CONTROLLER AND METHOD FOR AN ARTIFICIAL HEART - Patent application

Abstract

The present invention relates to a controller unit (100) and a method for controlling an intracardiac prosthesis (200), the prosthesis comprising: at least one pump section (202, 203, 602, 702); an inlet (210, 610, 710) connected to said at least one pump section; an outlet (213, 613, 713) connected to said at least one pump section; a pressure sensor (231; 232) configured to measure the pressure of a fluid flowing from the inlet to the outlet; and a pump actuator (221, 222) configured to induce a fluid flow. The controller unit further comprises a memory and a processing unit, and is configured to: obtain pressure values ​​from the pressure sensor, obtain a desired value of the pressure of the fluid flowing in the pump, calculate an error signal equal to the difference between the desired value of the pressure and the measured pressure, and control the output of the pump so that the measured pressure is close to or equal to the desired pressure by controlling the pump stroke speed and / or the pump stroke volume.

Description

[Technical Field]

[0001] In general, the present invention relates to control methods and devices for artificial hearts, and in particular to the response of artificial heart systems to a wide range of physiological demands, including mechanisms for adapting to actual flow imbalances between the pulmonary and systemic circulations. [Background technology]

[0002] Despite steady progress in the development of permanent artificial hearts for long-term implantation in patients as a replacement for failing natural hearts, several challenges remain to be overcome. Among the challenges that need to be addressed in untethered artificial heart systems are control strategies that respond to varying physiological demands and mechanisms for adapting to flow imbalances between the pulmonary and systemic circulations.

[0003] The difference in cardiac output between the left and right sides is well documented. Physiologically, blood flow pumped by the left side of the heart is greater than blood flow pumped by the right side. This difference is due in large part to a circulatory pathway known as the bronchial shunt. This flow originates in the left atrial system, passes through the bronchial tissue, and then returns directly to the left atrium. This difference typically appears to be a maximum of approximately 10% of cardiac output, with left-sided flow always greater than right-sided flow. Artificial heart systems must take this inherent physiological circulatory imbalance into account. In addition, sources of flow imbalance can be created. For example, differences in regurgitation through left- and right-sided artificial valves can cause flow imbalance. Artificial heart systems must also take these types of circulatory imbalance into account.

[0004] A blood pumping device by the same applicant is disclosed in Patent Document 1, which includes at least a first pump, a second pump, and left and right pump actuation means for inducing blood flow within the body's circulatory system. Each pump includes an upper chamber having an inlet channel and a lower chamber having an outlet channel. The upper and lower chambers are separated by a movable valve plane provided with a valve. The pump actuation means applies upward and downward movement of the valve plane between the upper and lower chambers in response to control signals from a control unit, such that when the valve plane moves upward, the valve provided in the valve plane is in an open position, allowing blood to flow from the upper chamber to the lower chamber, and when the valve plane moves downward, the valve is in a closed position, allowing blood to be ejected from the lower chamber through the outlet channel. A bag-like portion is provided at the bottom of the lower chamber.

[0005] FIG. 1 shows a schematic cross-sectional view of the four-chamber blood pumping device 1 of Patent Document 1. The four-chamber blood pumping device 1 includes two pumps, i.e., a first pump 2 and a second pump 3, and first and second pump actuators for inducing blood flow in the body's circulatory system. The first and second pumps 2 and 3 are identical in structure. Each pump includes an upper chamber 9 and a lower chamber 12. The upper chamber 9 has an inlet channel (not shown) that allows blood to enter the upper chamber 9. The upper chamber corresponds to the atrium of a natural heart. The lower chamber 12 has an outlet channel (not shown) that allows blood to exit the lower chamber 12. The lower chamber 12 corresponds to the ventricle of a natural heart. A movable valve plane 7 separates the upper and lower chambers 9 and 12. The valve plane corresponds to the atrioventricular (AV) plane (i.e., the plane of fibrous tissue) between the atria and ventricles of a natural heart. The valve 14 is located within the valve plane 7 and corresponds to the tricuspid or mitral valves that function in the pulmonary or aortic circulation, depending on whether it is located within the pump. The bottom of the lower chamber 12 is advantageously designed to have a shape similar to the anatomical shape of a ventricle in a natural heart. In the four-chamber blood pumping device 1, the bottom of the lower chamber 12 has a bag-like shape designed to mimic the internal shape of a ventricle in a natural heart. Blood flow (arrows) entering the lower chamber 12 from the upper chamber 9 through the valve 14 encounters a stop surface at the bottom of the bag-like shape and abruptly stops, whereupon the flow abruptly changes direction and continues along an outlet flow path (not shown). The internal turn of the bag-like portion at the bottom of the lower chamber 12 forms a bend of approximately 90 to 340°, more preferably 100 to 300°, more preferably 105 to 200°, and most preferably 110 to 150°, similar to the internal bend of a ventricle in a natural heart. The blood then continues into the outlet channel. The cross section of the back-shaped portion at the bottom of the lower chamber 12 advantageously has a triangular shape to allow optimal blood flow from the lower chamber 12 into the outlet channel.As in the natural heart, the triangular cross-section facilitates the formation of a flow path within the cavity of the lower chamber 12, allowing blood to approach the stop surface of the bag-shaped portion from different angles, stop, change direction, enter the outlet flow path, and then exit the blood pumping device through the outlet valve. Alternatively, the cross-section of the interior structure of the lower chamber 12 can have an oval or circular cross-section. The inner wall at the bottom of the lower chamber 12 and the outlet flow path are advantageously provided with a roughened surface simulating trabeculae, i.e., the muscular ridges that move vertically and horizontally and protrude from the inner wall of the ventricles of a natural heart. The roughened surface is covered with ridges and protrusions that protrude approximately 0.01 to 3 mm, preferably at least 0.5 to 2 mm, from the surface of the lower chamber 12. The outlet flow path and the bottom of the lower chamber 12 can also have a smooth surface. The outlet flow path from the lower chamber 12 can also have a continuously decreasing diameter, similar to the design of the outlet of the ventricles of a natural heart.

