Systems and methods for determining cardiac performance

Abstract

To determine cardiac performance.SOLUTION: Systems and methods described herein determine metrics of cardiac performance via a mechanical circulatory support device, and use the cardiac performance to calibrate, control and deliver mechanical circulatory support for the heart. The systems include a controller configured to operate the device, receive inputs indicative of device operating conditions and hemodynamic parameters, and determine vascular performance, including vascular resistance and compliance, and native cardiac output. The systems and methods operate by using the mechanical circulatory support device (e.g., a heart pump) to introduce controlled perturbations of the vascular system and, in response, determine heart parameters such as stroke volume, vascular resistance and compliance, and left ventricular end diastolic pressure, and ultimately determine native cardiac output.SELECTED DRAWING: Figure 3

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims priority to and benefit of U.S. Provisional Patent Application No. 62 / 687,133, entitled "METHODS AND SYSTEMS FOR IMPROVED ASSESSMENT OF VASCULAR AND CARDIAC STATE," filed June 19, 2018, U.S. Provisional Patent Application No. 62 / 863,136, entitled "SYSTEMS AND METHODS FOR SYSTEM IDENTIFICATION," filed June 18, 2019, and U.S. Provisional Patent Application No. 62 / 863,146, entitled "SYSTEMS AND METHODS FOR DETERMINING CARDIAC PERFORMANCE," filed June 18, 2019, the entire contents of which are incorporated herein by reference. [Background technology]

[0002] background Cardiovascular disease is a leading cause of illness, death, and healthcare burden worldwide. Various treatment modalities have been developed for cardiac health, ranging from pharmaceuticals to mechanical devices and transplants. Temporary cardiac assist devices, such as cardiac pump systems, provide hemodynamic support and promote cardiac recovery. Some cardiac pump systems are inserted percutaneously into the heart and can operate in parallel with the native heart to supplement cardiac output. Examples of such devices include the IMPELLA® family of devices (Abiomed, Inc., Danvers, Massachusetts). Such cardiac pump systems have sensors that detect blood pressure (or assess transmembrane pressure differentials) and can monitor motor current, using sensor and motor current readings to help identify the pump's location.

[0003] The cardiac support required by a given patient can vary from patient to patient. Cardiac output (CO) is the volumetric flow rate of blood pumped by the heart. Normal cardiac output is approximately 5 L / min in healthy adults but can vary depending on various factors, including a given patient's physical makeup. It is difficult for clinicians to quantitatively determine using known techniques what cardiac output a given heart provides, how much additional support a device should provide, when, and for how long. This determination can be particularly challenging for patients recovering from interventional or other cardiac procedures. Therefore, clinicians tend to rely on judgments and indirect estimates of cardiac function, such as measuring intracardiac or intravascular pressures using fluid-filled catheters. Quantifying cardiac output (CO) is particularly challenging. One approach uses a pulmonary artery catheter (PAC) to measure central venous pressure and pulmonary artery pressure in real time. PAC relies on estimating CO using Fick's law or bolus thermodilution via measurements of systemic oxygen consumption. However, the assumptions that must be made to arrive at CO metrics and the corresponding lack of precision may limit PAC's use in complex interventions and high-risk situations such as cardiogenic shock. PAC measurements do not consider dynamic changes in cardiac function and are not continuous, while they may not address nonlinear aspects of systemic ventriculovascular coupling. Summary of the Invention

[0004] overview The systems and methods described herein determine indicators of vascular and / or cardiac performance, such as CO2, via a mechanical circulatory assist device, such as an intravascular blood pump system, and use cardiac performance to calibrate, control, and provide mechanical circulatory support for the heart. The systems include a mechanical circulatory assist device and a controller configured to operate the device, receive inputs representing device operating conditions and hemodynamic parameters, and determine vascular performance, including vascular resistance and compliance, and intrinsic cardiac output. The systems and methods operate to introduce controlled perturbations in the vasculature by using the mechanical circulatory assist device (e.g., a heart pump), and in response, determine cardiac parameters such as stroke volume, vascular resistance and compliance, myocardial contractility, ventricular elastance, CO2, and left ventricular end-diastolic pressure, ultimately determining intrinsic cardiac output. These determined parameters are then used to calibrate and control further mechanical circulatory support for the heart. By determining the heart's intrinsic cardiac output, a therapy can be applied using mechanical circulatory assist (e.g., a blood pump). To implement therapy, the process control system activates or deactivates mechanical circulatory assist devices (eg, pumps) to provide and regulate the level of assist.

[0005] The system and method are configured with a time-variant nonlinear model of the vasculature, using device-arterial coupling to continuously determine systemic vascular resistance and compliance, thereby quantifying cardiac stroke volume. In some implementations, the system and method improves upon traditional linear approximations using a Windkessel model of the vasculature to provide dynamic changes in vascular response. In some embodiments, the system and method are configured as a cardiac output sensor that can directly determine the intrinsic cardiac output of a patient's heart.

[0006] In various applications, systems and methods are configured to "ping" the vasculature during a heartbeat using a mechanical circulatory assist system and then detect a response by the heart at one or more later periods or time points, e.g., at a later heartbeat. "Pinging" involves increasing or decreasing the output of the mechanical circulatory assist system (e.g., increasing or decreasing the pump speed of a cardiac pump system) for a short period of time, e.g., at a single heartbeat, thereby causing a spike in blood pressure and blood flow (e.g., aortic pressure and blood flow exiting the left ventricle). The ping causes a change in a hemodynamic parameter (e.g., aortic pressure) from baseline, and this change is detected and compared to the hemodynamic parameter at another time point (e.g., aortic pressure when the heart is not being pinged) to determine cardiac performance. Pings may involve changing the pump speed during a period (or time point) during a portion of a single heartbeat (e.g., a phase of the heartbeat) and comparing the hemodynamic parameters during this "change" time with those during a "normal" period or time point when no pings are applied (e.g., during a subsequent heartbeat).

[0007] Exemplary hemodynamic parameters include heart rate, blood pressure, arterial oxygen saturation, mixed venous saturation, central venous oxygen saturation, arterial blood pressure, mean arterial pressure, right arterial pressure, central venous pressure, right ventricular pressure, pulmonary artery pressure, mean pulmonary artery pressure, pulmonary artery occlusion pressure, left atrial pressure, aortic pressure, differential pressure, left ventricular end-diastolic pressure, stroke volume, stroke volume index, stroke volume variation, systemic vascular resistance, systemic vascular resistance index, pulmonary vascular resistance, pulmonary vascular resistance index, left ventricular These include ventricular stroke work, left ventricular stroke work index, right ventricular stroke work, right ventricular stroke work index, coronary artery perfusion pressure, right ventricular end-diastolic volume, right ventricular end-diastolic volume index, right ventricular end-systolic volume, right ventricular ejection fraction, arterial oxygen content, venous oxygen content, arterial-venous oxygen content difference, oxygen delivery, oxygen delivery index, oxygen consumption, oxygen consumption index, oxygen uptake rate, oxygen uptake index, total peripheral resistance, CO, cardiac index, and cardiac output (CPO).

[0008] Intrabeat ping (i.e., adjusting pump operation within a single heartbeat) allows for comparison of hemodynamic parameters across multiple heartbeats (either consecutively adjacent to each other or separated by other beats) while also minimizing noise (e.g., sympathetic response) that can occur when pump speed is changed over a relatively long period of time. As described above, in some embodiments, comparison of hemodynamic parameters is achieved by a control system processor programmed with a model of the vasculature, such as the two-element Windkessel model, which models and accounts for the changing, nonlinear interactions between pump flow and cardiac operation. The control system uses known terms (received as inputs) to approximate values ​​of pump flow and cardiac operation (e.g., aortic pressure), allowing such models to be easily validated and utilized in clinical applications. The systems and methods provide metrics representative of a patient's cardiac health, such as the resistance and compliance of the systemic vasculature, which allows for the determination of CO and other aspects of cardiac performance. In some applications, the systems and methods are deployed without the need for additional measurements or diagnostic catheters. The ability to continuously and accurately track changes in systemic vascular performance (e.g., resistance or compliance) and estimate cardiac stroke volume represents a significant advance over traditional measurements obtained from PAC or other diagnostics deployed in current clinical practice.

[0009] In some applications, the systems and methods described herein vary the pump speed of a cardiac pump system over a single heartbeat to detect its effect on vascular performance. This can be done by comparing changes in one or more hemodynamic parameters during a “normal” or “reference” heartbeat (e.g., a heartbeat during which the cardiac pump system is operating at a first pump speed) and during a “modulated” heartbeat (e.g., a heartbeat during which the cardiac pump system is operating at a pump speed different from the first pump speed for at least a portion of the heartbeat). The reference heartbeat may occur before or after the modulated heartbeat. By modulating or “pinging” the heartbeat, the systems and methods can capture and quantify differences in hemodynamic parameters between the normal and modulated heartbeats. These differences are then correlated to differences in blood flow, stroke volume, CO, or other useful indicators of vascular and / or cardiac performance. By varying the pump speed for a short period of time (e.g., a portion of a heartbeat), systemic resistance and compliance can be quantified at various pump speeds with greater real-time accuracy and without introducing additional noise into the system measurements.

[0010] In some implementations, a controller is provided and configured to perform any of the implementations, aspects, and methods described herein. For example, the controller may be an Automated Impella Controller (AIC) from Abiomed, Inc., or any other suitable controller programmed to perform the disclosed functions. In some implementations, the systems and methods use a mechanical circulatory assist device, such as a heart pump. An exemplary heart pump includes a catheter, a motor, a rotor operably coupled to the motor, a pump housing at least partially surrounding the rotor, one or more sensors, such as a differential pressure sensor, and a controller, such that actuation of the motor drives the rotor and pumps blood through the pump housing. For example, a heart pump system may include a blood pump with a cannula configured to be deployed within the heart and a motor positioned either inside or outside the heart and configured to drive the pump. The heart pump system may be an Impella 3.5 heart pump from Abiomed, Inc., connected to an AIC or any other suitable control system.

[0011] The systems and methods described herein vary the operation of a mechanical circulatory device within a heartbeat to compare one or more monitored hemodynamic parameters in that phase with the same parameters at different heartbeats, and calculate an index representing the patient's vascular performance. For example, a pump can be inserted into a blood vessel and operated at a first pump speed (or other output level) for a series of baseline heartbeats, including a first heartbeat. The heart is then "pinged" by increasing or decreasing the pump speed for a very short period at a second or target heartbeat or at a specific phase of the target heartbeat (e.g., located at or after the dicrotic notch of the target heartbeat). Aortic pressure or other hemodynamic parameters are measured at both the series of baseline heartbeats and the target heartbeat. In some applications, the hemodynamic parameters are measured at the same portion of the first and second heartbeats (e.g., during the systole of both beats, or at or after the dicrotic notch of both beats). The system then compares hemodynamic parameters, such as aortic pressure, determined during a series of baseline heartbeats (e.g., during the first heartbeat) with such hemodynamic parameters determined during a period of increased pump speed (e.g., during the second heartbeat), e.g., by using the same sensors, to calculate or characterize vascular resistance or compliance (which can be used to determine CO or modify pump operation to better treat the patient). The series of baseline heartbeats (e.g., including the first heartbeat) may occur after or before the ping. Such a method can be performed using the heart pump system 100 of FIG. 1, described below, or any other suitable pump.

