t) = Dao(t) in the absence of a mechanical cardiac assist device

[0141] There are many pathologies / causes of dysfunction that lead to a reduction in the heart's natural efficiency. A decrease in native cardiac output is then observed, which is called severe heart failure. When the pathology is related to a dysfunction of the heart's electrical control, pacemakers are used to restore proper control and sufficient cardiac function. On the other hand, when the dysfunction is non-electrical in origin, the failure is mechanical in nature. Pacemakers are then ineffective, and patients find themselves in a therapeutic dead end, which the invention aims to overcome.

[0142]

[0143] Although synergy is completely defined by an effect on the hydrodynamic energy transmitted to the cardiovascular system, it is possible to decompose it into effects on pressures and flow rates. As a reminder, hydrodynamic power is the product of a flow rate and a pressure. The hydrodynamic energy over a cardiac cycle is therefore the integration of this power over the duration of the cycle.

[0144]

[0145] Synergy effect on flow rates.

[0146] Let Dnat(t) be the patient's instantaneous native aortic cardiac output expressed in L / min. This is the cardiac output before device activation. Device activation is characterized by the use of an external energy source, for example, electrical, to generate blood flow at the outlet of the mechanical cardiac assist device. Ddisp(t) is defined as the instantaneous flow measured at the outlet of the mechanical cardiac assist device. This flow Ddisp(t) corresponds to the blood flow passing through the device. The positive direction is from the inlet to the outlet. By convention, the outlet of the mechanical cardiac assist device is distinguished from the inlet by the fact that it is closer to the aortic valve.

[0147] Let Dao(t) be the instantaneous aortic cardiac output of the patient equipped with the device expressed in L / min.

[0148] The flow coefficient Sd is defined by the following relation, defined during the ejection phase (systole):

[0149]

[0150] Dao(t) = Dnat(t) + Sd * Ddisp(t)

[0151]

[0152] Dnat(t) is therefore the measure of the aortic flow Dao(t) when the device is inactive, i.e. when Ddisp=0.

[0153] The specific feature of the present invention is that the flow coefficient Sd is > 100% during the ejection phase (systole). Thus, the instantaneous aortic flow is greater than the sum of the native flow and the instantaneous flow of the device. This is notably due to a non-linear fluid entrainment effect under conditions of alignment of the forced flow with the natural flow.

[0154] On the contrary, for anterior-art LVAD-Bypass devices, a competition effect is observed between the heart and the device, characterized by a flow coefficient < 100%. This flow coefficient of LVAD-Bypass devices is generally very low (<50%) in the first weeks post-implantation (typical post-implantation LVAD-Bypass example: Dao=15L / min, Dnat=9L / min, and Ddisp=12L / min, i.e., Sd=50%). This leads in the long term to atrophy of the cardiac muscle, which becomes completely dependent on the device. Dnat then tends towards 0 and Sd tends towards 100%, meaning that the LVAD-Bypass has replaced the action of the left ventricle (typical long-term example: LVAD-Bypass Dao=15L / min, Dnat=3L / min, and Ddisp=15L / min, i.e., Sd=80%). Thus, the increase over time in the Sd coefficient of an anterior artery bypass device is not at all a synergistic effect but, on the contrary, the effect of the replacement or total takeover of blood flow, since it is common practice for clinicians to state that an aspirating cardiac assist pump of the LVAD-Bypass type is particularly effective when the aortic valve remains closed (which means that the native contraction is too weak to cause the aortic valve to open, which is entirely unphysiological by definition).Thus, the patient's ventricular muscle becomes atrophied and their heart dependent on the cardiac assist pump alone (a life-threatening situation if the cardiac assist pump fails).

[0155] Conversely, a drive effect is used to amplify the flow rate provided by the device of the present invention post-implantation. The following post-implantation values ​​are obtained: Dao = 15 L / min, Dnat = 9 L / min, and Ddisp = 3 L / min, i.e., Sd = 200%. With the present invention, an increase in Dnat is observed, a phenomenon called cardiac restoration. One of the objectives of the invention is to maintain the flow coefficient of the mechanical cardiac assist device substantially constant. Thus, restoration is accompanied by a reduction over time in the flow rate Ddisp(t) instead of a reduction in the flow coefficient. The long-term objective is: Dao = 16 L / min, Dnat = 15 L / min, and Ddisp = 0.5 L / min, i.e., Sd = 200%. When the native flow rate has reached a level sufficiently close to a healthy flow rate, it is then possible to remove the mechanical cardiac assist device.This will obviously not be possible for all heart conditions, particularly when the cause of the actual impairment is related to a myocardial infarction. Some patients will therefore need to remain permanently fitted with the device, but with a significant improvement in their quality of life.