[0006] The same applicant's U.S. Pat. No. 6,269,623 relates to a blood pump housing device designed to enclose and protect a total artificial heart when implanted in a subject. The blood pump housing device includes first and second artificial heart pump receiving members configured to receive and partially enclose first and second artificial heart pumps of a total artificial heart (TAH), and first and second pump actuation enclosing members configured to partially house first and second pump actuation means. The artificial heart pump receiving member and the pump actuation means enclosing members are arranged to couple to each other in a leak-tight manner. [Prior art documents] [Patent documents]

[0007] [Patent Document 1] WO2016 / 020219 [Patent Document 2] WO2017 / 137486 Summary of the Invention [Problem to be solved by the invention]

[0008] Generally, there is a need for a control device and method for automatically controlling at least the heart rate and stroke volume of each blood pump device, and in particular the pump devices mentioned above. [Means for solving the problem]

[0009] The present invention also provides for the response of the artificial heart system to various physiological demands, including mechanisms to adapt to actual flow imbalances between the pulmonary and systemic circulation.

[0010] For these reasons, a method for controlling an intracardiac prosthesis is provided. The intracardiac prosthesis includes: at least one pump section; an inlet connected to the at least one pump section; an outlet connected to the at least one pump section; a pressure sensor configured to measure the pressure of a fluid flowing from the inlet to the outlet; a pump actuator configured to induce a fluid flow; and a controller unit. The method includes: obtaining a pressure value from the pressure sensor; obtaining a desired value of the pressure of the fluid flowing through the pump; calculating an error signal equal to the difference between the desired value of the pressure and the measured pressure; and controlling the output of the pump so that the measured pressure approaches or equals the desired pressure by controlling the pump actuator to control the pump stroke speed and / or pump stroke volume. In one embodiment, the fluid is blood. The output can be cardiac output. According to one embodiment, the pump includes a chamber corresponding to one of the right or left atrium.

[0011] In one embodiment, the intracardiac prosthesis may include two similar pumps, connected to the systemic and pulmonary circulations, respectively, and both controlled individually by setting a limit on the cardiac output of the pump connected to the pulmonary circulation. To this end, the method comprises the steps of: obtaining the cardiac output of the pump connected to the systemic circulation; setting a limit on the cardiac output when the pump is connected to the pulmonary circulation, given the cardiac output of the pump connected to the systemic circulation, the limit being updated as the cardiac output of the pump connected to the systemic circulation changes; and providing a control signal to said pump actuator. Further includes:

[0012] In one embodiment, two pumps are connected to the systemic and pulmonary circulations, respectively, having desired cardiac outputs, and a stroke volume and stroke rate for the two pumps are achieved, the method comprising: achieving the desired cardiac output for either pump; calculating a stroke rate using the desired cardiac outputs for the two pumps; achieving a stroke volume for either pump, given the achieved stroke rate, such that the product of the stroke rate and the stroke rate equals the respective desired cardiac output; and providing a control signal to said pump actuator.

[0013] In one embodiment, a pressure sensor is positioned in communication with the chamber. In an alternative embodiment, the pressure is measured at a location between the inlet opening and the end of the chamber forming the atrium. One embodiment includes measuring the pressure within the thoracic cavity as the reference pressure.

[0014] In one embodiment, the controller unit is configured to detect whether atrial pressure drops or intrathoracic pressure increases, i.e., a suction event, and function to prevent atrial pressure drops or intrathoracic pressure increases. To this end, the control unit compares the atrial pressure received from the pressure sensor, averaged over the course of a stroke, with the desired atrial pressure, and if the average atrial pressure is much lower than the desired pressure, a suction event is detected. In a further step, to reduce noise and prevent false suction event detection, the atrial pressure is averaged over the course of a cycle, and if a suction event is detected by the control unit, an inactive period is provided during which no suction event is detected. If a suction event is detected, further suction events are prevented, an alert is generated, and a gradual increase in the desired atrial pressure is implemented.

[0015] The present invention also relates to a controller unit for controlling an intracardiac prosthesis, the prosthesis comprising at least one pump section; an inlet connected to the at least one pump section; an outlet connected to the at least one pump section; a pressure sensor configured to measure the pressure of a fluid flowing from the inlet to the outlet; a pump actuator configured to induce a fluid flow, a memory, and a processing unit. The controller unit is configured to: obtain pressure values ​​from the pressure sensor, obtain a desired value of the pressure of the fluid flowing in the pump, calculate an error signal equal to the difference between the desired value of pressure and the measured pressure, and control the output of the pump so that the measured pressure approaches or equals the desired pressure by controlling the pump stroke speed and / or the pump stroke volume. In one alternative embodiment, the controller unit includes: a flow control functional block configured to determine the correct flow limit and desired atrial pressure as inputs to keep the cardiac output within range; a cardiac output determination functional block that determines the cardiac output of either pump to control the atrial pressure; and a heart rate determination functional block, where the heart rate and cardiac output of either pump provide the stroke volume of either pump.

[0016] In one embodiment, the intracardiac prosthesis includes first and second pumps, and the controller unit is configured to keep the flow rate of the first pump low enough so that the second pump does not reach its maximum flow rate. In an alternative embodiment, the intracardiac prosthesis includes two similar pumps, connected to the systemic and pulmonary circulations, respectively, both of which are individually controlled by setting a limit on the cardiac output of the pump connected to the pulmonary circulation, and the controller unit is further configured to: obtain the cardiac output of the pump connected to the systemic circulation; given the cardiac output of the pump connected to the systemic circulation, set a limit on the cardiac output when the pump is connected to the pulmonary circulation, the limit being updated as the cardiac output of the pump connected to the systemic circulation changes; and provide a control signal to the pump actuator.

[0017] In one embodiment, the intracardiac prosthesis includes two similar pumps connected to a systemic circulation and a pulmonary circulation, respectively, having a desired cardiac output, and achieving a stroke volume and a stroke rate for the two pumps, the controller unit being further configured to achieve the desired cardiac output for either pump, to calculate the stroke rate using the desired cardiac outputs of the two pumps, to achieve a stroke volume for either pump given the achieved stroke rate such that the product of the stroke rate and the stroke rate equals the respective desired cardiac output, and to provide control signals to said pump actuators.