[0012] To implement the systems and methods, a pump or other mechanical circulatory assist device is placed within a patient's vascular system (e.g., within the patient's heart) and is operable to alter the patient's hemodynamics. For example, operation of the device may increase the patient's aortic pressure by unloading the left ventricle or other means. In some implementations, the pump is an intravascular blood pump device placed within the patient's heart by percutaneous insertion. The pump may be a surgically implanted device, a left ventricular assist device, a counterpulsation device, an expandable heart pump, an extracorporeal membrane device, or any other suitable device. The pump may be appropriate because the patient is in cardiogenic shock or otherwise experiencing declining vascular health. The pump may be positioned across the aortic valve, with the blood inlet to the pump located within the left ventricle and the outlet from the pump located within the aorta. The pump may: CO=i h +i p (1) where CO is the total cardiac output and i h is the intrinsic cardiac output, and i p is the flow contributed by the pump. Such pumps can provide a life-saving benefit to patients in cardiogenic shock by increasing the flow of oxygenated blood from the heart to the coronary arteries and other areas of the vasculature.

[0013] Hemodynamic parameters are monitored during operation of the pump or other mechanical circulatory assist system (e.g., parameters can be monitored continuously during cardiac performance, and relevant data representative of the parameters at selected times during selected beats can be identified for use). Suitable hemodynamic parameters include parameters related to blood flow in organs and tissues of the body. Exemplary hemodynamic parameters are heart rate, blood pressure, arterial oxygen saturation, mixed venous saturation, central venous oxygen saturation, arterial blood pressure, mean arterial pressure, right arterial pressure, central venous pressure, right ventricular pressure, pulmonary artery pressure, mean pulmonary artery pressure, pulmonary artery occlusion pressure, left atrial pressure, aortic pressure, differential pressure, left ventricular end-diastolic pressure, stroke volume, stroke volume index, stroke volume variation, systemic vascular resistance, systemic vascular resistance index, pulmonary vascular resistance, pulmonary vascular resistance index, left ventricular vasoconstriction ... ventricular stroke work, left ventricular stroke work index, right ventricular stroke work, right ventricular stroke work index, coronary artery perfusion pressure, right ventricular end-diastolic volume, right ventricular end-diastolic volume index, right ventricular end-systolic volume, right ventricular ejection fraction, arterial oxygen content, venous oxygen content, arterial-venous oxygen content difference, oxygen delivery, oxygen delivery index, oxygen consumption, oxygen consumption index, oxygen uptake rate, oxygen uptake index, total peripheral resistance, CO, cardiac index, and cardiac output (CPO). In some implementations, differential pressure (P ) is used instead of aortic pressure for diastolic calculations. diff ) can be used (e.g., when the differential pressure is known and the aortic pressure is unknown). P diff is equal to the aortic pressure minus the left ventricular pressure (LVP). In many cases, the LVP is much smaller than the aortic pressure during diastole and does not change significantly during diastole compared to the aortic pressure. In these cases, P diff is close enough to aortic pressure (i.e., LVP can be ignored) to serve as a surrogate when AoP is unavailable. diff If is used, it may affect the accuracy of the results of certain calculations described herein.

[0014] Pump speed is the speed of operation of the pump and corresponds to the amount of blood flow produced by the operation of the pump. In some implementations, pump speed corresponds to the rotational speed of the rotor. For example, pump speed can be 10,000 RPM, 20,000 RPM, 30,000 RPM, 40,000 RPM, 50,000 RPM, 60,000 RPM, 70,000 RPM, 80,000 RPM, 90,000 RPM, 100,000 RPM, or any suitable speed. Pump speed can correspond to a power level or P level, as described above in connection with FIG. 1. For example, pump speed can be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, P-9, or any other suitable value. In some implementations, pump speed instead corresponds to the rate at which the pump's chamber fills with and expels blood. By monitoring hemodynamic parameters, the systems and methods described herein can detect changes in the hemodynamic parameters over time, which can be used to quantify cardiac performance.

[0015] In some implementations, a heartbeat detection method is provided by measuring hemodynamic parameters during cardiac performance (e.g., during multiple cardiac beats), identifying and predicting various cardiac phases and their sequential order, and then adjusting pump speed or otherwise timing cardiac pings based on the prediction of when subsequent cardiac phases will occur. The first (baseline) phase of a cardiac beat in a cardiac cycle may be identified as systole, diastole, or any other suitable phase or combination of phases of the heart occurring over a first period of time during cardiac performance. For example, the first period may be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable length of time for a given cardiac beat or other cycle.

[0016] In a second step, a second phase of the cardiac cycle is predicted or identified as the target phase in which the heart will experience the "ping." For example, the second phase can be the second heartbeat, a different systole, a different diastole, or any other suitable phase or combination of phases, provided that the second phase is selected to be the time when the "ping" will affect the heart. The system and method can predict when the second phase of the cardiac cycle will begin based on previously monitored hemodynamic parameters (e.g., aortic pressure measured in the first phase of the cardiac cycle). The second phase of the cardiac cycle is predicted to occur over a second period of time. For example, the second period of time can be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length of time. The second period of time can be set to correspond to the length of the second phase of the cardiac cycle for a particular patient. For example, if the second phase of the cardiac cycle is diastole, the second period of time can be set to the average diastolic period of the particular patient or to the predicted length of the next diastolic period. In some examples, the second period may be preset to a period shorter than the period of a heartbeat. The second phase is predicted based on the monitored hemodynamic parameters and the identified first phase of the cardiac cycle. A cardiac signal may be monitored, and based on the monitored signal, the system and method predicts when the next diastole or systole will occur. By predicting the timing of the cardiac cycle phase, the system and method can determine the timing of an increase or decrease ("ping") in the pump speed to begin (or affect) exactly when the second phase of the cardiac cycle begins. For example, the system and method may be configured to ping the heart with a brief increase in pump output (e.g., by unloading the heart with a higher pump speed) so that the resulting increase in blood flow coincides with the beginning of a preferred time point or period within the target heartbeat, such as at or shortly after the onset of the dicrotic notch in the subsequent heartbeat, mid-systole of the subsequent heartbeat, throughout diastole, or a predetermined portion of the subsequent heartbeat.

[0017] In certain implementations, the first phase of a cardiac cycle is the diastole of a first cardiac beat, and the second phase of the cardiac cycle is the diastole of an immediately following second cardiac beat. The first phase of a cardiac cycle can be the systole of a first cardiac beat, and the second phase of the cardiac cycle is the systole of a second cardiac beat. In some implementations, the second phase is during a cardiac beat that is several beats removed from the first cardiac beat, and in other implementations, the second phase is during a cardiac beat adjacent to the first cardiac beat.

[0018] After establishing the baseline and target heartbeat phase, the pump speed is changed to ping the heart, for example, by operating the pump at a second pump speed different from the first pump speed during a second heartbeat. The pump speed adjustment can be an increase or decrease in pump speed to ping the heartbeat. For example, the pump can be adjusted to temporarily increase the pump speed during one period of the heartbeat, such as the diastole of the heartbeat. Furthermore, the pump can be adjusted to return or otherwise decrease the speed to the baseline during the same period of the heartbeat or at some other time during the same heartbeat. Pinging can also be configured to occur in reverse by temporarily decreasing the pump speed from the baseline.

[0019] The heart is momentarily "pinged," for example, at a higher pump speed, by temporarily changing the pump speed and returning it to the baseline. In some implementations, the heart pump operates at a first baseline pump speed, then momentarily changes to a second, higher pump speed during systole or diastole (or other target period) of a subsequent heartbeat, and then immediately changes back to the first pump speed. In some implementations, the change in pump speed lasts for a time shorter than the length of the heartbeat, so that the pump returns to the baseline during the same heartbeat as the pinged heartbeat. For example, the entire ping may occur within the target heartbeat, such that the length of the ping is shorter than the duration of the target heartbeat. Changing the pump speed within a single heartbeat reduces the impact of noise on the collection of hemodynamic data during the first and second phases, improving accuracy.

[0020] The pump speed is adjusted to effect a pump speed change at a desired portion. For example, the speed change can be effected at systole, diastole, or both within a heartbeat. In some implementations, to adjust the pump speed, the controller sends a signal to the pump to change the pump speed before the start of the target phase in time to account for any time delay between sending the control signal and the change in pump speed. The pump must effect the actual increase or decrease in speed at a desired portion of the target heartbeat (e.g., diastole, at or after the dicrotic notch). The ping is timed so that the increase in pump speed occurs temporally during a known period of the heartbeat. For example, the start of the speed ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. The end of the speed ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. In some implementations, the ping is achieved by increasing or decreasing the pump speed for a set period of time. For example, the ping can be synchronized to the beginning of diastole so that it occurs during diastole. Alternatively, the ping can be synchronized to the end of diastole so that it occurs during systole of the next heartbeat. In other applications, the ping is synchronized to the beginning of systole, the end of systole, the peak of systolic blood pressure, or any other suitable time. The ping is configured to last for a set period of time. In some applications, the ping is set to last for a period corresponding to the length of a phase of the heartbeat. For example, the ping can be set to last approximately 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable time.

[0021] Pinging the heart or other vascular system by momentarily adjusting mechanical circulatory support (e.g., pump speed) imposes perturbations on the heart or other vascular system, thereby enabling the determination of additional indices of cardiac performance, including systemic resistance, compliance, and cardiac output. This determination can be made without introducing additional hardware (other than the pump providing hemodynamic support) into the patient's body (although such additional hardware can still be used, if desired). To make this determination, the hemodynamic waveforms (e.g., aortic pressure (or intraventricular pressure) waveforms) during normal and pinged heartbeats are compared via a nonlinear model, such as the Windkessel model described below. Changes in the pressure waveform (or other hemodynamic parameter) between the normal and pinged heartbeats are reflected in different values ​​in the model during two periods (the period of the baseline (normal) heartbeat and the period of the target (ping) heartbeat), which generates two model equations. Knowing the pressure waveforms during the two periods reduces the number of unknown variables between the two model equations, allowing the resistance and compliance to be calculated. The calculated resistance and compliance values, and the aortic pressure waveform are then used to calculate Equation (2):

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[0022] The adaptation of the above-described technique can be applied in various ways. In some implementations, hemodynamic parameters are monitored during the second phase of the second heartbeat. For example, the heart pump system can continuously monitor aortic pressure or any other hemodynamic parameter. In some implementations, the hemodynamic parameters monitored during the first phase are compared with the hemodynamic parameters monitored during the second phase. For example, a first blood volume pumped by the heart during the first phase and a second blood volume pumped by the heart during the second phase can be calculated. A numerical difference between the first blood volume and the second blood volume can be calculated to quantitatively compare the hemodynamic parameters during the first phase with those during the second phase. For example, the area under the curve (AUC) of the flow rate curve can represent the blood volume. The difference in AUC during the first phase and the second phase can indicate the difference in the blood volume pumped at the first pump speed and the second pump speed. Furthermore, comparing the hemodynamic parameters between the first and second phases can include evaluating the linearity of the change in the hemodynamic parameters. For example, aortic pressure may not change linearly between pump speeds, i.e., the change in aortic pressure from one pump speed to the next may not be a linear progression. Depending on how the change in aortic pressure progresses between pump speeds, the aortic pressure at different pump speeds can be predicted.