[0156]

[0157] The following table illustrates examples of flow coefficients Sd of different cardiac mechanical assist devices. The flow rates are instantaneous quantities during the systolic phase.

[0158]

[0159] Dao Dnal MK Technology— Sd in % Comment Total artificial heart 151.unin 0 L / min Sd — -100% Total replacement of mechanical activity LVAD-Bypass or Trans-aortic (anterior artery) 15L / min 9L / min Sd < 100% Pure example Sd = 58% Daat decreases over time to tend towards 0 because the device atrophies the heart. S increases over time to tend towards 1. The device provides almost 108% of the ICOMS cardiac output (document FR2955499). Anterior artery 15L / min 9L / min Sd = 100% Dnat increases over time via a mechanism of restoration of native activity. Ddisp is gradually reduced according to this restoration. Device according to the present invention 15L / min 9L / min Sd ■ 108% Example Sd=200% Similar to ICOMS but with a synergistic effect: there is an amplification of the blood flow provided by the device

[0160] Synergistic effect on pressures

[0161] Pdisp(t) is the total pressure difference (static and dynamic) between the outlet and inlet of the hydraulic module during a cardiac cycle. It is important to distinguish the static pressure difference of the device, denoted PdispStat(t), from the dynamic pressure difference, denoted PdispDyn(t).

[0162]

[0163] Pdisp(t)=PdispStat(t) + PdispDyn(t)

[0164]

[0165] The static pressure difference represents, for the displaced volume of fluid (blood), the increase in force exerted uniformly in all directions on the external surface of said displaced volume. This static pressure is represented by a divergent force field, capable, for example, of inflating or deflating the walls of the cardiovascular system according to their elasticity (their compliance). The static pressure difference accounts for a pressure source.

[0166] The dynamic pressure difference, for a given volume of displaced fluid (blood), represents the increase in its kinetic energy, i.e., its velocity. This dynamic pressure is represented by a force field with zero divergence, capable of circulating said volume of fluid through the cardiovascular system. The dynamic pressure difference accounts for a source of flow.

[0167] Bernoulli's equations explain how static pressure can be transformed into dynamic pressure and vice versa, thanks to the principle of conservation of total hydrodynamic energy. Thus, a source of static pressure can be transformed into a source of flow and vice versa. This is achieved simply by opening or closing a circuit, which is what valves do.

[0168] In practice, the cardiovascular system as a whole is complex because it alternates between a source of pressure and a source of flow. The transition from one to the other occurs when a valve opens or closes.

[0169] During the isovolumetric phases of compression or decompression, the left ventricle behaves as a source of static pressure, with all valves closed. This static pressure is positive due to the active contraction of the myocardium during systole, and negative due to the passive suction effect caused by myocardial compliance during diastole. It is the difference in static pressure between two chambers that opens the valves (aortic or mitral). After the opening of one of the two valves, the ventricle behaves as a source of flow: blood moves rapidly (velocity > 1.5 m / s) to pass from one chamber to the other. The left ventricular volume (LVV) then varies considerably, decreasing from approximately 200 mL to 120-140 mL during diastole (ventricular blood intake) and conversely during systole (ventricular blood ejection).During the flow source phase, the static pressure remains almost constant, but the dynamic pressure increases considerably. This is the phase during which the core consumes the most energy.

[0170] The walls of the arterial system, for their part, are essentially a source of static pressure, the average blood flow velocities being relatively slow (< 0.5 m / s) in the aorta, and even much slower in the arterioles and then the capillaries. The volume of blood stored in the arterial system varies very little, on the order of 2.5 L + / - 40 mL.