[0018] According to one embodiment, the controller unit includes a signal receiver that receives a signal and functions to detect whether atrial pressure is low or intrathoracic pressure is increasing, i.e., a suction event, and prevent the low atrial pressure or intrathoracic pressure increase. In one embodiment, the processing unit of the controller unit is configured to compare the atrial pressure received from the pressure sensor, averaged over the course of a stroke, with the desired atrial pressure, and detect a suction event if the average atrial pressure is much lower than the desired pressure. In another embodiment, the processing unit of the controller unit is configured to average the atrial pressure over the course of a cycle to reduce noise and prevent false suction event detection, and to provide an inactive period during which no suction event is detected when a suction event is detected by the control unit. In one embodiment, when a suction event is detected, the controller unit is configured to implement a gradual increase in the desired atrial pressure and generate an alert to prevent further suction events.

[0019] The invention also relates to an intracardiac prosthesis including a controller unit as described above.

[0020] The present invention further relates to a pressure sensor for use in an intracardiac prosthesis, the prosthesis comprising a housing having an upper part and a body, an inlet, an outlet and a chamber between the inlet and the outlet, the pressure sensor comprising: a flexible membrane covering an opening in the upper part or body, a pressure transmission medium, a pipe containing the pressure transmission medium and a pressure-sensitive sensor.

[0021] The invention also relates to an intracardiac prosthesis including such a pressure sensor.

[0022] Reference is now made to the accompanying drawings, in which elements having the same reference numbers may represent similar elements throughout. [Brief explanation of the drawings]

[0023] [Figure 1] FIG. 1 is a diagram of a four-chamber blood pumping device according to the prior art. [Figure 2] 1 is a schematic diagram of a control device according to the present invention connected to a blood pumping device. [Figure 3] FIG. 2 is a functional block diagram of a control unit according to the present invention; [Figure 4a] FIG. 1 is a schematic diagram of a PID controller according to one embodiment of the present invention. [Figure 4b] FIG. 1 is a schematic diagram of a PID controller according to one embodiment of the present invention. [Figure 5] 1 is a graph of an exemplary ratio between heart rate and blood flow. [Figure 6] 1 is a perspective view of an artificial heart according to one embodiment. FIG. [Figure 7] Figure 7a is a cross-sectional view of an artificial heart pump according to one embodiment of the present invention, and Figure 7b is an enlarged view of the circled area in Figure 7a. [Figure 8] FIG. 2 is a schematic diagram of a control unit according to the present invention. [Figure 9] 1 shows a perspective view of one pump embodiment that is part of an intracardiac prosthesis. [Figure 10] 10 shows a cross-sectional view of the pump of FIG. 9. [Figure 11] 1 is a highly schematic illustration of a cross section of the rib cage and thoracic cavity. FIG. DETAILED DESCRIPTION OF THE INVENTION

[0024] The following detailed description refers to the accompanying drawings, in which the same reference numbers in different drawings may represent the same or similar elements.

[0025] The term artificial heart, as used herein, relates to a pumping device that is connectable to a subject.

[0026] The term "stroke volume" as used herein relates to a specific amount of fluid displaced in a specific time interval, and the term "heart rate" as used herein relates to the rate at which a pump displaces a certain amount of fluid.

[0027] The terms "right atrium" and "left atrium" as used herein refer to the compartments within the pump housing into which fluid is provided and discharged.

[0028] In the following, a control device and method for setting the heart rate and stroke volume of one side of an artificial heart, such as those described in U.S. Pat. No. 5,629,497 or U.S. Pat. No. 5,629,497, will be described. However, it should be clear that the device and method according to the invention can be applied to any pumping device having a chamber through which a fluid can pass. The controller and method described herein can also be used, mutatis mutandis, in cardiac assist pumps.

[0029] 2 shows a highly schematic control unit 100 and blood pump device 200 according to one exemplary embodiment of the present invention. The blood pump 200 includes a first pump 202 and a second pump 203 and first and second pump actuation means 221, 222 for inducing blood flow. Each pump includes one upper chamber 209 with an inlet channel 210 and one lower chamber 212 with an outlet channel 213. The upper and lower chambers are separated by a movable valve plane 207 provided with a valve 214. Each pump actuation means 221, 222 is configured to apply upward and downward movement of the valve plane 207 between the upper and lower chambers in response to a control signal from the control unit 100, so that when the valve plane 207 moves upward, the valve 214 is in an open position, allowing blood flow from the upper chamber 209 through the inlet 210 into the lower chamber 212, and when the valve plane moves downward, the valve is in a closed position, allowing blood to be expelled from the lower chamber 212 through the outlet channel 213. Preferably, the bottom of the lower chamber 212 is provided with a bag-like portion (not shown, but shown schematically in FIG. 1).

[0030] The actuation means 221 and 222 may include magnetic or electromagnetic motors, or any other suitable drive system capable of receiving a signal and driving or displacing the pump parts directly or in cooperation with a gearbox (not shown).

[0031] The blood pump system includes pressure sensors 231 and 232. The sensors and their functions are described in more detail below. The pressure sensors are configured to measure the average atrial pressure over each heartbeat. This average can help reduce any noise in the measurement signal.

[0032] The intrathoracic pressure changes continuously during breathing. Since the intrathoracic pressure influences the pressure inside the atria (209), it is desirable to also measure the intrathoracic pressure (baseline pressure). In this way, it is possible to compensate for changes in the intrathoracic pressure. The intrathoracic pressure can be measured by: 1. Using a pressure sensor that is integrated into the pump but is designed to measure the pressure inside the thoracic cavity; 2. Using a separate pressure sensor placed inside the thoracic cavity and connected to the pump by wiring, or 3. Placing the pressure sensor outside the thoracic cavity, but inside the chest wall and strictly behind the pleural layer (the innermost layer of the chest wall).