[0023] Indicators of cardiac performance can be calculated based on changes in hemodynamic parameters between the first and second phases of the heart. For example, indicators of cardiac performance can be determined from different pressure waveforms of the cardiac cycle during the first and second heartbeats. Such indicators can include systemic resistance, systemic compliance, CO2, CPO2, stroke volume, stroke work, ejection fraction, myocardial contractility, ventricular elastance, cardiac index, and predictive value of patient survival. Many indicators of cardiac performance are interrelated. For example, CO2 is determined based on the flow rate of blood through an intravascular pump placed within the patient's heart. Stroke volume is an indicator of left ventricular function and is expressed by the formula SV = CO2 / HR, where SV is stroke volume, CO2 is cardiac output, and HR is heart rate. Stroke work, which is the work performed by the ventricle to eject blood, can be calculated from stroke volume using the formula SW = SV * MAP, where SW is stroke work, SV is stroke volume, and MAP is mean arterial pressure. Cardiac work is calculated by multiplying stroke work by heart rate. CPO is a measure of cardiac function that represents the heart's pumping power and is expressed in watts. CPO is calculated using the following formula: CPO=MAP*CO / 451(3) where CPO is cardiac output, MAP is mean arterial pressure, CO is cardiac output, and 451 is a constant used to convert mmHg x L / min to watts. The ejection fraction can be calculated by dividing the stroke volume by the ventricular blood volume. Other parameters, such as chamber pressure, preload state, afterload state, cardiac recovery, flow loading state, variable volume loading state, and / or cardiac cycle flow state, can be calculated from or determined via these values. In some implementations, indices representing cardiac performance are calculated via a two-element Windkessel model of the vasculature (e.g., the Windkessel model of FIG. 5) to model dynamic and nonlinear cardiac and vascular interactions. Thus, this process employs a time-varying nonlinear model of the vasculature and uses the coupling between the intravascular blood pump device and the patient's hemodynamic function, a well-controlled analog of ventricular-vascular coupling, to continuously determine systemic vascular resistance and compliance and quantify cardiac stroke volume without the need for additional external measurements.

[0024] The operation of the pump can be adjusted based on the indicator representing cardiac performance. Adjustments to the pump operation can include increasing the pump speed, decreasing the pump speed, adjusting the pump placement, turning the pump off, or any other suitable adjustment. For example, if the indicator representing cardiac performance is stroke volume, the pump speed can be increased if the stroke volume is below a threshold, while the pump speed can be decreased if the stroke volume is above the threshold.

[0025] In some implementations, a CO sensor is provided for determining the cardiac output of a patient's native heart. The CO sensor may include one or more hardware, software, and firmware elements configured to perform the methods described herein. In some implementations, the CO sensor includes a mechanical circulatory assist device (e.g., an intravascular blood pump) with a pressure sensor and a processor configured to receive measurements from the pressure sensor and determine the native cardiac output using intrapulsatile pumping as described herein. The mechanical circulatory assist device can be configured to be at least partially disposed within the patient's heart. In some applications, the intravascular blood pump includes a cannula, a rotor configured to pump blood through the cannula, and a drive mechanism configured to provide power to rotate the rotor. In some implementations, the cannula is configured to extend across the aortic valve such that a distal end of the cannula is in the left ventricle and a proximal end of the cannula is in the aorta. For example, a heart pump system may be considered "in place" when the cannula is positioned across the aortic valve so that the blood inlet to the pump is in the left ventricle and the outlet from the pump is in the aorta. The drive mechanism may include an internal motor, a drive cable, a drive shaft, or any other suitable element, or combination thereof.

[0026] The CO sensor may include an elongated catheter body coupled to the cannula. The elongated catheter may include a drive cable, electrical wiring connecting the blood pump to a control system, any suitable element, or any combination thereof. In some implementations, the pump includes a pump housing and a motor housing, the motor housing coupled to the cannula at a distal end of the motor housing. A rotor may rotate within the pump housing to induce blood flow into the cannula.

[0027] The CO sensor may include a hemodynamic parameter sensor operably disposed on the blood pump (or proximal or distal to the blood pump) and configured to detect pressure within the blood vessel resulting at least in part from the pumping of blood therein. For example, the pressure sensor may be an optical sensor located on or near the pump housing or cannula. As another example, the pressure sensor may include a pressure measurement lumen configured to measure aortic pressure. A differential pressure sensor may also be used, where one side or surface of the differential pressure sensor may be exposed to aortic pressure and a second side or surface of the differential pressure sensor may be exposed to intraventricular pressure, and the differential pressure sensor may measure the difference between the aortic pressure and the intraventricular pressure.

[0028] The CO sensor includes a controller electrically connected to the pressure sensor and configured to detect a signal from the sensor representative of blood pressure. All or part of the controller may be in a controller unit separate / remote from the intravascular blood pump. In some implementations, the control system is internal to the intravascular blood pump.

[0029] The controller can be configured to calculate CO based on a nonlinear model that correlates CO with vascular resistance and compliance based on changes in hemodynamic values ​​as a result of cardiac pinging. For example, the nonlinear model can be the Windkessel model or a simplified Windkessel model used in conjunction with a cardiac pump system placed across the patient's aortic valve. The governing equation of this model is:

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[0030] In some implementations, P0 is

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[0031] In some aspects, mechanical circulatory support can be provided to a patient using a blood pump in accordance with the systems and methods described herein, including operating a blood pump within the patient's vasculature, determining the patient's cardiac CO using any of the systems and methods described herein, and adjusting the pumping speed of the blood pump based on the determined CO.

[0032] In some aspects, a mechanical circulatory assist system can include an intracardiac blood pump having a cannula configured to extend into a left ventricle of the heart and a pressure sensor configured to detect left ventricular end-diastolic pressure. The system can be configured to determine CO according to any of the methods described herein.

[0033] In some implementations, a pump is placed within a patient's heart. The pump can be introduced into a patient because they are in cardiogenic shock, undergoing a coronary intervention, experiencing a heart attack, or otherwise experiencing declining cardiac health. The pump can be placed across the aortic valve so that the blood inlet to the pump is located in the left ventricle and the outlet from the pump is located in the aorta. The pump contributes to the natural heart's operation so that CO2 from the heart equals natural CO2 plus pump output.

[0034] A first aortic pressure wave can be detected. The first aortic pressure wave reflects multiple heartbeats, each reflected beat including a dicrotic notch. The pressure waveform can be measured by a pressure sensor. In some implementations, the pressure sensor can be built into the pump. In some implementations, the pressure sensor is located external to the pump and receives a fluid or electrical signal. The pressure sensor can be in communication with a controller configured to control operation of the pump.

[0035] Hemodynamic support can be applied to the heart at a first pumping rate during a first beat of the plurality of beats. For example, the first pumping rate can be a first rotor speed, such as the P level described above. Hemodynamic support to the heart is adjusted during a second beat of the plurality of beats (e.g., during systole or after the dicrotic notch) by providing a second pumping rate to the heart during the second beat. The first pumping rate can be set to be different from the second pumping rate.

[0036] A second aortic pressure wave of the heart can be detected at the second heartbeat. The second aortic pressure wave can be compared to a portion of the first aortic pressure wave corresponding to the second heartbeat to detect changes in the second aortic pressure wave. In some examples, the second aortic pressure wave can be compared to the first aortic pressure wave by comparing the area under the curve (AUC) for a length of time represented by a portion of the second aortic pressure wave with the same length of time represented by a portion of the first aortic pressure wave. In some examples, the overall maximum and minimum values ​​of the first and second aortic pressure waves can be compared. The shapes or slopes of the first and second aortic pressure waves can be compared, i.e., through changes in the wave derivatives over time. In some implementations, the first aortic pressure wave is compared to the second aortic pressure wave via a nonlinear model (e.g., the Windkessel model described below). Different waveforms provide two sets of values ​​to the model, resulting in two different equations, one for each aortic pressure waveform. The change between the first and second aortic pressure waves can be used to determine the resistance and compliance of the systemic vasculature. Furthermore, comparing the hemodynamic parameters between the first and second aortic pressure waves can include assessing the linearity of the change in the aortic pressure wave between the first and second pump speeds. For example, aortic pressure may not change linearly between pump speeds, i.e., the change in aortic pressure from one pump speed to the next may not be a linear progression. Depending on how the change in aortic pressure progresses between pump speeds, the aortic pressure at different pump speeds can be predicted.

[0037] In some implementations, CO is determined based on a nonlinear transfer function relating CO to systemic resistance and compliance. In some implementations, the nonlinear transfer function includes a Windkessel model. In some implementations, the transfer function is further related to an aortic pressure waveform.

[0038] The system and method can compare a hemodynamic parameter at a first heartbeat with a hemodynamic parameter at a second heartbeat to calculate a change in the hemodynamic parameter between the first and second heartbeats. This change is caused, at least in part, by the difference between the first and second output levels of the mechanical circulatory assist device. For example, if the hemodynamic parameter is aortic pressure, increasing the output level of the device will increase the measured aortic pressure, and decreasing the output level will decrease the measured aortic pressure. This change in aortic pressure from the first pump output level to the second pump output level indicates the contribution of the mechanical circulatory assist device to the change in aortic pressure.

[0039] An index representing cardiac vascular and / or cardiac performance can be calculated based on changes in hemodynamic parameters between the first and second heartbeats. For example, the hemodynamic parameters at the first and second heartbeats can be compared via a nonlinear model, such as the Windkessel model described below. Because changes in hemodynamic parameters between a normal heartbeat and a pinged heartbeat are reflected in different values ​​in the model for two periods (one for the first heartbeat and one for the second heartbeat), cardiac performance can be determined using two model equations. In some implementations, the index indicating cardiac performance is cardiac output. To calculate cardiac output, vascular resistance and compliance can be determined based on changes in hemodynamic parameters between the first and second heartbeats, as described above. Knowing the hemodynamic parameter waveforms for the two periods can reduce the number of unknown variables between the two model equations, allowing the resistance and compliance (and ultimately, cardiac output) to be calculated.