[0171] Compression of the heart muscle thus transfers its energy to the arterial network (including the aorta), which stores it in essentially static form. This transformation is achieved by the so-called Windkessel effect, i.e., the elastic deformation of the aortic walls during systole, which allows the accumulation of static energy that is then used to continuously circulate blood in the cardiovascular system.

[0172] The hydrostatic energy transferred from the left ventricle to the cardiovascular system via the aortic valve during a cardiac cycle (energy called Stroke Work in English) is equal to the area of ​​the static pressure curve PVG vs ventricle volume VolVG, called the PV-loop curve of the LV.

[0173] The hydrostatic energy stored in the entire cardiovascular system is much greater, about 60 times greater, and varies very little over time. It takes about 1 minute (60 cycles) for the left ventricle to transfer the energy equivalent of all that is stored in the arterial system. This roughly corresponds to the fact that the entire blood volume (5 L) completes the circuit every minute.

[0174]

[0175] Thus, static pressure essentially represents a potential energy stored (in elastic tissues during their deformation by swelling, deflation) while dynamic pressure represents the kinetic energy that circulates the blood.

[0176]

[0177] In general, when designing a pump, the aim is to prioritize either a static pressure source or a flow source. The impact on design choices is fundamental. It is very difficult to design a good source that provides both flow and pressure because their characteristics are very different.

[0178] Among the dimensioning elements are the maximum pressure at zero flow and the maximum flow at zero pressure.

[0179] In the case of an LVAD-Bypass, the maximum pressure is the value PvalveAOmax, the maximum pressure difference at the aortic valve. As a reminder, this static pressure is on the order of 100 mmHg.

[0180]

[0181] Static pressure

[0182] The static pressure function of the mechanical cardiac assistance device, denoted Sp(t), is defined by the following relation:

[0183]

[0184] PdispStat(t) = Sp(t) * PvalveAOmax

[0185]

[0186] This pressure function is a characteristic that evolves throughout the entire cardiac cycle (systole and diastole). Spmin and Spmax are introduced as being respectively the minimum and maximum values ​​of the static pressure function of the device Sp(t).

[0187]

[0188] Spmin < Sp(t) < Spmax

[0189]

[0190] The following table illustrates examples of Spmin and Spmax pressure coefficients of different mechanical cardiac assistance devices.

[0191]

[0192] PvalveAÔmax Technology Commentary Total Artificial Heart 100 mmHg The total heart replaces the native heart and imposes exactly the aortic pressure LVAD-Bypass or Trans-aortic (anterior artery) 100 mmHg The bypass replaces native function and must even oppose the native pressure to avoid backflow in the ICOMS device (document FR2955499), anterior artery 100 mmHg The ICOMS is not subjected to the full trans-aortic pressure, thus reducing its energy consumption Device according to the present invention 100 mmHg The FlowMaker implements a synergy at the pressure level, making it possible to eliminate the need for a feeding propeller (reduction in the length of the propeller) and to further reduce the pressure until it is reversed (at least transiently). The FlowMaker can be a hybrid device between a pump and a turbine

[0193] The present invention is thus distinguished from the prior art:

[0194] - Either LVAD-Bypasses whose outlet-inlet pressure difference is equal to the Transaortic pressure, which is virtually constant, is very high. SpminBYPASS > 80%

[0195] - Either of the ICOMS whose outlet-inlet pressure difference is small compared to to transaortic pressure but constantly positive. SpminICOMS < 60% and SpminICOMS > 0

[0196]

[0197] Fig. 3 illustrates flow / static pressure operating point curves measured in a static performance characterization bench of the ICOMS and the cardiac mechanical assistance device according to the invention, named FlowMaker.

[0198] A first representation of the instantaneous flow rate Ddisp (L / min) as a function of the static pressure difference PdispStat (mmHg) is shown by the solid-lined curves. For the mechanical cardiac assistance device according to the invention, FlowMaker, the flow rate varies from 6 L / min to zero when the pressure difference varies from zero to approximately 30 mmHg. For Icoms, the flow rate varies from 9 L / min to zero while the pressure difference varies from zero to 90 mmHg.

[0199] A second representation of the instantaneous mechanical power Pmi (watt) as a function of the static pressure difference PdispStat (mmHg) is shown by the dashed curves. The FlowMaker reaches a maximum of 0.0 watts for 10 mmHg, while the icoms reaches a maximum of 6.6 watts for 50 mmHg.