[0033] For this purpose, the pressure sensor can be connected to the pleura (pleural) membrane, which acts as a natural flexible membrane between the chest cavity and the pressure sensor. To eliminate any influence of breathing, it is sufficient to use one pressure sensor to measure the pressure inside the chest cavity for both the right and left pumps, the measurement being made outside the heart itself. This is shown in Figure 11.

[0034] FIG. 11 is a highly schematic illustration of a cross section of the human rib cage and thoracic cavity.

[0035] Reference numeral 501 denotes ribs, 502 denotes skin, 503 denotes superficial fascia, 504 denotes intercostal muscles, 505 denotes parietal pleura, 506 denotes pleural cavity, 507 denotes visceral pleura, and 508 denotes lungs.

[0036] In this application, pressure sensor 510 is positioned between ribs 501 and in contact with parietal pleura 507 to measure pressure within pleural cavity 506. Electrical connections that may extend through tissue are not shown. The pressure sensor could also be positioned near the costal cartilage. Pressure sensor 510 functions in a similar manner to the pressure sensors described, but the electronics, sensor element, flexible membrane, and pressure transmission medium are located within a single housing.

[0037] The control unit 100 has input signals from the pressure sensors 231 and 232 and output signals to the pump actuation means 221, 222. The signals can be received and provided directly or via an interface, wirelessly or wired. The power supply 110 can be connected to or integrated into the control unit 100.

[0038] 3 shows a schematic diagram and functional blocks of an exemplary control unit 100. The control unit 100, according to this embodiment, comprises the following functional blocks: - a cardiac output control unit 101 for the right atrium, - a cardiac output control unit 102 for the left atrium; - a flow control block 103, and processing block 104, may include:

[0039] The main function of the control unit 100 is to set the heart rate and stroke volume for one side of the blood pump 200 (ie, the left pump and the right pump) to values ​​appropriate for the subject's current physiological state.

[0040] The control problem can be solved by breaking it down into multiple subproblems, including: - Flow Control: This determines the correct flow limit and desired atrial pressure, which is input and keeps cardiac output within the appropriate range. The desired atrial pressure can be overridden by allowing the operator of the prosthesis to manually set the desired atrial pressure. - Determining the cardiac output of one pump to control the atrial pressure: this is done individually, without interdependence between the pumps. - Determine heart rate: The heart rate and cardiac output of either pump (calculated in the previous step) will now provide the stroke volume of either pump.

[0041] Inputs to the function block include: - Right atrial pressure (RAP) measured by pressure sensor 231 in the right atrium to the (right) cardiac output control unit 101; and (Left) Left atrial pressure (LAP) measured by pressure sensor 232 in the left atrium to cardiac output control unit 102.

[0042] Determination of cardiac output for either pump can be performed using only atrial pressure as input.

[0043] As mentioned above, stroke volume can be calculated as stroke length multiplied by a constant and therefore may not match the actual pumped stroke volume, which implies that the reported flow rate is also only an approximation.

[0044] The function block flow control 103 determines the flow limit and the desired atrial pressure. The flow limit is the limit on the cardiac output of the right pump.

[0045] The purpose of flow control is to keep the flow rate of the right pump 202 low enough so that the left pump 203 never reaches its maximum flow rate. If this occurs, i.e., if the maximum flow rate is reached, the left pump will not be able to maintain the LAP in a restricted state, and there is a risk of pulmonary edema. By restricting the capacity of the right pump, the RAP may increase, which can be considered acceptable. Higher RAP also increases central venous pressure (CVP), which helps reduce venous return. The flow control block 103 receives output from the left cardiac output controller 102 and sets the desired RAP, desired LAP, and right cardiac output limits.

[0046] The actual flow rate of either half of the pump, i.e., 202 and 203, even given the same heart rate and stroke length, may depend on various factors such as outflow and inflow pressures, etc. Due to this, the flow restriction of the right pump (202) cannot be constant but must vary dynamically.

[0047] The following parameters can be considered:

[0048] Flow State: The flow state is a variable that depends on the flow rate of the left pump 203. Table 1 shows some example values ​​for different flow states and left pump flow rate in liters / minute: [Table 1]

[0049] Flow Limitation: The flow limit is the maximum allowable cardiac output of the right pump. The following rules can be applied: - When transitioning from a low or normal flow state to a high or very high flow state (see Table 1), the flow limit can be set equal to the current right cardiac output. This is the only state transition that directly results in a change in flow limit. - When in high flow conditions, the flow restriction is constant. - During very high flow conditions, flow restriction is reduced at a constant rate, but not below half of maximum cardiac output. - When in low flow or normal flow conditions: The flow limit is increased by a fixed percentage only if the output of the right pump is limited by the flow limit. - Flow restrictions can be reduced at a faster rate than they can be raised. - The limit should never be set higher than the maximum cardiac output. - The limit should never be set lower than half of the maximum cardiac output. Desired Atrial Pressure: In low flow conditions, the desired atrial pressure can be reduced to help improve venous return.

[0050] During low flow conditions, the desired atrial pressure can be determined by the left flow rate and can be set according to Table 2. Table 2 shows exemplary values ​​for desired left and right atrial pressure (mmHg) in relation to left flow rate (liters / minute): [Table 2]

[0051] If a low flow or extreme flow condition is detected, an alert can be generated.

[0052] As a result, the flow control block 103 receives output from the left cardiac output control unit 102 and sets the desired RAP, the desired LAP and the right cardiac output limit.

[0053] A cardiac output controller for the right pump 101 receives the RAP, the desired RAP, and a limit on the right cardiac output, and outputs the cardiac output for the right pump. A cardiac output controller for the left pump 102 receives the LAP and the desired LAP, and outputs the cardiac output for the left pump, as will be further described below.

[0054] Processing block 104 uses the right and left cardiac output levels to generate the correct stroke volume, heart rate and left stroke volume, which are provided by the controller to the pump actuation means.

[0055] In one embodiment, cardiac output is controlled by a proportional-integral-derivative (PID) controller that controls the derivative of cardiac output to keep the actual atrial pressure close to (preferably equal to) the desired atrial pressure, as shown schematically in Figure 4a.