[0040] In some implementations, the systems and methods described herein include modeling a patient's heartbeat and representing it as a series of sinusoids that a processor can use to construct one or more heartbeats representative of the patient's cardiac function. The constructed heartbeats are then used by the processor to adjust the pump speed. As described above, the blood pump operates at a first pump speed (or other operating parameter), then adjusts to a second pump speed (or other operating parameter) to ping the heart, and then quickly ramps down to the baseline first speed or parameter. A hemodynamic parameter (e.g., aortic pressure) is monitored during the pump's operation, including the ping period. The processor calculates an index representing the heart's cardiac performance based on (i) the first operating parameter (pump speed), (ii) the second operating parameter (e.g., pump speed during ping), and (iii) the hemodynamic parameter during the first and second periods, e.g., the first and second diastole periods. The index is used in a transfer function or set of equations, such as those described above in connection with the Windkessel model. A mathematical representation of the hemodynamic parameter is determined by the controller processor for the first and second diastole phases. For example, the mathematical representation may be a sum of sinusoids or other waveform function representing the hemodynamic parameter at a given pump speed.

[0041] Cardiac performance is then calculated by a processor from the sum of the sinusoids or other waveforms. The calculation may include decomposing a first waveform representing the hemodynamic parameters for a first diastole (when the pump is operating at a first pump speed) to determine a first set of sinusoids, and decomposing a second waveform representing the hemodynamic parameters for a second diastole (when the pump is operating at a second pump speed) to determine a second set of sinusoids. These decompositions may include applying a Fourier transform to the first waveform, the second waveform, or both. The set of sinusoids may include one or more summed sinusoids.

[0042] The blood flow in the aorta is due to the pump contribution (i p) plus the contribution of the original heart (i h ), so comparing the first set of sinusoids and the second set of sinusoids, the patient's cardiac contribution to the blood flow in the aorta (i h For example, if the hemodynamic parameter is aortic pressure, it can be determined as a sum of sinusoids resulting from a Fourier transform:

number

[0043] The decomposition of the hemodynamic parameter over time resulting from changes in the device's operating parameters into its constituent frequencies allows the hemodynamic parameter to be characterized using a complex mathematical expression or series of equations. In some implementations, the mathematical expression is an exponential equation based on a comparison of sinusoids. After the hemodynamic parameter waveform is characterized by the mathematical expression, cardiac parameters such as vascular resistance and compliance can be determined from the equation. For example, if the hemodynamic parameter waveform is

number

[0044] Heartbeats representing a patient's cardiac function can be simulated based on a comparison of sinusoidal curves representing the evolution of hemodynamic parameters resulting from changes in pump operating parameters. For example, a blood pump can be operated at a range of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), with each pump speed corresponding to a rotor rotation speed within the pump and a corresponding frequency (e.g., 100 Hz, 200 Hz, 300 Hz, 400 Hz, 500 Hz, 1000 Hz, 2000 Hz, 3000 Hz, etc.). Changing the pump speed (or frequency) changes the blood flow within the vasculature resulting from the pump's operation, thereby changing the value of the hemodynamic parameter. By gradually passing through multiple pump speeds (or operating parameters, such as blood flow, resulting from the pump) to identify corresponding changes in one or more hemodynamic parameters, forming hemodynamic waveforms, and decomposing the hemodynamic waveforms resulting from each pump speed, a relationship between pressure and flow during diastole can be established. The patient's overall cardiac function can then be mapped (as a function of the measured hemodynamic parameters) as a mathematical representation that can be used to simulate future cardiac function and inform the delivery and control of mechanical circulatory support to the patient. For example, the measured aortic pressure waveform of a recorded heartbeat can be constructed using the methods described below, and the CO for that heartbeat can be calculated.

[0045] As mentioned above, in some implementations, a brief change in pump speed can be applied to the pump within one heartbeat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heartbeat can be compared to the aortic pressure of a heartbeat without this brief speed change or impulse stimulus. The difference between the two (the aortic pressure of the altered heartbeat and the aortic pressure of the "normal" heartbeat) is called the aortic pressure impulse response: Δp(t)=p1(t)-p2(t) where P1(t) is the pressure waveform measured with impulse stimulation, P2(t) is the pressure waveform without impulse stimulation, and ΔP(t) is the impulse response of the aortic pressure. If this impulse stimulus is applied only during diastole, the difference in total cardiac blood flow for the two heartbeats is: Δi(t)=i1(t)-i2(t) where i(t) and i(t) are the pump flows for the heartbeats with and without impulse stimulation, respectively, and Δi(t) is the impulse response of cardiac blood flow. The relationship between aortic pressure and pump flow is then calculated in the frequency domain as follows:

number

number

[0046] [Figure 1] 1 illustrates an exemplary heart pump system inserted into a patient's blood vessel. [Figure 2] 1 illustrates a process for calculating an index representative of cardiac performance of a heart according to a particular implementation. [Figure 3] 3 shows a plot 300 of pressure versus time for a heart pump system according to certain implementations. [Figure 4] 1 shows plots of pressure, motor speed, and flow rate versus time according to certain implementations. [Figure 5] 1 illustrates a Windkessel model according to a specific implementation. [Figure 6] 1 illustrates a CO sensor coupled to a patient according to certain implementations. [Figure 7] 1 illustrates a process for determining CO according to a particular implementation. [Figure 8] 1 illustrates a process for determining a change in a hemodynamic parameter between two heartbeats according to a particular implementation. [Figure 9] 1 illustrates a process for determining CO according to a particular implementation. [Figure 10] 1 shows two illustrative graphs showing aortic pressure and cardiac blood flow over the same 10 second period according to certain embodiments. DETAILED DESCRIPTION OF THE INVENTION

[0047] Detailed Description Certain exemplary aspects will be described to provide an overall understanding of the systems, methods, and devices described herein. While the aspects and features described herein are specifically described for use in connection with percutaneous heart pump systems, it will be appreciated that the components and other features outlined below can be combined with one another in any suitable manner and can be adapted and applied to other types of cardiac therapy and heart pump systems, such as heart pump systems implanted using surgical incisions and intra-aortic pumps.

[0048] The systems, devices, and methods described herein enable assessment of organ function using an assist device located completely or partially within the organ. In particular, the systems, devices, and methods enable assessment of cardiac function using a cardiac pump system, such as a percutaneous ventricular assist device. For example, such devices can be used to treat cardiogenic shock.

[0049] Using a heart pump system to assess cardiac function can alert medical professionals to changes in cardiac function and allow them to adjust the degree / level of assistance provided by the assist device (i.e., the flow rate of blood pumped by the device) based on the needs of a particular patient. For example, the degree of assistance can be increased when the patient's cardiac function is deteriorating, or decreased when the patient's cardiac function recovers and returns to a baseline of normal cardiac function. This can allow the device to dynamically respond to changes in cardiac function to promote cardiac recovery and can allow the patient to be gradually weaned from therapy. Furthermore, assessment of cardiac function can indicate when it is appropriate to terminate use of the heart pump system. While some embodiments presented herein are directed to heart pump systems implanted across the aortic valve and located partially in the left ventricle, the concepts are applicable to devices within the heart, cardiovascular system, or other regions of the body.

[0050] Assessment of cardiac function can include determining cardiac parameters utilizing interactions between the heart and devices. Using the Windkessel model of the vasculature to improve upon traditional linear approximations and provide dynamic variations in vascular response, the systems and methods described herein introduce controlled perturbations of the vasculature with a heart pump system and, in response, calculate cardiac parameters such as stroke volume, vascular resistance and compliance, CO, and left ventricular end-diastolic pressure. In particular, the systems, devices, and methods described herein "ping" the heart using a mechanical circulatory assist system. "Pinging" involves increasing the pumping speed of the heart pump system for a period of time, for example, during a single heartbeat, to create a spike in aortic pressure and flow. During pinging, hemodynamic parameters change, which can be detected and compared to hemodynamic parameters at another time (i.e., when the heart pump system is not pinged) to calculate other hemodynamic parameters or measure cardiac performance.

[0051] Continuously measuring vascular and cardiac performance using the action of the heart pump system can provide additional clinical data useful for titrating appropriate device support. The systems and methods also provide for the use of device-arterial coupling to determine cardiac and vascular status, including determining native cardiac output. The mechanical circulatory assist systems presented herein are located within the heart and work in parallel with native ventricular function. This allows the systems to be sensitive enough to detect native ventricular function, unlike some more invasive devices. Thus, the systems, devices, and methods enable the use of mechanical circulatory assist systems not only as assist devices but also as diagnostic and prognostic tools. By hydraulically coupling to the heart, the heart pump system can function as a sensor to extract information about cardiac function. In some implementations, the heart pump system operates at a constant level (e.g., a constant rotor rotational speed) while the power delivered to the assist device is measured. In certain implementations, the rotor speed of the heart pump system can be varied (e.g., as a delta, step, or ramp function) to further investigate native cardiac function.

[0052] FIG. 1 illustrates an exemplary heart pump system inserted into a patient's blood vessel. As an example, a heart pump system compatible with the present disclosure is disclosed in U.S. Patent Application Publication No. 2018-0078159-A1, the contents of which are incorporated herein by reference in their entirety. Generally, any other heart pump system or other mechanical circulatory assist system (as well as sensors for acquiring physiological data from a patient) can be used with the present disclosure. In some implementations, the systems and methods described herein can use deployable pumps (e.g., the Heartmate PHP™ family of devices (Thoratec Corporation)) or left atrium-to-femoral artery bypass pumps (e.g., the TandemHeart family of devices (LivaNova, PLC)). In some implementations, the systems and methods described herein can use the IMPELLA® family of devices (Abiomed, Inc., Danvers, Massachusetts).

[0053] The heart pump system 100 may operate within the heart, partially within the heart, external to the heart, partially external to the heart, partially external to the vasculature, or at any other suitable location in the patient's vasculature. The heart pump system may be considered "in place" when the cannula 173 is positioned across the aortic valve such that the blood inlet to the pump (e.g., blood inlet 172) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170) is located in the aorta. The heart pump system 100 includes a heart pump 106 and a control system 104. All or part of the control system 104 may be in a controller unit separate / remote from the heart pump 106. In some implementations, the control system 104 is internal to the heart pump 106. The control system 104 and the heart pump 106 are not shown to scale. The pump system 100 includes an elongated catheter body 105, a motor housing 102, and a drive shaft from which pump elements are formed. Pump 100 includes a pump housing 134 and a motor housing 102, which rotates impeller blades on a drive shaft coupled to a cannula 173 at a distal end 111 of the motor housing 102 within the pump housing 134 to generate blood flow into the cannula 173 at a suction head 174. The suction head 174 provides a blood inlet 172 at a distal end portion 171 of the cannula 173. A blood flow 109 passes through the cannula 173 in a first direction 108 and exits the cannula 173 at one or more outlet openings 170 of the cannula 173.