[0200] The mechanical power is estimated with a conversion efficiency of 10%.

[0201] It is observed that the FlowMaker differs from the ICOMS by the strong reduction in pressure at zero flow and by the possibility of operating at negative static pressure in a transient manner.

[0203] Figure 4 shows a typical example of the hydrostatic pressure developed by the device according to the invention implanted in a left ventricle as a function of time over a cardiac cycle. This is the static pressure difference between the outlet and inlet of the hydraulic module implanted in the left ventricle during a cardiac cycle. Systole lasts 300 ms between times t=100 ms and t=400 ms. A maximum static pressure difference of 6 mmHg is observed, significantly lower than that of an LVAD-Bypass or an ICOMS. A negative transition is even observed at the start of the second systole (t=0.85 s), due to the design of the present invention.

[0204]

[0205] Dynamic pressure.

[0206] The FlowMaker is designed as a flow source and not as a pressure source. The FlowMaker's operation relies on synergistic effects to increase the efficiency of the native pump, not to replace it. Its action occurs primarily during systole and contributes to increasing aortic flow. Under no circumstances does the FlowMaker contribute to generating the pressure necessary to open the aortic valve. On the other hand, the increase in aortic flow helps restore static pressure in the arterial system, which performs the transformation of kinetic energy into static pressure.

[0207]

[0208] Dynamic pressure is a misnomer because it represents a dynamic phenomenon that has nothing to do with pressure: it represents the kinetic energy of the fluid, i.e., its velocity. The equation is non-linear since the kinetic energy of a volume of fluid is proportional to the square of its velocity.

[0209] Note that the behavior of the fluid must be considered at different scales to be understood. At the atomic scale and even at the scale of hemoglobin cells, the fluid motion (1 Opm) is governed by random laws of the Brownian motion type. For scales between 1 Opm and 1 mm, the fluid is subject to turbulent motions. These local motions are not considered in the problem that interests us, namely displacements on scales between 1 mm and 10 cm.

[0210] The typical diameter of an aortic valve varies from 20 mm to 30 mm. The length of the outflow tract is of the same order of magnitude. The length of the left ventricle is on the order of 50 mm to 120 mm depending on the patient. Its width varies considerably during contraction between 10 mm and 50 mm. For patients with heart failure, the width increases significantly in most patients, this which leads to a reduction in compliance and therefore a vicious cycle of degradation in the heart's efficiency.

[0211] The timescale is also important to consider. On very short timescales, less than 1 ms, there are also generally turbulent movements, i.e., the fluid swirls around itself quite rapidly and chaotically. We are interested in timescales between 1 ms and 1 s. As a reminder, a complete cardiac cycle lasts approximately 1 s, and systole itself lasts only about 300 ms with an initial flow peak lasting on the order of 100 ms.

[0212] On the considered dimensional and time scales, the "average" behavior of the fluid during the ejection phase, in a healthy patient and at the level of the left ventricular outflow tract, is virtually uniform. The average fluid velocity at the considered time and space scales is on the order of m / s to 1.5 m / s. It is understood that this behavior is averaged both in time (say, over 1 ms) and in space (say, over 1 mm³). It is also understood that this does not concern the behavior during diastole, i.e., the filling of the left ventricle, during which the behavior is highly turbulent at these scales, including at the level of the left ventricular outflow tract.

[0213] It is observed that for most sick patients suffering from severe heart failure, there are two recurring (but not exhaustive) configurations:

[0214] Type 1. Either the flow is relatively uniform over the time and space scales considered, during the ejection phase and in the outflow tract, but with a blood velocity that is too low, typically less than 0.6 m / s or even 0.4 m / s. This can be due to either an increase in systemic resistance, a reduction in arterial compliance, or a reduction in myocardial contractility, for example. In this case, the work of the myocardium is reduced compared to that of a healthy patient. The internal hydraulic resistance of the native pump is "normal," but the contractile force is insufficient to oppose the load impedance (systemic resistance in parallel with arterial compliance).