[0056] Figure 4b shows a simplified PID controller diagram in which the derivatives and integrals are assumed to cancel each other. This eliminates the need for numerical differentiation, which can introduce noise. The PID controller continuously calculates an error value e(t) as the difference between the desired set point (SP) and the measured process variable (PV) and applies a correction based on the proportional (P), integral (I), and derivative (D) terms. As a result, SP is the desired atrial pressure, and PV is the cardiac output. The measured atrial pressure is subtracted from the desired atrial pressure to generate the error value e(t). The calculated error signal e(t) is truncated before being fed to the PID controller. This is done to prevent the controller from "overreacting" when a large difference exists between the desired and actual atrial pressure, which can occur transiently. The error is simply multiplied by the P, I, and D actions. The resulting "error control actions" are then summed together, integrated, and output as cardiac output (left or right).

[0057] In one embodiment, two pumps 202 and 203 can be connected to the systemic and pulmonary circulations, respectively. These can be controlled individually by setting a limit on the cardiac output of the pump connected to the pulmonary circulation. As a result, the cardiac output of the pump connected to the systemic circulation is obtained, and given the cardiac output of the pump connected to the systemic circulation, an appropriate limit on the cardiac output of the pump connected to the pulmonary circulation is obtained, and the limit is updated as the cardiac output of the pump connected to the systemic circulation changes.

[0058] In another embodiment, two pumps can be connected to the systemic and pulmonary circulations, each with a desired cardiac output. The appropriate stroke volumes and stroke rates of the two pumps are achieved; the desired cardiac output of either pump is obtained; and using the desired cardiac outputs of the two pumps, appropriate stroke rates are calculated to achieve the appropriate stroke volume of either pump, given the achieved stroke rates, such that the product of the stroke rate and either stroke volume equals either desired cardiac output.

[0059] The cardiac portion integrated in function block 104 has two inputs, left and right cardiac outputs Q left and Q right and heart rate (HR) and left and right stroke volume (SV) left and SV right The solution to the following simultaneous equations is required: Q left =HR*SV left Q right =HR*SV right

[0060] The heart rate procedure is limited by several factors: Maximum heart rate, Minimum heart rate, ·Maximum stroke volume, ·Minimum stroke volume, This means that there is a maximum and a minimum cardiac output.

[0061] Flow rate (Q left and Q right ) is small, the stroke volume is also small. Only when the flow rate is high will the stroke volume be set closer to its maximum.

[0062] Figure 5 shows a graph of heart rate (beats / min) and flow rate (liters / min). This shows the same height (same heart rate) and Q left and Q rightThis allows us to interpret the problem as obtaining two points of flow rate equal to

[0063] To simplify the terminology, Q big Q left and Q right is the maximum value of Q small is the minimum value.

[0064] As shown in Figure 5, the function f is small can be used to adjust the heart rate to reach a reasonable heart rate. The function f is given by way of example only; other functions may be used. The curve f(Q) in FIG. 5 may remain high even relatively far to the left. This means that during times of low flow (which may occur if the patient has little blood volume due to dehydration or bleeding, for example), the heart rate remains relatively high and the stroke volume remains low, which constitutes a form of tachycardia.

[0065] The heart rate HR is: HR=f(Q small ) Given the following: SV small =Q small / HR SV big =Q big / HR However, this is SV big >SV max Unless otherwise requested, the heart rate is set according to the maximum flow rate as follows: HR=Q big / SV max

[0066] The above values, heart rate and stroke volume, are truncated to stay within those limits.

[0067] The controller unit 100 converts the HR and SV values ​​into control signals and provides them directly or indirectly to the actuation means 221 and 222 .

[0068] The controller unit 100 can be an embedded or integrated microcomputer or electronic chip. The microcomputer can provide control signals to the pump actuation means to modify the pumping activity. If, for any reason, the microcomputer is not receiving any input information, the pump actuation means can continue at a constant level of activity.

[0069] In some situations, atrial pressure may become too low or intrathoracic pressure may increase (e.g., due to ventilatory support or other reasons). In the presence of such conditions, the atria and / or connected veins may collapse. This may completely or partially block the flow of blood into the pump (this may occur individually in any pump). This is known as a suction event.

[0070] The controller unit of the present invention can detect these events and act to prevent them.

[0071] The control unit detects an aspiration event by comparing the atrial pressure received from the pressure sensor, averaged over the course of one stroke, with the desired atrial pressure. If the average atrial pressure falls too far below the desired pressure (by a certain margin, which may be on the order of tens of mmHg), an aspiration event is detected.

[0072] To reduce noise and prevent false positives (detecting aspiration when there is none), the atrial pressure is averaged over the course of a cycle (e.g., on the order of 1 / 2 second to 1 second). When a aspiration event is detected by the control unit, there will be a period of inactivity (on the order of several seconds) during which no aspiration event is detected. This is to avoid detecting the same aspiration event multiple times.

[0073] The exact value of the atrial pressure detection margin and the length of the inactive period can be set to different values: the detection margin is on the order of tens of mmHg, and the inactive period is on the order of a few seconds.

[0074] Once a suction event is detected, a stepwise increase in the desired atrial pressure (on the order of one to several mmHg) can be implemented to prevent further suction events. The increase in the desired atrial pressure can be appropriately selected.

[0075] Upon detection of suction events, these can be reported to the device's user interface.

[0076] FIG. 6 shows one of the pumps 602 of an artificial heart (artificial object) according to one embodiment of the present invention, and includes a cylindrical housing 660, an upper chamber 609 inside the housing, an inlet flow path 610, an outlet flow path 613, a pump actuator housing 661, a pressure sensor housing 662, a pressure transmission medium pipe 663, and a sensor cover 664.

[0077] Figure 7a is a cross-sectional schematic view of an alternative pump 702. Figure 7b is an enlarged view of the circled section including the sensor 732.