[0054] Rotation of the drive shaft within the pump housing 134 rotates the pump elements within the bearing gap. A hemocompatible fluid is pumped through the elongate catheter 105, through the motor housing 102, and to the proximal end portion of the cannula 173, where the fluid lubricates the pump. The flow of the hemocompatible fluid has a second direction 122 through the bearing gap of the pump. After exiting the bearing gap, the hemocompatible fluid follows flow direction 123, mixing with the flowing blood and flowing with the blood into the aorta.

[0055] The heart pump 100 is inserted into the patient's blood vessel through a sheath 175. A pump housing 134 surrounds the rotor and internal bearing and may be sized for percutaneous insertion into the patient's blood vessel. In some implementations, the pump is advanced through the vascular system and past the aortic arch 164. While the pump is illustrated in the left ventricle, the pump may instead be positioned in the right heart to pump blood from the patient's inferior vena cava or right atrium through the right ventricle and into the pulmonary artery.

[0056] To stabilize the heart pump 100 in a blood vessel or ventricle, the distal end portion 171 of the cannula 173 includes a flexible protrusion 176 distal to the suction head 174. The flexible protrusion 176 is atraumatic and helps prevent the suction head 174 from approaching the vessel wall and potentially becoming stuck to it due to suction. Because the flexible protrusion 176 does not have suction, the flexible protrusion 176 extends the pump 100 mechanically, but not fluidically. In some implementations, the flexible protrusion can be formed as a pigtail. In some aspects, the pump does not necessarily include a flexible protrusion.

[0057] The elongate catheter 105 houses a connection 126 with a fluid supply line and an electrical connection cable. In addition, the connection 126 supplies the hemocompatible fluid from a fluid reservoir to the pump and is housed within the control system 104.

[0058] The control system 104 includes a controller 182 that controls the pump 106 by sending power to the motor and controlling the motor speed. The control system 104 includes circuitry to monitor the motor current for a drop in current indicating air in the line, monitor a change in the differential pressure signal, monitor flow position, monitor suction, or any other suitable measurement. In some implementations, the control system 104 includes a display screen for displaying measurements such as the differential pressure signal and motor current. The control system 104 can include a warning sound, warning light, or indicator to alert the operator to a sensor failure, a disconnected or broken connection 126, or a sudden change in the patient's health.

[0059] The motor 108 is configured to operate at a speed necessary to maintain the rotor at a set speed. As a result, the motor current drawn by the motor to maintain the rotor speed can be monitored and used to understand the underlying cardiac condition, as described further below. The control system 104 is configured to vary the pump speed within the cardiac cycle of the assisted heart to effect changes in blood flow through the pump, with the pump speed changes synchronized to the heartbeat with at least one event per cardiac cycle related to a predetermined event within the cardiac cycle; i.e., the systems, devices, and methods described herein use the heart pump system to "ping" the heartbeat. When a "ping" occurs, the pump speed of the heart pump system (or other mechanical circulatory assist device) is increased or decreased for a relatively short period of time, such as during a phase of the cardiac cycle, and then changed to a baseline or other speed. The pump speed can be increased for a period within a single cardiac cycle or for a period spanning multiple cardiac cycles.

[0060] Heart pumps can operate at various pump speeds, or P levels. P levels are performance levels of the heart pump system and are related to the flow control of the system. Higher P levels increase the flow rate, motor current, and revolutions per minute for the heart pump system; therefore, higher P levels correspond to higher flow rates and higher revolutions per minute for the heart pump system. For example, power level P-1 can correspond to a first rotor revolutions per minute (RPM) number, while power level P-2 can correspond to a second RPM number. In some examples, the pump operates at 10 different power levels ranging from P-0 to P-9. These P levels can correspond to 0 RPM to 100,000 RPM, or any suitable number. As shown in FIG. 3 and described below, changing the rotor speed changes cardiac CO2.

[0061] In some implementations, the pump speed is increased during systole, diastole, or both during a single heartbeat. The pings are timed so that the increase in pump speed occurs over a known period of the heartbeat. For example, the start of the rate ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. The end of the rate ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. In some implementations, the pump speed is increased or decreased for a set period of time. For example, the start of the ping can be synchronized with the start of diastole, the dicrotic notch or post-dicrotic notch, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. The ping may last for a set period of time. For example, the ping may last for 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable time.

[0062] The control system 104 includes a current sensor (not shown). The controller 182 supplies current to the motor 108 via connection 126, such as via one or more electrical wires. The current supplied to the motor 108 via connection 126 is measured by the current sensor. The load experienced by a mechanical pump motor corresponds to the force of the pressure head, or the difference between aortic and left ventricular pressure. The heart pump 106 experiences a nominal load during steady-state operation at a given pressure head; variations from this nominal load are the result of changes in external load conditions, such as the mechanics of left ventricular contraction. Changes in dynamic load conditions alter the motor current required to operate the pump rotor at a constant or substantially constant speed. As described above, the motor can operate at a speed required to maintain the rotor at a set speed, and monitoring the motor current drawn by the motor to maintain the rotor speed can be used to detect underlying cardiac conditions. Cardiac conditions can be accurately quantified and understood by simultaneously monitoring the pressure head during the cardiac cycle using pressure sensor 112. A cardiac parameter estimator 185 receives the current signal from the current sensor and the pressure signal from the pressure sensor 112. The cardiac parameter estimator 185 uses these current and pressure signals to determine cardiac function. The cardiac parameter estimator 185 can access a stored lookup table to obtain additional information to determine cardiac function based on the pressure and current signals. For example, the cardiac parameter estimator 185 can receive the aortic pressure from the pressure sensor 112 and use the lookup table to determine the delta pressure using the aortic pressure. The cardiac parameter estimator 185 can be software programmed into the controller 182 or can be separate hardware connected to the controller 182 by a wired or wireless connection. The cardiac parameter estimator 185 is configured to execute the algorithms described herein. For example, the cardiac parameter estimator 185 can be configured to estimate pump flow based on the current delivered to the pump and to determine intrinsic cardiac output according to the methods described herein.

[0063] Various implementations of the pressure sensor can be used. One example is an optical sensor or a differential sensor. The differential pressure sensor is a flexible membrane integrated into the cannula 172. One side of the sensor is exposed to blood pressure outside the cannula, and the other side is exposed to blood pressure inside the cannula. The sensor generates an electrical signal (differential pressure signal) proportional to the difference between the pressure outside and inside the cannula, which may be displayed by the heart pump system. When the heart pump system is correctly positioned across the aortic valve, the top (external surface) of the sensor is exposed to aortic pressure, and the bottom (internal surface) of the sensor is exposed to intraventricular pressure. Therefore, the differential pressure signal is approximately equal to the difference between the aortic pressure and the intraventricular pressure. Other sensors, such as optical sensors or fluid-filled columns, can also be used.

[0064] FIG. 2 illustrates a process 200 for determining cardiac performance of a heart. The process includes a series of steps related to modifying (pinging) the operation of a pump within a patient's heart to compare monitored hemodynamic parameters and calculate an index (e.g., CO2) representative of the heart's cardiac performance. For example, the process described below may increase the pump speed for a short period of time and compare the aortic pressure during the increased pump speed with the aortic pressure during normal pump operation to calculate or determine vascular resistance or compliance, which can be used to determine CO2 and / or modify pump operation to better treat the patient. Vascular resistance or compliance is determined via the Windkessel model (described below in FIG. 5 ) or other nonlinear time-dependent model by constructing a system of two equations, one for normal pump operation and one for high-speed pump operation, which can be solved using measured or estimated pressure and flow values ​​to calculate resistance and compliance values ​​for the systemic vasculature. Process 200 can be performed using cardiac pump system 100 of FIG. 1 or any other suitable mechanical circulatory assist system.

[0065] In step 202, a pump (e.g., pump 102 in FIG. 1 ) is placed within the patient's heart. In some implementations, the pump is an intravascular blood pump device placed within the patient's heart by percutaneous insertion. In some implementations, the pump may be a surgically implanted device, a left ventricular assist device, a counterpulsation device, an expandable heart pump, or any other suitable device. The pump may be introduced into a patient because the patient is in cardiogenic shock or otherwise experiencing declining health. The pump may be positioned across the aortic valve so that the blood inlet to the pump (e.g., blood inlet 172 in FIG. 1 ) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 in FIG. 1 ) is located in the aorta. The pump is CO=i h +i p (1) where CO is the total cardiac output and i h is the intrinsic cardiac output, and i p is the flow rate contributed by the pump.

[0066] In step 204, a hemodynamic parameter is monitored while the pump is operated at the first pump speed. The hemodynamic parameter may be any parameter related to blood flow in the body. For example, the hemodynamic parameter may be heart rate, blood pressure, arterial oxygen saturation, mixed venous saturation, central venous oxygen saturation, arterial blood pressure, mean arterial pressure, right arterial pressure, central venous pressure, right ventricular pressure, pulmonary artery pressure, mean pulmonary artery pressure, pulmonary artery occlusion pressure, left atrial pressure, aortic pressure, pressure difference, left ventricular end-diastolic pressure, stroke volume, stroke volume index, stroke volume variation, systemic vascular resistance, systemic vascular resistance index, pulmonary vascular resistance, pulmonary vascular resistance index, pulmonary vascular resistance, pulmonary vascular resistance index, left ventricular stroke work. , left ventricular stroke work index, right ventricular stroke work, right ventricular stroke work index, coronary artery perfusion pressure, right ventricular end-diastolic volume, right ventricular end-diastolic volume index, right ventricular end-systolic volume, right ventricular ejection fraction, arterial oxygen content, venous oxygen content, arterial-venous oxygen content difference, oxygen delivery, oxygen delivery index, oxygen consumption, oxygen consumption index, oxygen uptake rate, oxygen uptake index, total peripheral resistance, CO2, cardiac index, and CPO2. Pump speed is the rate at which the pump operates and corresponds to the amount of blood flow provided by operation of the pump. In some implementations, pump speed may correspond to the rotational speed of the rotor. For example, the pump speed may be 10,000 RPM, 20,000 RPM, 30,000 RPM, 40,000 RPM, 50,000 RPM, 60,000 RPM, 70,000 RPM, 80,000 RPM, 90,000 RPM, 100,000 RPM, or any other suitable speed. The pump speed may correspond to a power level or P level, as described above in connection with FIG. 1. For example, the pump speed may be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, P-9, or any other suitable value. In some implementations, the pump speed may instead correspond to the rate at which the pump's chamber fills with and expels blood.

[0067] In step 206, a first phase of a first heartbeat of the heart is identified. For example, the first phase may be systole, diastole, or any other suitable phase. The first phase of the first heartbeat is identified from the shape of a hemodynamic parameter waveform. For example, the hemodynamic parameter may be aortic pressure. Process 200 includes identifying a minimum in the aortic pressure waveform and determining a dicrotic notch from the minimum, where the onset of the dicrotic notch represents the onset of diastole. The first phase of the first heartbeat occurs during a first period. For example, the first period may be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable length of time.