[0215] Type 2. Either the flow is significantly turbulent, including on the time and space scales considered, and including during the ejection phase and in the outflow tract. This may be due to left ventricular hypertrophy, for example. The consequence is that the blood, although its velocity is "normal" in intensity, tends to vortex in the outflow tract and reduce the volume ejected by the left ventricle. The direct consequence is a sharp reduction in aortic flow, even though the work done by the myocardium is greater than that of a healthy patient. The internal hydraulic resistance of the native pump is much higher than that of a healthy patient.

[0217] It is possible that type 1 pathologies may eventually transform into type 2 pathologies, thus accumulating all the negative effects. In the vast majority of cases, the left ventricle's contractile capacity is impaired to such an extent that the physician is forced to prescribe the maximum doses of medications intended to lower aortic pressure (below 70 mmHg), which induces clear clinical symptoms requiring hospitalization in intensive care with identification of a state of shock for the patient.

[0218] It should be noted that the internal hydraulic resistance of the outflow tract, although very low in absolute value compared to the systemic resistance, is indeed one of the factors directly limiting flow during the ejection phase. In fact, the systemic resistance is bypassed by the impedance of arterial compliance, the impedance of which is very low over the time scales considered. If this compliance is too low, then it also limits aortic flow.

[0219]

[0220] The objective of the present invention is to significantly reduce the internal hydraulic resistance of the left ventricle, thanks to the synergistic action of the cardiac mechanical assist device. To this end, the device must be capable of generating, during the ejection phase, a "jet-stick" fluid flow, i.e., over a substantially cylindrical column of small diameter compared to the diameter of the aortic valve. The definition of a jet-stick could be "more than 50% of the power delivered in the form of a rectilinear velocity field contained within a small-diameter cylinder over a large distance."

[0221] The advantage of such a "stick throw" is twofold, since it allows one to: - Eliminate any type 2 turbulence by literally pushing the fluid towards the aortic valve. With the "plug" being greatly reduced (or even eliminated) at the start, ejection is properly initiated and there follows a significant increase in flow throughout the duration of systole, almost independently of the duration of forced ejection. - Significantly reduce the hydraulic resistance of the outflow tract through an amplification effect. The fluid velocity cannot be discontinuous (Bernoulli's principle), but is zero at the walls. Thus, an increase in velocity in the blood column exiting the pump, even over a small diameter, leads to an increase in velocity across the entire cross-section of the outflow tract. This is the amplification effect; highly non-linear. In this case, the increase in aortic flow is proportional to the duration of the forced ejection. Note, however, that the amplification is greater the more viscous the fluid. Thus, unlike conventional devices that require treatment with anti- coagulant (a lifelong drug treatment), it is possible to reduce or even eliminate this constraint by using the present invention.

[0222]

[0223] To be effective, the jet diameter is preferably significantly smaller than the flushing chamber diameter. This has a dual effect: - to increase the fluid velocity at the outlet of the mechanical cardiac assistance device, and - to increase the amplification phenomenon.

[0224] The effective diameter of the pump outlet is preferably less than 12 mm, or even 10 mm, or even 8 mm, or even 6 mm, or even 5 mm. The fluid velocity at the device outlet is preferably greater than 0.8 m / s, or even 1 m / s, or even 1.5 m / s, or even 2 m / s.

[0225]

[0226] Figure 5 shows the evolution over time of the instantaneous flow rate and the average blood velocity in the forced fluid column exiting the mechanical cardiac assistance device according to the invention. The diameter of the cylindrical forced column is approximately 6 mm. The maximum instantaneous flow rate is 6 L / min. The average velocity during the ejection phase is approximately 2.5 m / s.

[0227] A blood flow rate of approximately 3 L / min is observed at the device outlet during diastole. This corresponds to a continuous "washing" of the mechanical cardiac assistance device to prevent stagnation and thrombosis.

[0228]

[0229] The hydrodynamic pressure of the mechanical cardiac assist device according to the invention is shown in [Fig. 6]. It has a maximum of 27 mmHg, significantly higher than the 7 mmHg static pressure. It can be observed that, although the device's "wash" flow rate is only half that required to force ejection, the dynamic pressure is four times lower, on the order of 7 mmHg, since it varies with the square of the velocity. Note that this dynamic pressure cannot disrupt the opening of the mitral valve since it is both low and spatially localized to the outflow tract. Care must be taken not to direct the fluid column towards the mitral valve, which is achieved by means of the adjustable fixation system.