[0078] The sensor structure 732 includes an opening or container 738 in the wall of the upper chamber 709 , a flexible membrane 733 , a pipe 763 , an electrical sensor 734 , a circuit board 735 , a pressure transmission medium 736 , and a mounting mechanism 737 .

[0079] The pressure-sensing sensor 734 may include a MEMS (microelectromechanical system) sensor disposed on an electrical circuit board 735 (PCB) along with corresponding electronics. The pipe 763 may include a metal cylinder or the like, attached to the electronic circuitry from one side with the pipe's lumen 7631 directly connected to and communicating with the MEMS sensor 734. The pressure-receiving portions of the pressure sensors are disposed in the left atrium 709 and the right pump atrium (or the upper half of each pump), as described above. The lumen 7631 of the pipe 763 is connected to a flexible membrane 733 in the atrial wall 709 of each respective pump, either directly or through a separate cylindrical structure (assembly tube), to extend the length of the lumen and make assembly of the pump with the pressure sensor possible and satisfactory. The flexible membrane 733 is disposed as part of the inner wall of the housing and part of a container 738 containing the pressure-transmitting medium 736. The wall of the assembly tube may be made of a hard plastic or metal material. The connection of the pipe lumen or the lumen of the assembly tube 763 to the membrane 733 of the atrial wall is sealed to prevent leakage, for example, using glue or a sealing washer. The pipe 763 and the container 738 can be filled with a biocompatible, implantable oil, such as medical-grade silicone oil, as the pressure transmission medium 736. The container 738 can be a part of the pipe 763 with the same or slightly different dimensions. There can be means assembled to the metal pipe lumen or the lumen of the assembly pipe to facilitate filling the lumen with oil, and there can be other means connected to the metal cylinder lumen or the lumen of the assembly tube to facilitate evacuation of air while filling the lumen with the pressure transmission medium.

[0080] In operation, when the membrane 733 of the atrial wall 760 is affected by the pressure of blood inside the atrium 709 of each pump, the flexible membrane 733 will either bulge when the pressure inside the atrium increases or deflect inward into the atrial cavity when the pressure decreases. In this way, the pressure transmission medium inside the reservoir and the lumen of the pipe 763 will compress or expand due to the influence of the pressure inside the atrium 709. The pressure is transmitted from the membrane inside the pipe 763 to the surface of the sensor 734. Thus, the sensing surface of the MEMS sensor 734 is affected by the pressure changes in the pressure transmission medium, and the MEMS sensor converts the pressure into a digital value, and an electrical signal representative of the pressure is generated and provided to the controller unit 100.

[0081] According to this exemplary embodiment, the structure of the sensor assembly conserves space and utilizes the available space within the pump housing. Therefore, the sensor structure described above is one example of a sensor, and other sensors may be used in the pump, for example, the sensor may be located entirely within the atrial wall. Pressure measurements may also be taken between the inlet opening and the end of the chamber forming the atrium.

[0082] While many of the illustrated embodiments above relate to pulsating or displacement pumps, the methods and control units of the present invention are equally applicable to systems with different types of pumps, such as centrifugal pumps. Centrifugal pumps can include one or more of canned, radial, side channel, regenerative turbine, axial, and mixed flow pump types. Displacement pumps can include one or more of dosing, axial screw, gear, multi-screw, piston-diaphragm, plunger-and-piston, rotary lobe, vacuum, and hose pump types.

[0083] 8 is a diagram of an exemplary controller unit 100 capable of implementing the methods described herein. Controller unit 100 may include a bus 110, a processor 120, a memory 130, a read-only memory (ROM) 140, a storage device 150, an input device 160, an output device 170, and a communication interface 180. Bus 110 enables communication between components of controller unit 100. Controller unit 100 may also include one or more power supplies (not shown). Those skilled in the art will understand that controller unit 100 may be configured in several other ways and may include other or different elements.

[0084] Processor 120 may include any type of processor or microprocessor that interprets and executes instructions. Processor 120 may also include logic that can decode media files and generate output to, for example, a speaker, a display, etc. Memory 130 may include random access memory (RAM) or another dynamic storage device that stores information and instructions for execution by processor 120. Memory 130 may also be used to store temporary variables or other intermediate information during execution of instructions by processor 120.

[0085] ROM 140 may include a conventional ROM device and / or another static storage device that stores static information and instructions for processor 120. Storage device 150 may include a magnetic disk, solid-state drive, or optical disk and its corresponding drive, and / or some other type of recording medium and its corresponding drive for storing information and instructions. Storage device 150 may also include a flash memory (e.g., an electrically erasable programmable read-only memory (EEPROM)) device for storing information and instructions.

[0086] The input device 160 may include one or more conventional mechanisms that allow a user to input information into the controller unit 100, such as a keyboard, keypad, directional pad, mouse, pen, voice recognition, touch screen, and / or biometric mechanisms. The output device 170 may include one or more conventional mechanisms that output information to a user, including a display, printer, one or more speakers, etc. The communication interface 180 may include any transceiver-like mechanism that allows the controller unit 100 to communicate with other devices and / or systems. For example, the communication interface 180 may include a modem or an Ethernet interface to a LAN. Alternatively, or in addition, the communication interface 180 may include other mechanisms for communicating over a network, such as a wireless network. For example, the communication interface may include a radio frequency (RF) transmitter and receiver and one or more antennas for transmitting and receiving RF data.

[0087] Consistent with the present invention, controller unit 100 provides a platform as described above. According to an exemplary embodiment, controller unit 100 can perform various processes in response to processor 120 executing sequences of instructions contained in memory 130. Such instructions can be read into memory 130 from another computer-readable medium, such as storage device 150, or from a separate device via communication interface 180. It should be understood that a computer-readable medium can include one or more memory devices or carrier waves. Execution of the sequences of instructions contained in memory 130 causes processor 120 to perform the operations described above. In alternative embodiments, hardwired circuitry can be used in place of or in combination with software instructions to implement aspects consistent with the present invention. Thus, the present invention is not limited to any specific combination of hardware circuitry and software.