[0068] In step 208, a second phase of the second heartbeat is predicted based on the monitored hemodynamic parameters. For example, the second phase may be systole, diastole, or any other suitable phase, and the second phase may be the same phase as the first phase (e.g., diastole). The second phase is predicted by monitoring the hemodynamic parameters over time and determining a pattern in the hemodynamic parameters to anticipate when the second phase of the cardiac cycle will begin. In some implementations, the prediction of the second phase may be further based on the identified first phase of the cardiac cycle. For example, if the first phase is the diastole of the first heartbeat and the second phase is the diastole of the second heartbeat immediately following the first heartbeat, the second phase can be predicted by determining the average length of the heartbeat and calculating the start of the second phase by adding the length of the heartbeat to the start time of the first phase. The second phase of the second heartbeat occurs over a second period of time. For example, the second period may be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length of time. By estimating when the second heartbeat (and subsequent heartbeats) will begin, the system can time the pump speed change so that its effect (increased flow from the left ventricle, increased aortic pressure, etc.) occurs during the desired second phase of the second heartbeat.

[0069] In one example, the first phase is the diastole of the first heartbeat and the second phase is the diastole of the second heartbeat. In another example, the first phase is the systole of the first heartbeat and the second phase is the systole of the second heartbeat. In one example, the first phase is the diastole of the first heartbeat and the second phase is the systole of the first heartbeat.

[0070] In step 210, the pump speed is changed so that the pump operates at a second pump speed during a second phase of the cardiac cycle, thereby "pinging" the heartbeat during this second phase. The pump speed may be increased or decreased. As shown in FIG. 3 and described below, the pump speed may be increased during diastole, i.e., the first phase may be the systole of the first cardiac cycle and the second phase may be the diastole of the first cardiac cycle. To accomplish the pump speed change, a controller (e.g., controller 104 of FIG. 1) may send a signal to the pump to change the pump speed before the start of the second phase, taking into account any time delay that may exist between sending the signal and the physical change in pump speed, so that the pump speed change occurs during the second phase. Changing the pump speed during a single cardiac cycle ensures that there is little or no noise or external factors affecting the collection of hemodynamic data during the first and second phases.

[0071] The pump speed changes after the "ping." In some implementations, the pump speed is changed back to the first pump speed after the second phase of the second heartbeat. For example, the heart pump can operate at the second pump speed only during systole or diastole, and then return to the first pump speed again during that period.

[0072] In step 212, hemodynamic parameters are monitored during the second phase of the second heartbeat while the ping occurs. For example, the heart pump system may continuously monitor aortic pressure or any other hemodynamic parameter. In step 214, the hemodynamic parameters monitored during the first phase are compared to the hemodynamic parameters monitored during the second phase. For example, a first blood volume pumped by the heart during the first phase and a second blood volume pumped by the heart during the second phase may be calculated. A numerical difference between the first and second blood volumes may be calculated to quantitatively compare the hemodynamic parameters during the first phase of the first heartbeat with those during the second phase of the second heartbeat.

[0073] In step 216, an index representing cardiac performance of the heart is calculated based on the change in hemodynamic parameters between the first and second phases. For example, the index representing cardiac performance may be systemic resistance, cardiac compliance, CO2, CPO2, stroke volume, stroke work, ejection fraction, myocardial contractility, ventricular elastance, cardiac index, or a prediction of patient survival. For example, the numerical difference between a first blood volume pumped by the heart in the first phase of the cardiac cycle and a second blood volume pumped by the heart in the second phase of the cardiac cycle may be calculated. The numerical difference in blood volume may be used to determine the stroke volume of an individual heartbeat or the average cardiac output (CO2) over a desired period of time. Many indexes representing cardiac performance are interrelated. For example, CO2 is determined based on the flow rate of blood passing through the pump. Stroke volume is an index of left ventricular function, and its formula is SV=CO2 / HR, where SV is stroke volume, CO2 is cardiac output, and HR is heart rate. Stroke work is the work done by the ventricle to eject blood and can be calculated from stroke volume according to the formula SW = SV * MAP, where SW is stroke work, SV is stroke volume, and MAP is mean arterial pressure. Cardiac work is calculated by multiplying stroke work by heart rate. CPO is a measure of cardiac function that represents the heart's pumping ability and is measured in watts. CPO is calculated using the formula CPO = mAoP * CO2 / 451, where CPO is cardiac output, mAoP is mean arterial pressure, CO2 is cardiac output, and 451 is a constant used to convert mmHg x L / min to watts. Ejection fraction can be calculated by dividing the stroke volume by the ventricular blood volume. Other parameters, such as chamber pressure, preload, afterload, cardiac recovery, flow loading, variable volume loading, and / or cardiac cycle flow, can be calculated from or determined via these values. In some implementations, indices representing cardiac performance of the heart are calculated via a two-element Windkessel model of the vascular system (e.g., the Windkessel model of Figure 5) to model dynamic and nonlinear cardiac and vascular interactions.Thus, this process employs a time-varying nonlinear model of the vasculature and leverages device-arterial coupling, a well-controlled analog of ventricle-vascular coupling, to continuously determine systemic vascular resistance and compliance and quantify cardiac stroke volume without the need for additional external measurements.

[0074] In optional step 218, the operation of the pump is adjusted based on the index representative of cardiac performance. In some implementations, the pump speed is increased or decreased based on the index representative of cardiac performance.

[0075] FIG. 3 illustrates a pressure versus time plot 300 of a heart pump system according to a specific implementation. The y-axis of plot 300 represents aortic pressure (mmHg), while the x-axis represents time as a percentage of heartbeat length. In particular, plot 300 illustrates the effect of pinging on aortic pressure. t1 represents the time of the first heartbeat, and t2 represents the time of the second heartbeat after the first heartbeat. Times t1 and t2 occur when the heart pump system is at least partially disposed within a patient's heart. Point 310 represents the peak systolic pressure of the first heartbeat, and point 320 represents the peak systolic pressure of the second heartbeat. Point 312 represents the dicrotic notch of the first heartbeat, and point 322 represents the dicrotic notch of the second heartbeat. Diastolic periods t3 and t4 represent the diastolic periods of the first and second heartbeats, respectively. At time t1, the pump operates at a first pump speed. At time t4, the pump operates at a second pump speed that is faster than the first pump speed.

[0076] At higher pump speeds, the measured aortic pressure and total flow rate are greater compared to lower pump speeds. Thus, during diastole t4, when the pump is operating at a second pump speed higher than the first pump speed, the aortic pressure is higher than during diastole t3, when the pump is operating at the first pump speed. The difference in aortic pressure between diastole t3 and t4 is illustrated by the shaded region 324. This difference correlates to an increase in flow rate and CO during the same period t4.

[0077] FIG. 4 shows plots of pressure, motor speed, and flow rate versus time. The y-axis of pressure plot 410 represents aortic pressure (mmHg), the y-axis of motor speed plot 420 represents motor speed in terms of P level, and the y-axis of flow rate plot 530 represents flow rate (mL / s). The x-axis of plots 410, 420, and 430 represents time as a percentage of heartbeat length. For all three plots, t1 represents the time of the first heartbeat, and t2 represents the time of the second heartbeat after the first heartbeat. Times t1 and t2 occur when the cardiac pump system is at least partially disposed within the patient's heart. At time 440, the second heartbeat begins. At time 450, the diastolic phase t3 of the second heartbeat begins. At time 460, the second heartbeat ends.

[0078] Pressure plot 410 is similar to plot 300 described above. Point 410 represents the dicrotic notch of the first heartbeat, point 414 represents the dicrotic notch of the second heartbeat, and point 416 represents the start of the systolic upstroke of the second heartbeat. At time 450, which corresponds to point 410 (the dicrotic notch of the first heartbeat) on plot 410, the pump speed is increased, as shown in motor speed plot 420. During time period t1, the pump operates at pump speed P-4. During diastole t3, the pump operates at pump speed P-6. There may be a time delay between when the controller sends a signal to the pump to change the pump speed and when the pump speed increases. As can be seen in pressure plot 410 and flow plot 430, as the pump speed increases to P-6 during time period t3, both flow and pressure increase.

[0079] 5 shows a Windkessel model 500. The Windkessel model 500 includes a current source 510, a current source 520, a resistance 530, and a compliance 540. The governing equation of the model is:

number

number

number

[0080] In this way, vascular status can be determined through analysis of the aortic pressure waveform measured by the heart pump system by measuring the difference in aortic pressure caused by changes in heart pump speed, based on the assumption that vascular status remains stable during this period. Systemic vascular resistance is determined using the above equation at two different Impella operating points and the difference in estimated Impella flow. These vascular status values ​​can then be used with the measured aortic pressure in the general equation above to calculate flow from the heart, thereby determining cardiac performance. The pulsatile ejection component of the flow waveform is numerically integrated over the ejection phase of the cardiac cycle to estimate cardiac output or CO.

[0081] FIG. 6 illustrates a CO sensor 610 coupled to a patient 600, the CO sensor configured to determine intrinsic cardiac output. The CO sensor 610 may include various hardware elements configured to perform the methods described herein. In some implementations, the CO sensor includes an intravascular blood pump (e.g., pump 202 of FIG. 1 ) and a controller for operating the pump, receiving inputs representing the pump's operating state and intravascular pressure, and determining intrinsic cardiac output. The intravascular blood pump can be configured to be disposed at least partially within the patient's heart. In some implementations, the intravascular blood pump includes a cannula, a rotor configured to rotate within the blood vessel and pump blood through the cannula, and a drive mechanism configured to provide power to rotate the rotor. In some implementations, the cannula can be configured to extend across the aortic valve such that a distal end of the cannula is in the left ventricle and a proximal end of the cannula is in the aorta. For example, a heart pump system may be considered "in place" when the cannula is positioned across the aortic valve so that the blood inlet to the pump is located in the left ventricle and the outlet from the pump is located in the aorta. The drive mechanism may include an internal motor, a drive cable, a drive shaft, or any other suitable element, or combination thereof.

[0082] In some implementations, the CO sensor 610 includes an elongated catheter body coupled to a cannula. The elongated catheter may include a drive cable, electrical wiring connecting the blood pump to a control system, any suitable element, or any combination thereof. In some implementations, the blood pump includes a pump housing and a motor housing, the motor housing coupled to the cannula at a distal end of the motor housing. A rotor can rotate within the pump housing to induce blood flow into the cannula.