[0230]

[0231] In the case of an LVAD-Bypass, the static pressure (100mmHg) is much greater than the dynamic pressure, on the order of 10mmHg.

[0232] In the case of an ICOMS, the static pressure was of the same order of magnitude as the dynamic pressure, around 50mmHg.

[0234] Figure 7 shows the flow amplification effect. The instantaneous flow rate at the device outlet is 6 L / min. A maximum flow rate of 12 L / min is recorded when the device is inactive. A maximum flow rate of 21 L / min is recorded when the device is activated. This results in a synergistic flow coefficient Sd of approximately 150%.

[0235]

[0236] Device operation.

[0237] The instantaneous hydraulic power of the device is determined by the product of the pressure difference and the flow rate.

[0238] Pow_disp(t) = Ddisp(t)*Pdisp(t) in Watt

[0239]

[0240] Instantaneous power can also be decomposed into two terms, corresponding to static and dynamic power:

[0241] Pow_disp(t) = Pow_dispStat(t) + Pow_dispDyn(t)

[0242]

[0243] The first part corresponds to the device's ability to increase the hydraulic potential energy of the ejected fluid volume, while the second part corresponds to the increase in the kinetic energy of the ejected fluid volume, and therefore its velocity. These quantities depend strongly on the device design, the rotational speed of the device's rotor, and the device's load conditions, i.e., its environment.

[0244] In the present invention, dynamic power has been favored since the blood flow has a high exit velocity.

[0245] Figure 8 represents the time evolution of the two components of the instantaneous hydraulic power. The device's rotational speed typically varies from QMax = 4000 rpm during the ejection phase when the aortic valve is open to QMin = 2000 rpm during the diastolic phase when the aortic valve is closed, representing a factor of 2 between these two rotational speeds. In this simulation, the device activation time (Q = Qmax) corresponds to the ejection time, approximately 300 ms.

[0246] It is already observed that the dynamic power is significantly greater than the static power. In particular, during the ejection phase, there is a ratio of approximately 1 to 5. For an LVAD-Bypass, the ratio is almost reversed (more static and less dynamic). For an ICOMS, the ratio is almost one.

[0247] It is also observed that the effect of the device's "washing" during the systolic phase does not consume much energy, although the speed is only halved. Indeed, the power varies globally as the cube of the rotational speed. The speed QMin can obviously be reduced, for example to 1000 rpm in order to Divide this ratio by 8 again. However, a minimum rotation is necessary to avoid stagnation in the hydraulic module.

[0248]

[0249] The hydraulic work of the device over a cardiac cycle of duration T can finally be defined as the hydraulic energy supplied by the device to the fluid during said cycle. Again, it can be decomposed into two terms related to static and dynamic work. The work is expressed in Joules and represents the energy transferred to the fluid, and consequently to the cardiovascular system, by the action of the device. The device's power consumption is a function of this work and the design. From the perspective of an electrical energy source, for example, an implantable battery, the total electrical consumption of the device is defined as the product of the hydraulic work transmitted to the fluid and a conversion efficiency.Efficiency can be broken down into three parts: electrical conversion efficiency for the electronics located between the power source (battery) and a motor, electromechanical conversion efficiency for the motor, and mechanical-hydraulic conversion efficiency for the fluidic part of the device. Typically, at the optimal operating point, an electrical efficiency of 90%, an electromechanical efficiency of 85%, and a mechanical-hydraulic efficiency of 15% can be achieved for a rotary axial pump, resulting in a total efficiency of approximately 10%. The actual power consumption of a device is therefore roughly ten times greater than the work it performs.

[0250] The present invention makes it possible to improve the mechanical-hydraulic efficiency.

[0251] Travan_diSp(T) = f^Pow _distffydt

[0252]

[0253] Available_Work(t) = Available_WorkStat(t) + Available_WorkDyn(t)

[0254]

[0255] Figure 9 illustrates the evolution of the hydraulic work transmitted to the fluid by the cardiac mechanical assistance device according to the invention. It shows that this work is essentially hydrodynamic in nature. Approximately 80% of the energy is converted into kinetic energy. Recall that for an LVAD-Bypass, the opposite is true: 80% is converted into static energy. For an ICOMS, it is around 50%. It also shows that the device in this example operates primarily during the ejection phase (between 100 ms and 400 ms).

