System and method for determining cardiac performance

Abstract

The invention relates to a system and method for determining heart performance. The systems and methods described herein determine metrics of heart performance via a mechanical circulation support device, and use the heart performance to calibrate, control, and deliver mechanical circulation support to the heart. The system includes 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 a native heart output. Systems and methods operate by using a mechanical circulation support device (e.g., a heart pump) to introduce controlled perturbations of the vascular system and in response to determine cardiac parameters (such as pulsation volume, vascular resistance and compliance, left ventricular end diastolic pressure) and ultimately the native heart output.

Description

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application is a divisional application, the parent application number of which is 201980054368.7, the application date is June 19, 2019, and the invention name is “SYSTEMS AND METHODS FOR DETERMINING CARDIAC PERFORMANCE”. This application claims priority and benefits of the following: U.S. Provisional Patent Application No. 62 / 687,133, entitled “METHODS AND SYSTEMS FOR IMPROVED ASSESSMENT OF VASCULAR AND CARDIAC STATE”, filed on June 19, 2018; U.S. Provisional Patent Application No. 62 / 863,136, entitled “SYSTEMS AND METHODS FOR SYSTEM IDENTIFICATION”, filed on June 18, 2019; and U.S. Provisional Patent Application No. 62 / 863,146, entitled “SYSTEMS AND METHODS FOR DETERMINING CARDIAC PERFORMANCE”, filed on June 18, 2019. The entire contents of the above-referenced applications are incorporated herein by reference. Background Art

[0002] Cardiovascular disease is a leading cause of morbidity, mortality, and healthcare burden worldwide.

[0003] A variety of treatment modalities have been developed for heart health, ranging from drugs to mechanical devices and transplants. Temporary cardiac support devices, such as heart pump systems, provide hemodynamic support and promote cardiac recovery. Some heart 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 TRANSPORT family of devices (Abiomed, Inc., Danvers, Mass.) This type of heart pump system has sensors that detect blood pressure (or assess the pressure differential across the membrane) and can monitor motor current, and uses the sensors and motor current readings to help identify pump positioning.

[0004] The cardiac support required for a given patient can vary from patient to patient. Cardiac output (CO) is the volume flow of blood delivered by the heart. Normal cardiac output is about 5L / min in healthy adults, but can vary based on various factors, including the body structure of a given patient. It is difficult for clinicians to quantitatively determine how much cardiac output a given heart provides, how much additional support the device should deliver, when to deliver it, and how long it takes using known techniques. For patients recovering from interventions or other cardiac care, this determination can be particularly challenging. Therefore, clinicians tend to rely on judgment and indirect estimates of cardiac function, such as using fluid-filled catheters to measure intracardiac or intravascular pressure. Cardiac output (CO) is particularly difficult to quantify. One technology uses a pulmonary artery catheter (PAC) to provide real-time measurements of central venous pressure and pulmonary artery pressure. PAC relies on estimates of CO using Fick's law through measurements of systemic oxygen consumption or bolus thermodilution. However, due to the assumptions that must be made to derive CO measurements and the corresponding lack of precision, the use of PAC in high-risk situations (such as complex interventions and cardiogenic shock) can be limited. Measurements made by PAC ignore dynamic changes in cardiac function and are therefore discontinuous, while nonlinear aspects of the systemic ventricular-vascular connection may not be adequately addressed. Summary of the invention

[0005] The systems and methods described herein determine a measure of vascular and / or cardiac performance (such as CO) via a mechanical circulatory support device (e.g., an intravascular blood pump system), and use the cardiac performance to calibrate, control, and deliver mechanical circulatory support to the heart. The system includes a mechanical circulatory support device and a controller configured to operate the device, receive input indicating device operating conditions and hemodynamic parameters, and determine vascular performance and native cardiac output, including vascular resistance and compliance. The system and method operate by the following steps: using a mechanical circulatory support device (e.g., a heart pump) to introduce a controlled perturbation of the vascular system, and in response determine cardiac parameters (such as stroke volume, vascular resistance and compliance, cardiac contractility, ventricular elastance, CO, and left ventricular end-diastolic pressure), and ultimately determine native cardiac output. Those determined parameters can then be used to calibrate and control further mechanical circulatory support to the heart. By determining the native cardiac output of the heart, a treatment regimen can be applied using mechanical circulatory support (e.g., a blood pump). To achieve therapy, the process control system activates or deactivates a mechanical circulatory support device (eg, a pump) to deliver and regulate the level of support.

[0006] The systems and methods are constructed with a time-varying nonlinear model of the vascular system and use the device-arterial coupling to continuously determine systemic vascular resistance and compliance and thereby quantify cardiac stroke volume. In some embodiments, the systems and methods use a Windkessel model of the vascular system to improve upon traditional linear approximations and provide dynamic changes in vascular response. In some embodiments, the systems and methods are constructed as a cardiac output sensor that can directly determine the native cardiac output of the patient's heart.

[0007] In various adaptations, systems and methods are configured to "ping" the vasculature using a mechanical circulatory support system during a heart beat, and then detect the response of the heart at one or more later periods or time points (e.g., during a later heart beat). "Pinging" involves increasing or decreasing the output of the mechanical circulatory support system (e.g., increasing or decreasing the pump speed of the heart pump system) for a brief period of time (e.g., within a single heart beat), thereby generating spikes in blood pressure and blood flow (e.g., aortic pressure and blood flow leaving the left ventricle). Pinging modifies a hemodynamic parameter (e.g., aortic pressure) from its baseline, and the change is detected and compared to a hemodynamic parameter at another time (e.g., aortic pressure when the heart was not pinged) to determine cardiac performance. The test may entail changing the pump speed during a certain time period (or time point) within a portion of a single cardiac beat (e.g., during a phase of the cardiac beat) and comparing the hemodynamic parameters during this "changed" time period with the hemodynamic parameters during a "normal" operating period or time point when the test is not applied (e.g., during a subsequent cardiac beat).