[0088] Figures 9 and 10 show a perspective view of the pump of Figure 7a, which is a further extension of the prior art pump by the same applicant and is therefore part of an enhanced artificial heart. Figure 10 is a cross-sectional view of Figure 9. This embodiment includes a first blood-receiving member 702 of the artificial heart and a drive / actuator system 750 that drives / actuates the artificial heart to generate the pumping mechanism. The artificial heart mainly consists of two pumps 702, a left pump and a right pump (not shown). Each pump includes one blood-receiving member and one driver / actuator system. The blood-receiving members are easily assembled to the drive system by screws, glue, or a combination thereof. The blood-receiving members include an artificial atrium 709 and an artificial ventricle 712. Between the artificial atrium 709 and the artificial ventricle 712 is a connecting cylinder 720 equipped with one-way valves corresponding to the mitral valve on the left side of a natural heart and the tricuspid valve on the right side of a natural heart. The connecting cylinder can be made from a flexible blood compatible material such as polyurethane, silicone or any other blood compatible material.

[0089] Each atrium 709 consists primarily of a wall covering the atrium, which is largely composed of two layers: a rigid outer layer and a flexible inner layer. The rigid outer layer 7091 is preferably made of hard polyurethane, hard silicone, a biocompatible metal such as titanium or stainless steel, or any other biocompatible rigid material. The inner layer is a flexible, blood-compatible membrane 7092 made of polyurethane, silicone, or any other blood-compatible material. The inner flexible membrane is an extension of the flexible atrial membrane, which is an extension of a flexible membrane 7093 that covers the inner surface of the connecting AV cylinder 720, which is an extension of the flexible ventricular membrane 7121. In addition, there is an atrial protective flexible membrane 7094 that protects the atrial flexible membrane 7092. Additionally, there is an atrial mounting ring 721, which is rigid and made of rigid polyurethane, rigid silicone, biocompatible metal such as titanium or stainless steel, or any other biocompatible hard material. The atrial mounting ring is attached to the upper edge of the drive / actuation system by screws, glue, or a combination thereof. The atrium has an inlet opening 710 for blood flow into the atrium. There is also a pressure window 7641, which may be circular, oval, or any other shape, for example, with a diameter of at least 5-30 mm. The pressure window wall consists of only a flexible membrane without a rigid wall layer. This pressure window is part of the pressure sensor structure, as described above.

[0090] Each ventricle 712 primarily consists of a wall covering the ventricle, which is largely composed of two layers. The outer layer is rigid and made of rigid polyurethane, rigid silicone, a biocompatible metal such as titanium or stainless steel, or any other biocompatible rigid material. Additionally, there is an inner flexible blood-compatible membrane 7121 made of polyurethane, silicone, or any other blood-compatible material. The inner flexible membrane is an extension of the flexible ventricular membrane, which in turn is an extension of the flexible membrane 7093 that covers the inner surface of the connecting AV cylinder 720. Additionally, there is a ventricular protective flexible membrane 7095 that protects the flexible membrane 7121. Additionally, there is a ventricular mounting ring 722 that is rigid and made of rigid polyurethane, rigid silicone, a biocompatible metal such as titanium or stainless steel, or any other biocompatible rigid material. The ventricular mounting ring is assembled to the lower edge of the drive / actuation system by screws, glue, or a combination thereof.

[0091] The driver / actuator system 750 consists of a gearbox 753 and an electric motor 751 in a housing 762. The housing is made of a rigid plastic biocompatible material, such as PEEK or any other biocompatible plastic material, or a biocompatible metal. The housing encloses the gearbox and electric motor. The gearbox consists of multiple gears and a metal shaft. The electric motor is a brushless motor type or any other type of electric motor with or without an encoder. The motor 751 is disposed inside the pump housing between the outlet 713 and the pump. A spacer 760 can be disposed between the outlet 713 of the housing and the motor 751. The spacer 760 is made of a material with good heat transfer capacity so that when blood (or other liquid) flows through the outlet pipe (700), heat from the drive mechanism is transferred to the blood by the spacer 760 and the wall of the outlet 713, thereby transferring and reducing heat from the drive mechanism.

[0092] The center of the drive / actuation system has a cylindrical shape and surrounds the AV-cylinder 720, which is made of a rigid plastic biocompatible material such as rigid polyurethane, rigid silicone, or a biocompatible metal such as titanium, stainless steel, or any other biocompatible rigid material. The AV-cylinder is covered by a flexible membrane 7093, which is an extension of the flexible atrial membrane 7092 and the flexible ventricular membrane 7021. The AV-cylinder surrounds the valve 714. There are two racks 725 and 726, one on each side of the AV-cylinder 720. Each rack is articulated by gears in a gearbox 753 that are actuated upwards and downwards.

[0093] There may be wiring (not shown) connected to the drive / actuation system to provide power and control signals to the electric motor. Additionally, there may be a coating (not shown) that is a layer of biocompatible plastic material such as polyurethane or silicone that surrounds the entire drive / actuation system and allows the wiring to the electric motor and the pressure sensor pipes to pass through this coating.

[0094] It should be noted that the word "comprising" does not necessarily exclude the presence of other elements or steps than those listed, and the word "a" or "an" preceding an element does not exclude the presence of a plurality of such elements. Furthermore, it should be noted that any reference signs do not limit the scope of the claims, and that the invention may be implemented at least partly by means of both hardware and software, and that several "means," "units," or "devices" may be represented by the same item of hardware.

[0095] The embodiments described and illustrated above are given as examples only and should not be limitations on the present invention. Other solutions, uses, objects and functions within the scope of the present invention, as claimed in the following claims, should be apparent to those skilled in the art.

[0096] Various embodiments of the invention described herein are described in the general context of method steps or processes, which in one embodiment may be implemented by a computer program product including computer-executable instructions, such as program code, embodied in a computer-readable medium for execution by computers in a network environment. Computer-readable media may include removable and non-removable storage devices, including, but not limited to, read-only memory (ROM), random-access memory (RAM), compact discs (CDs), digital versatile discs (DVDs), solid-state drives, and the like. Generally, program modules may include routines, programs, objects, components, data structures, and the like that perform particular tasks or implement particular abstract data types. Computer-executable instructions, associated data structures, and program modules represent examples of program code for executing steps of the methods disclosed herein. The particular sequence of such executable instructions or associated data structures represents examples of corresponding acts for implementing the functions described in such steps or processes.