[0083] The CO sensor 610 includes a pressure sensor configured to detect pressure within a blood vessel resulting at least in part from the pumping of blood within the blood vessel. For example, the pressure sensor may be an optical pressure sensor that is part of a blood pump, or a differential pressure sensor may be used. One side or surface of the differential pressure sensor may be exposed to aortic pressure, and a second side or surface of the differential pressure sensor may be exposed to intraventricular pressure, and the differential pressure sensor may measure the difference between the aortic pressure and the intraventricular pressure. As another example, the pressure sensor 612 may include a pressure measurement lumen configured to measure aortic pressure.

[0084] The CO sensor 610 includes a controller 614. The controller 614 is coupled to a pressure sensor 612. The controller 614 can be directly or indirectly coupled to the pressure sensor 612. For example, the controller 614 can be connected to the pressure sensor 612 via electrical wiring, a wireless signal, or any other suitable means. The controller 614 is configured to detect a signal from the pressure sensor that is indicative of blood pressure. All or part of the controller 614 can be in a controller unit separate / remote from the intravascular blood pump. In some implementations, the control system is internal to the intravascular blood pump.

[0085] In some implementations, the controller 614 is configured to calculate CO based on a nonlinear model that correlates CO to vascular resistance and vascular compliance. For example, the nonlinear model can be the Windkessel model, as described above with respect to FIG. 5.

[0086] FIG. 7 illustrates a process 700 for determining CO. Process 700 can be performed using the heart pump system 100 of FIG. 1 or any other suitable pump. In some implementations, the pump is an intravascular blood pump device that is placed within a patient's heart by percutaneous insertion. The pump may be introduced into a patient because the patient is in cardiogenic shock or otherwise experiencing declining health. The pump may be positioned across the aortic valve so that the blood inlet to the pump (e.g., blood inlet 172 of FIG. 1) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 of FIG. 1) is located in the aorta. The pump may: CO=i h +i p (1) where CO is the total cardiac output and i h is the intrinsic cardiac output, and i p is the flow rate contributed by the pump.

[0087] In step 702, a first aortic pressure wave is detected. The first aortic pressure wave reflects multiple heartbeats, each reflected beat including a dicrotic notch. The pressure waveform can be measured by a pressure sensor. In some implementations, the pressure sensor can be built into the pump. In some implementations, the pressure sensor can be located external to the pump. The pressure sensor can be in communication with a controller configured to control operation of the pump.

[0088] In step 704, hemodynamic support is applied to the heart at a first pumping rate during a first beat of the plurality of beats. For example, the first pumping rate may be a first rotor speed, such as the P level described above. In step 706, hemodynamic support to the heart is adjusted during a second beat of the plurality of beats by providing a second pumping rate to the heart after the dicrotic notch during the second beat. The first pumping rate is different from the second pumping rate.

[0089] A second aortic pressure wave of the heart is detected at the second beat in step 708. The second aortic pressure wave is compared to the portion of the first aortic pressure wave corresponding to the second beat in step 710 to detect a change in the second aortic pressure wave. The change between the first and second aortic pressure waves can be used to determine the resistance and compliance of the systemic vasculature.

[0090] In step 712, CO is determined based on a nonlinear transfer function relating CO to systemic resistance and compliance. The transfer function may be further related to the aortic pressure waveform. In some implementations, the nonlinear transfer function includes a Windkessel model, such as that described above in connection with FIG. 5.

[0091] FIG. 8 shows a process 800 for determining a change in a hemodynamic parameter between two heartbeats. In step 802, a mechanical circulatory assist device is placed within a patient's vasculature. In some implementations, the device is an intravascular blood pump device placed within the patient's heart by percutaneous insertion. The device may be introduced into a patient because the patient is in cardiogenic shock or otherwise experiencing declining vascular health. The device may be a left- or right-heart device. In some implementations, the device is positioned across the aortic valve so that the blood inlet to the device (e.g., blood inlet 172 in FIG. 1 ) is located in the left ventricle and the outlet from the device (e.g., outlet opening 170 in FIG. 1 ) is located in the aorta.

[0092] The device is operable to alter hemodynamic parameters in the patient. For example, operation of the device can affect the patient's aortic pressure by pumping blood from the left ventricle to the aorta. The device operates while the heart is beating at a first power level. The first power level corresponds to a first blood flow rate due to the contribution of the mechanical circulatory assist device to the patient's natural blood flow when the device operates at the first power level. For example, the first power level can be associated with a first motor speed, such as the P level described above.

[0093] The device operates at a first output level for a period that includes the period of a first heartbeat, and the patient's hemodynamic parameters are monitored during device operation, with the results of the monitoring determined as a function of time within each heartbeat and stored in the device's memory (or other data storage device). As described above, the hemodynamic parameters may be any parameters related to blood flow in a bodily organ or tissue. In step 804, the hemodynamic parameters are detected during the first heartbeat and coincide in time with the first heartbeat. Measurements of the device output level and hemodynamic parameters during that heartbeat (or any other time at the first output level) coincide with events in the cardiac cycle (systole, diastole, dicrotic notch, etc.). As a result, the hemodynamic parameters and device output levels can be correlated to cardiac cycle events at various times during the heartbeat. For example, it can be readily detected that a pump operating at a first output level has a first measured hemodynamic parameter (e.g., aortic pressure) at or after the dicrotic notch of the first heartbeat. In some implementations, the hemodynamic parameter is aortic pressure, and the mechanical circulatory assist device includes a pressure sensor configured to detect the aortic pressure. In some applications, the pressure sensor is included on a cannula extending partially into the patient's left ventricle.

[0094] In step 806, the device operates to output a second power level during a second period, which includes one or more periods within the second heartbeat. The second power level (delivered during the second heartbeat) may be greater than or less than the first power level (delivered during the first heartbeat). For example, the second power level may be associated with a second motor speed or P level greater than or less than the first motor speed, and may be delivered during the second heartbeat at the same phase point (e.g., at or after the dicrotic notch) as the first power level.

[0095] In step 808, a hemodynamic parameter is sensed during the second heartbeat (during the second power level) at or near the same time in the cardiac phase of the second heartbeat as the first heartbeat. The hemodynamic parameter may be measured throughout the first or second heartbeat, or for a portion of each beat. For example, the hemodynamic parameter may be measured during systole or diastole of the second heartbeat, or at the dicrotic notch.

[0096] In step 810, the hemodynamic parameter measured at the first heartbeat is compared to the hemodynamic parameter measured at the second heartbeat. These two measurements are taken at two different heartbeats but at approximately the same point in the cardiac cycle. The difference in the hemodynamic measurement occurs due to the change in pump speed between the first and second heartbeats. For example, if the hemodynamic parameter is aortic pressure, increasing the output level will increase the measured aortic pressure, and decreasing the output level will decrease the measured aortic pressure. This change in aortic pressure from the first output level to the second output level correlates to the contribution of the mechanical circulatory assist device to the change in total cardiac output.

[0097] FIG. 9 illustrates a process 900 for determining cardiac output through a cardiac "ping" process. Process 900 can be performed using the cardiac pump system 100 of FIG. 1 or any other suitable pump. The pump is placed into the patient's heart by percutaneous insertion. The patient may be in cardiogenic shock or otherwise experiencing compromised vascular health. The pump may be a left- or right-heart device. The pump is positioned across the aortic valve so that the blood inlet to the pump (e.g., blood inlet 172 of FIG. 1) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 of FIG. 1) is located in the aorta. The pump: CO=i h +i p (1) where CO is the total cardiac output and i h is the intrinsic cardiac output, and i p is the flow rate contributed by the pump.

[0098] In step 902, the pump operates at a first pump speed for a first period of time that includes the period of a first heartbeat. In step 904, a hemodynamic parameter is monitored during operation of the cardiac pump at the first pump speed during the first diastole of the first heartbeat. The hemodynamic parameter relates to blood flow through the body. The pump speed is the speed of operation of the pump and corresponds to the amount of blood flow provided by the operation of the pump. In some implementations, the pump speed corresponds to the rotational speed of the pump rotor. For example, the pump speed may be 10,000 RPM, 20,000 RPM, 30,000 RPM, 40,000 RPM, 50,000 RPM, 60,000 RPM, 70,000 RPM, 80,000 RPM, 90,000 RPM, 100,000 RPM, or any suitable speed or greater. The pump speed may correspond to a power level or P level, as described above in connection with FIG. 1. For example, the pump speed may be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, P-9, or any other suitable value. In some implementations, the pump speed corresponds to the rate at which the pump's chamber fills and expels blood. By monitoring hemodynamic parameters, the systems and methods described herein can identify changes in the hemodynamic parameters over time, including the phases of the first and second heartbeats. Such comparisons can be used to quantify cardiac performance (e.g., via CO), as discussed more fully herein.

[0099] In step 906, a first operating parameter of the intravascular blood pump during diastole is determined. For example, the operating parameter may be a current supplied to the pump, a rate of blood flow provided by the pump, or a placement of the pump in the patient's vasculature. Specifically, determining the first operating parameter may include determining a first blood flow rate provided by the blood pump during diastole. This first operating parameter and the measured hemodynamic parameter may be identified at a particular point in a cardiac cycle of a first heartbeat. The flow rate from the pump is estimated based on a motor current supplied to a blood pump motor to maintain the pump speed.

[0100] For a given intravascular blood pump system, the flow output i p can be determined by the pump's speed (revolutions per minute, or RPM) and the motor current supplied to the pump to maintain operation at that pump speed. This mathematical calculation from pump speed and motor current to flow rate can be accomplished by setting up a lookup table where the pump speed and motor current are indexes into the table and the flow rate values ​​in the table are pre-populated by bench testing. Another approach is to pre-determine flow rates for some of the possible combinations of pump speed and motor current. For example, if flow rate i1 represents the flow rate at a pump speed of 40,000 RPM and a motor current of 500 mA, and flow rate i2 represents the flow rate at a pump speed of 40,000 RPM and a motor current of 510 mA, then flow rate i3 at a pump speed of 40,000 RPM and a motor current of 505 mA can be calculated by taking the average of i1 and i2.

[0101] In step 908, the first pump speed is changed to a second pump speed so that operation of the heart pump produces a second output level during a second diastole of the second heartbeat. The second pump speed may be greater or less than the first pump speed. In some implementations, the timing of the pump speed increase is determined so that the pump speed increase occurs during a predicted period of the heartbeat. For example, the start of the speed increase can be synchronized with the start of diastole to account for the time delay between sending an instruction to the pump to change speed and the physical occurrence of the speed change. The end of the speed increase can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other suitable time. In some implementations, the system is configured to increase or decrease the pump speed for a set period of time. For example, the speed change may last approximately 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable time. The second heartbeat is different from the first heartbeat. The hemodynamic parameter is measured during the second heartbeat, and this may occur at the same time in the beat (such as at the dicrotic notch) as the first hemodynamic parameter was measured in the first heartbeat. In some implementations, the detection and measurement is applied to the second heartbeat, which occurs after the first heartbeat.