[0256]

[0257] We can now analyze the hydraulic energy transferred by the left ventricle to the arterial network during a cardiac cycle. We begin by defining Pow_Ao(t) as the total instantaneous hydraulic power transferred to the arterial network via The aorta. It is the product of the instantaneous total aortic pressure Pao(t) and the instantaneous aortic flow rate Dao(t). In reality, the integration must be performed over the entire surface of the aorta because the surface pressure density and the fluid velocity are not uniform quantities.

[0258]

[0259] I \JLj7 IJH,... _ZA---------- = I , . dS » Pao(t) * Dao(t) ace aorta dS ' "• 7

[0260] Figure 10 shows this instantaneous hydraulic power during a cycle, with and without the active device. It can be observed that the maximum power is almost doubled, increasing from 2.5W to 5W. However, the device only provides 0.3W. The instantaneous power perspective can be misleading because synergy can only be discussed when calculating the cumulative effects over an entire cardiac cycle. This is why it is preferable to define synergy as an energy ratio.

[0261]

[0262] The energy supplied by the left ventricle to the arterial network over one cycle is then calculated as the time integral of the instantaneous hydraulic power. The final value of this energy, denoted Ener_ao (T), calculated over a duration T, is equal to the work done. The difference in work done via the aorta with and without the device can then be calculated, and the additional work input deduced. The ratio between this additional work and the work done by the device can then be calculated to deduce the synergy coefficient Sen. Figure 10 illustrates the three energies. The work done via the aorta is measured to be 0.8 J when the device is inactive and 1.35 J when the device is active, an increase of 0.55 J. The work done by the device over this same cycle is 0.13 J. The synergy coefficient is therefore 400%. 102631 Sen = EnerV)avecDevice(.T)-EnerAOsaiisDevice(T) Work_Disp(T)

[0264]

[0265] The increase in energy supplied by the left ventricle to the aorta can be estimated quite accurately in practice. It is necessary to begin by determining the stroke volume, denoted ΔLVL, which corresponds to the difference in the left ventricular volume at the end and beginning of systole. ΔLVL can be measured by simple, conventional echocardiographic techniques. Next, the mean aortic pressure, denoted PaoMoy, is determined. The work done by the ventricle via the aorta is the product of the two quantities mentioned above.

[0266] To determine the work done by the device, its electrical consumption can be measured and multiplied by the electro-mechanical-fluid efficiency determined in vitro.

[0268] DeltaVoIVG.PaoMqy|arec dispcisjt^ disjMsüif Work_Dis^T}

[0269]

[0270] Thus, with the mechanical cardiac assistance device according to the invention, at each cardiac cycle, the hydraulic energy transmitted to the cardiovascular system via the aorta when the hydraulic module is actuated is strictly greater than the sum of the hydraulic energy transmitted by the ventricle when the hydraulic module is deactivated and the hydraulic work provided by the hydraulic module.

[0271]

[0272] Of course, the invention is not limited to the examples just described. Many modifications can be made to these examples without departing from the scope of the present invention as described.

Claims

Demands

1. 1. A mechanical cardiac assistance device (10) implantable in a heart (1) and comprising a hydraulic module (11) intended to be disposed entirely inside the left ventricle (3) of the heart, this hydraulic module (11) comprising an inlet (18) for blood flow and an outlet (16) of the blood flow into the ventricle and to the aortic valve (7) of the heart, a base (12) intended to be fixed on the ventricle (3), a controller (13) capable of controlling the hydraulic module, characterized in that: - the outlet (16) of the hydraulic module has an internal surface area of ​​less than 113.1mm2 - the controller (13) is configured to automatically and dynamically control the velocity of the blood flow according to a velocity profile synchronized with the heart's pulse, and to control the blood flow velocity to a value greater than or equal to vmax for an ejection time Tpulse;vmax being greater than or equal to 0.6m / s and the ejection time Tpulse being less than or equal to a duration Tsystole of a systole of the cardiac cycle, - a first measuring device to measure an evolution of the volume of the ventricle during a cardiac cycle, the ejection time Tpulse and the maximum velocity vmax being a function of this evolution of the volume of the ventricle, and - a second measuring device to measure the viscosity of the blood, the ejection time Tpulse and the maximum velocity vmax being a function of this viscosity of the blood.;