[0097] Software and web implementations of various embodiments of the present invention can be accomplished with standard programming techniques using rule-based logic and other logic to accomplish the various database lookup steps or processes, correlation steps or processes, comparison steps or processes, and decision steps or processes. It should be noted that as used herein and in the claims that follow, the terms "component" and "module" are intended to encompass implementations using one or more lines of software code, and / or hardware implementations, and / or apparatus for receiving manual input.

[0098] The foregoing description of embodiments of the present invention has been presented for purposes of illustration and description. The above description is not intended to be exhaustive or to limit the embodiments of the present invention to the precise form disclosed, as modifications and variations are possible in light of the above teachings or may be acquired from the practice of various embodiments of the present invention. The embodiments discussed herein have been chosen and described to explain the principles and properties of various embodiments of the present invention and their practical applications, and to enable those skilled in the art to utilize the present invention in various embodiments, with various modifications suited to the particular use contemplated. Features of the embodiments described herein may be combined in all possible combinations of methods, apparatus, modules, systems and computer program products.

Claims

1. A controller unit (100) configured to control an intracardiac prosthesis (200), the prosthesis comprising: first and second pump sections (202, 203, 602, 702), each pump section including a first and second chamber (209; 212); an inlet (210, 610, 710) connected to the first chamber (209) and an outlet (213, 613, 713) connected to the second chamber (212) in each pump section; a pump actuator (221, 222) in communication with each pump section, each pump actuator configured to induce fluid flow within a corresponding pump section; a pressure sensor (231; 232) arranged in each pump section between the inlet and the outlet and configured to measure a pressure value of the fluid flowing from the inlet to the outlet; The controller unit is further comprising a memory (130) and a processing unit (120); The processing unit: receiving a measured pressure value from a pressure sensor (231; 232) located in the first pump portion, the measured pressure value corresponding to the pressure of the fluid flowing between the inlet and the outlet; receiving a determined value of the pressure of the fluid flowing within the pump portion; controlling the actuator of each pump segment so that the measured pressure value approaches or equals the determined pressure value by controlling the pump stroke rate and pump stroke volume; wherein the product of pump speed and stroke volume, referred to as cardiac output, of at least one pump section provides the pump stroke volume of any pump section.

2. a flow control functional block configured to calculate a first flow restriction of the first chamber and the determined pressure as inputs for keeping cardiac output within range; a cardiac output determination functional block configured to determine a cardiac output of at least one pump portion to control the pressure of the first chamber; a pump stroke rate determination function block; The controller unit of claim 1 , comprising:

3. A controller unit as described in claim 1, wherein the controller unit is configured to maintain the flow rate from the first pump (202) at a first level so that the second pump (203) does not reach a maximum flow rate level.

4. The first and second pump portions are configured to be connected to the systemic circulation and the pulmonary circulation, respectively, and each pump portion is individually controlled by setting a limit on the cardiac output of the pump portion connected to the pulmonary circulation, and the controller unit: Obtain the cardiac output of the pump connected to the systemic circulation; setting a limit on cardiac output when the pump section is connected to the pulmonary circulation given the cardiac output of the pump section connected to the systemic circulation, the limit being updated as the cardiac output of the pump section connected to the systemic circulation changes; providing a control signal to a pump actuator; The controller unit of claim 1 further configured to:

5. The first and second pump portions are configured to connect to a systemic circulation and a pulmonary circulation, respectively, having a determined cardiac output, to achieve a stroke volume and a stroke rate of the two pump portions, and the controller unit: Achieving a determined cardiac output for any pump segment; Calculating the stroke rate using the determined cardiac outputs of the two pump segments; Given the stroke rate achieved, achieve a stroke volume for any pump portion such that the product of the stroke rate and the stroke volume is equal to the respective determined cardiac output; providing a control signal to the pump actuator of each pump section; The controller unit of claim 1 further configured to:

6. 10. The controller unit of claim 1, including a signal receiver operable to receive the signal, detect whether the pressure in the first chamber reaches a determined level or the thoracic pressure increases, and prevent the pressure in the first chamber or the thoracic pressure from increasing.

7. 7. The controller unit of claim 6, wherein the processing unit (120) of the controller unit is configured to compare the first chamber pressure value received from the pressure sensor, averaged over the course of one stroke, with the determined pressure value, and a suction event is detected if the average first chamber pressure value is lower than the determined pressure.

8. 8. The controller unit of claim 7, wherein the processing unit (120) of the controller unit is configured to average the pressure in the first chamber over the course of a cycle to reduce noise and prevent false detection of a suction event, and when a suction event is detected by the control unit, to provide an inactive period during which no suction event is detected.

9. 9. The controller unit of claim 8, wherein when a suction event is detected, the controller unit is configured to perform a step increase in the determined pressure of the first chamber to prevent further suction events and generate an alert.

10. The controller unit of claim 1 , wherein the fluid is blood.

11. A controller unit as described in claim 1, wherein the first chamber (209) corresponds to one of the right atrium or the left atrium.

12. The controller unit of claim 11 , wherein a pressure sensor is disposed in communication with the first chamber.

13. The controller unit of claim 1 , wherein the pressure sensor is configured to measure pressure at a position between an inlet opening and an end of a chamber forming the atrium.

14. The controller unit of claim 1 , wherein the sensor is configured to measure the pressure in a cavity of the pump that represents the thoracic cavity as a reference pressure.

15. An intracardiac prosthesis (200) comprising a controller unit according to claim 1.

16. A flexible membrane covering an opening; a pressure transmission medium; a pipe containing the pressure transmission medium; A pressure sensor; 10. A pressure sensor configured to be connected to the controller unit of claim 1, comprising:

17. An intracardiac prosthesis (200) including a pressure sensor according to claim 16 and configured to be connected to a controller unit according to claim 1.

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