[0102] A hemodynamic parameter is monitored during a second diastole of the second heartbeat, for example, at the diastolic dicrotic notch, at step 910. A second operating parameter of the intravascular blood pump during the second diastole is determined at step 912. Determining the second operating parameter may include determining a second blood flow rate (or a second level of the motor operating parameter) provided by the blood pump during the second diastole.

[0103] In step 914, an index representative of the cardiac performance of the heart is calculated. The index is based on (i) the first operating parameter, (ii) the second operating parameter, and (iii) the hemodynamic parameter in the first diastole and the second diastole (e.g., at the dicrotic notch during both periods). The index can be used in a transfer function or set of equations such as those described above in connection with the Windkessel model. In some implementations, a mathematical expression of the hemodynamic parameter is determined for the first and second diastole. For example, the mathematical expression can be a sum of sinusoids.

[0104] The indices are used to construct a waveform that can be used to determine cardiac output. Calculating cardiac performance may include decomposing a first waveform representing the hemodynamic parameter for a first diastole to determine a first set of sinusoids and decomposing a second waveform representing the hemodynamic parameter for a second diastole to determine a second set of sinusoids. These decompositions may include applying a Fourier transform to the first waveform, the second waveform, or both. The set of sinusoids may include one or more summed sinusoids.

[0105] When a pump is operating in the patient's vascular system, the blood flow in the aorta is determined by the pump contribution (i p ) plus the contribution of the original heart (i h ) plus the patient's heart (i h For example, aortic pressure may be the hemodynamic parameter, and the aortic pressure may be calculated as a sum of sinusoids resulting from a Fourier transform:

number

[0106] By decomposing the hemodynamic parameter over time into its constituent frequencies, the hemodynamic parameter can be determined using a mathematical expression or series of equations. In some implementations, the mathematical expression is an exponential equation based on a comparison of sinusoids. After the hemodynamic parameter waveform is characterized by the mathematical expression, cardiac parameters such as vascular resistance and compliance can be determined from the equation. For example, if the hemodynamic parameter waveform is

number

[0107] In some implementations, a model heartbeat representing a patient's cardiac function can be simulated based on a comparison of sinusoids and used to determine cardiac output, the timing and level of mechanical circulatory support, and other parameters. For example, a blood pump can be operated at a range of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), with each pump speed corresponding to a rotor rotation speed and a corresponding frequency (e.g., 100 Hz, 200 Hz, 300 Hz, 400 Hz, 500 Hz, 1000 Hz, 2000 Hz, 3000 Hz, etc.). Changing the pump speed (or frequency) changes the blood flow in the vasculature resulting from the pump's operation, thereby changing the value of a hemodynamic parameter. By gradually passing through multiple pump speeds (or operating parameters such as blood flow resulting from the pump) to identify corresponding changes in one or more hemodynamic parameters, forming a hemodynamic waveform, and decomposing the hemodynamic waveform resulting from each pump speed, a relationship between pressure and flow during diastole can be established. The patient's overall cardiac function can then be mapped (as a function of the measured hemodynamic parameters) as a mathematical expression that can be used to simulate future cardiac function and inform the delivery and control of mechanical circulatory support to the patient. For example, the measured aortic pressure waveform of a recorded heartbeat can be constructed using the methods described below, and the CO for that heartbeat can be calculated.

[0108] As mentioned above, in some implementations, a brief change in pump speed can be applied to the pump within one heartbeat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heartbeat can be compared to the aortic pressure of a heartbeat without this brief speed change or impulse stimulus. The difference between the two (the aortic pressure of the altered heartbeat and the aortic pressure of the "normal" heartbeat) is called the aortic pressure impulse response: Δp(t)=p1(t)-p2(t) It can be considered as where P1(t) is the pressure waveform measured during impulse stimulation, P2(t) is the pressure waveform without impulse stimulation, and ΔP(t) is the impulse response of the aortic pressure. If this impulse stimulation is applied only during diastole, the difference in total cardiac blood flow for the two heartbeats can be calculated as: Δi(t)=i1(t)-i2(t) It can be expressed as where i1(t) and i2(t) are the pump flows for the impulse-stimulated and non-impulse-stimulated heartbeats, respectively, and Δi(t) is the impulse response of cardiac blood flow. Next, the relationship between aortic pressure and pump flow is calculated in the frequency domain as

number

number

[0109] FIG. 10 shows two plots for the same 10-second period, one for aortic pressure and one for cardiac blood flow. The y-axis of the upper plot represents aortic pressure (mmHg), while the x-axis represents time (seconds). The y-axis of the lower plot represents calculated total cardiac blood flow (liters / minute), while the x-axis represents time (seconds). In this example, systemic vascular resistance R and compliance C are known. For example, R and C can be calculated using aortic pressure measurements taken over the illustrated 10-second period in combination with pump data as described above. Total cardiac blood flow i h +i p is expressed as equation (2):

number

[0110] The foregoing is merely illustrative of the principles of the present disclosure, and the device may be embodied in other aspects than those described above, which are presented for purposes of illustration and not limitation. It should be understood that while the device disclosed herein is shown as being used in the percutaneous insertion of a heart pump, it is applicable to devices for other applications requiring hemostasis.

[0111] Those skilled in the art will appreciate variations and modifications upon reviewing this disclosure. The disclosed features may be implemented in any combination and subcombination (including multiple dependent combinations and subcombinations) with one or more other features described herein. The various features described or illustrated above, including any and all components thereof, may be combined or integrated into other systems. Furthermore, certain features may be omitted or not implemented.

[0112] The systems and methods described above may be implemented locally on the heart pump system or on a controller of the heart pump system, such as an AIC. The heart pump system may include a data processing device. The systems and methods described herein may be performed remotely on a separate data processing device. The separate data processing device may be directly or indirectly connected to the heart pump system via a cloud application. The heart pump system may communicate with the separate data processing device in real time (or near real time).

[0113] In general, aspects of the subject matter and functional operations described herein can be implemented in digital electronic circuitry, or computer software, firmware, or hardware, including the structures disclosed herein and their structural equivalents, or one or more combinations thereof. Aspects of the subject matter described herein can be implemented as one or more computer program products, i.e., one or more modules of computer program instructions encoded on a computer-readable medium for execution by or control the operation of a data processing device. The computer-readable medium can be a machine-readable storage device, a machine-readable storage substrate, a memory device, a composition affecting a machine-readable propagated signal, or one or more combinations thereof. The term "data processing device" encompasses all apparatuses, devices, and machines for processing data, such as, for example, a programmable processor, a computer, or multiple processors or computers. In addition to hardware, an apparatus can include code that creates an execution environment for a given computer program, such as code constituting processor firmware, a protocol stack, a database management system, an operating system, or one or more combinations thereof. A propagated signal is an artificially generated signal, such as a machine-generated electrical, optical, or electromagnetic signal generated to encode information for transmission to an appropriate receiver apparatus.

[0114] A computer program (also known as a program, software, software application, script, or code) can be written in any form of programming language, including compiled or interpreted languages, and can be deployed in any form, for example, as a stand-alone program or as a module, component, subroutine, or other unit suitable for use in a computing environment. A computer program may correspond to a file in a file system. A program can be stored as part of a file that holds other programs or data (e.g., one or more scripts stored in a markup language document), in a single file dedicated to the program, or in multiple cooperating files (e.g., files that store one or more modules, subprograms, or portions of code). A computer program can be deployed to be executed on one computer or on multiple computers located at one site or distributed across multiple sites and interconnected by a communications network.

[0115] The processes and logic flows described herein may be performed by one or more programmable processors executing one or more computer programs to perform functions by manipulating input data and generating output. The processes and logic flows may also be performed by, and apparatus may be implemented as, special purpose logic circuitry, such as an FPGA (field programmable gate array) or an ASIC (application-specific integrated circuit).

[0116] Processors suitable for the execution of a computer program include, by way of example, both general-purpose and special-purpose microprocessors, and any one or more processors of any kind of digital computer. Generally, a processor receives instructions and data from a read-only memory or a random-access memory, or both. The essential elements of a computer are a processor for executing instructions and one or more memory devices for storing instructions and data. Typically, a computer also includes one or more mass storage devices for storing data, such as magnetic, magneto-optical, or optical disks, or is operatively coupled to such mass storage devices for receiving data, transferring data, or both. However, a computer need not necessarily have such devices.

[0117] Examples of changes, substitutions, and modifications will be apparent to those skilled in the art and can be made without departing from the scope of the information disclosed herein. All references cited herein are incorporated by reference in their entirety and form part of this application.

Claims

1. operating a pump configured to be positioned within the heart and configured to operate at an adjustable pump speed at a first pump speed; monitoring a hemodynamic parameter during pumping at said first pump speed; Identifying a first phase of a first heartbeat of the heart over a first time period; predicting a second phase of a second heartbeat of the heart over a second time period based on the monitored hemodynamic parameter; changing the pump speed to a second pump speed during the second phase of the second heartbeat; monitoring the hemodynamic parameters during the second phase; comparing the monitored hemodynamic parameters during the first phase with the monitored hemodynamic parameters during the second phase to calculate a change in the hemodynamic parameters between the first phase and the second phase; and calculating an index representative of vascular performance based on the change in the hemodynamic parameter between the first phase and the second phase; The controller is configured as follows:

2. The controller of claim 1 , wherein the controller is further configured to change the second pump speed to the first pump speed after the second phase.

3. The controller according to any one of claims 1 to 2, wherein the hemodynamic parameter is aortic pressure.

4. The controller of claim 1 , wherein the first phase is one of a systole and a diastole, and the second phase is one of a systole and a diastole.

5. The controller of claim 4 , wherein the first phase is a first diastole and the second phase is a second diastole.

6. The controller of claim 4 , wherein the first phase is a first systole and the second phase is a second systole.

7. 7. The controller of claim 1, wherein the identification of the first phase of the first heartbeat is based on changes in the monitored hemodynamic parameter over time while pumping at the first pump speed.

8. The controller of claim 1 , wherein the prediction of the second phase of the second beat of the heart is based on the identified first phase of a cardiac cycle.

9. In the comparison of the hemodynamic parameters in the first phase with the second phase, the controller: calculating a first volume of blood pumped by the heart in the first phase; calculating a second volume of blood pumped by the heart in the second phase; and determining a numerical difference between the first blood volume and the second blood volume; The controller of any one of claims 1 to 8, further configured to:

10. The controller of claim 8 , wherein the controller is further configured to assess the linearity of the change over time of the hemodynamic parameter between the first phase and the second phase.

11. Determining the cardiac performance of the heart includes: calculating vascular compliance and vascular resistance of the systemic vasculature based on the changes in the hemodynamic parameters between the first phase and the second phase; Calculating the cardiac output of the heart using the Windkessel model; The controller of any one of claims 1 to 10, wherein the controller is determined by:

12. The controller of any preceding claim, wherein the first pump speed is less than the second pump speed.

13. The controller of any preceding claim, wherein the first pump speed is greater than the second pump speed.

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