2. 2. Device according to claim 1, characterized in that the outlet (16) of the hydraulic module has an internal surface area less than 78.5mm2, 50.24mm2, 28.26mm2 or 19.62mm2, corresponding to an equivalent internal diameter of the outlet (16) of the hydraulic module less than 10mm, 8mm, 6mm or 5mm.

3. 3. Device according to claim 1 or 2, characterized in that the maximum velocity vmax of the blood flow exiting the hydraulic module is greater than or equal to 0.8m / s, 1m / s, 1.5m / s, or 2m / s.

4. 4. Device according to any one of the preceding claims, characterized in that the hydraulic module (11) comprises a helical rotor (19) whose diameter-to-length ratio (19) is greater than or equal to 0.

54.

5. 5. Device according to any one of the preceding claims, characterized in that the base (12) comprises a reversible electromechanical converter capable of converting both electrical energy into mechanical energy according to motor operation, and mechanical energy into electrical energy according to generator operation.

6. 6. Device according to any one of the preceding claims, characterized in that the hydraulic module (11) comprises a box having an external profile straight or flared towards the outlet.

7. 7. Device according to claim 6, characterized in that the box has an external surface traversed by a helical groove (20).

8. 8. Device according to claim 6 or 7, characterized in that the box has an internal surface traversed by a helical groove.

9. 9. Device according to any one of the preceding claims, characterized in that the hydraulic module is of axial type and comprises a helical rotor (19) of length less than or equal to 25 millimeters.

10. 10. Device according to claim 4 or 9, characterized in that the helical rotor (19) is a propeller whose ratio between the diameter of the blades, including the hub, and the diameter of the hub is greater than 2.

11. 11. Device according to any one of the preceding claims, characterized in that the first measuring device comprises at least one cardio-impedance measuring probe (21) for measuring the evolution of the left ventricle volume.

12. 12. Device according to any one of claims 1 to 10, characterized in that the first measuring device comprises an external echocardiography system for measuring the evolution of ventricular volume.

13. 13. Device according to claim 10 or 11, characterized in that the first measuring device is configured to measure a final systolic volume (VolVGmin) and a final diastolic volume (VolVGMax) of the ventricle.

14. 14. Device according to any one of the preceding claims, characterized in that the second measuring device comprises at least one fluid viscosity measuring probe.

15. 15. Device according to any one of the preceding claims, characterized in that the second measuring device comprises an external viscometer system for measuring blood viscosity.

16. 16. Device according to any one of the preceding claims, characterized in that the ejection time Tpulse is less than 300ms, 200ms or 100ms.

17. 17. Device according to any one of the preceding claims, characterized in that the hydraulic module comprises a stator in which is placed at least one first magnet (22) intended to cooperate magnetically with at least one second magnet (23) placed in the helical rotor (19) of the hydraulic module so as to maintain the rotor in the axis of rotation.

18. 18. Device according to any one of claims 4 to 17, characterized in that the controller (13) is configured to control a dynamic pressure at the outlet of the hydraulic module according to a predetermined time profile, the dynamic pressure being a function predetermined by the design of the hydraulic module, this function depending on the rotational speed Q of the rotor and the viscosity of the blood.

19. 19. Device according to any one of the preceding claims, characterized in that the controller (13) is configured to activate the hydraulic module only every "n" cardiac cycles, "n" being a function of a ventricular end-diastolic volume.

20. 20. Device according to any one of the preceding claims, characterized in that it comprises a fastening device (15) in which the hydraulic module is inserted in such a way as: - orientable for an inclination of the axis of rotation of the hydraulic module relative to the axis of the flushing chamber, and - removable for withdrawal or displacement along the axis of rotation of the hydraulic module.

21. 21. Device according to any one of the preceding claims, characterized in that it comprises a third measuring probe for measuring the electrical activity of the heart and thus synchronizing the speed of the hydraulic module with the cardiac activity.

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