[0008] Example hemodynamic parameters include: heart rate, blood pressure, arterial oxygen saturation, mixed venous oxygen 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 gradient, left ventricular end pressure, stroke volume, stroke volume index, stroke volume change, 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 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, arteriovenous oxygen content difference, oxygen delivery, oxygen delivery index, oxygen consumption, oxygen consumption index, quantitative oxygen extraction, oxygen extraction index, total peripheral resistance, CO, cardiac index and cardiac power output (CPO).

[0009] Intra-beat interrogation (i.e., adjusting pump operation within a single cardiac beat) allows comparison of hemodynamic parameters occurring between cardiac beats (either sequentially adjacent to one another or separated by other beats) while also minimizing noise (e.g., resonance) that may occur if pump speed is changed over a longer period of time. As mentioned above, in some embodiments, hemodynamic parameter comparison is accomplished by a control system processor that is programmed with a model of the vascular system (such as a two-element Windkessel model) that models and accounts for the nonlinear interaction of changes between pump flow and cardiac operation. The control system uses known terms (received as inputs) to approximate values ​​for pump flow and cardiac operation (e.g., aortic pressure), which allows such models to be easily validated and utilized in clinical applications. The systems and methods provide metrics indicative of a patient's cardiac health, such as resistance and compliance of the systemic vasculature, which allows determination of CO and other aspects of cardiac performance. In some applications, the systems and methods are deployed without the need for additional measurement or diagnostic catheters. The potential to continuously and accurately track changes in systemic vascular performance (eg, resistance and compliance) and estimate stroke volume represents a significant advance compared to traditional measures obtained from PACs or other diagnostics currently deployed in clinical practice.

[0010] In some adaptations, the systems and methods described herein change the pump speed of the heart pump system within a single heart beat 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" heart beat (e.g., a heart beat when the heart pump system is operating at a first pump speed) and during a "modulated" heart beat (e.g., a heart beat when the heart pump system is operating at a pump speed different from the first pump speed within at least a portion of the heart beat). The reference heart beat may occur before or after the modulated heart beat. By modulating or "checking" the heart beat, the system and method can capture and quantify the differences in hemodynamic parameters between the normal heart beat and the modulated heart beat. These differences are then correlated with differences in flow, stroke volume, CO, or other useful metrics of vascular and / or cardiac performance. By changing the pump speed within a short period of time (e.g., a small portion of a heart beat), systemic resistance and compliance can be quantified with higher real-time accuracy at multiple pump speeds without introducing additional noise into the system measurements.

[0011] In some embodiments, a controller is provided, and the controller is configured to perform any of the embodiments, aspects, and methods described herein. For example, the controller may be an automated leaf Impella controller (AIC) from Abiomed, Inc, or any other suitable controller programmed to implement the disclosed functionality. In some embodiments, the systems and methods use a mechanical circulatory support device, such as a heart pump. An example heart pump includes: a catheter; a motor; a rotor operatively coupled to the motor; a pump housing at least partially surrounding the rotor so that the actuated motor drives the rotor and pumps blood through the pump housing; one or more sensors, such as a differential pressure sensor; and a controller. For example, a heart pump system may include: a blood pump having a cannula configured to be deployed within the heart; and a motor positioned within 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 the AIC, or any other suitable control system.

[0012] The systems and methods described herein modify the operation of a mechanical circulatory device within a heart beat so that one or more of the mentioned hemodynamic parameters during that phase are compared with the same parameters during different beats, and thus a metric indicating the patient's vascular performance is calculated. For example, a pump may be inserted intravascularly, and the pump may be operated at a first pump speed (or other output level) during a baseline series of heart beats (including the first heart beat), and then the heart may be "checked" by increasing or decreasing the pump speed during a very short period of time during a second or target heart beat or during a specific phase of the target heart beat (e.g., at or after the dicrotic notch during the target heart beat). Aortic pressure or other hemodynamic parameters are measured both during the baseline series and during the target heart beat. In some adaptations, hemodynamic parameters are measured during the same portion of the first heart beat and the second heart beat (e.g., during contraction in both beats, or at or after the dicrotic notch in both beats). The system then compares the hemodynamic parameter (e.g., aortic pressure) identified during the baseline series (e.g., during the first heart beat) with the hemodynamic parameter identified during the time when the pump speed is increased (e.g., during the second heart beat), for example by using the same sensor and calculating or characterizing vascular resistance or compliance, which can be used to determine CO or modify the operation of the pump to better treat the patient. The baseline series of heart beats (e.g., including the first heart beat) can occur after the interrogation or before the interrogation. The system can use the following described Figure 1 The heart pump system 100 or any other suitable pump may be used to perform such methods.

[0013] To implement the system and method, a pump or other mechanical circulatory support device is positioned within the patient's vascular system (e.g., in the patient's heart) and is operable to change the patient's hemodynamics. For example, the operation of the device can increase the patient's aortic pressure by unloading the left ventricle or other means. In some embodiments, the pump is an intravascular blood pump device placed in the patient's heart via percutaneous insertion. The pump can also be a surgically implanted device, a left ventricular assist device, a counterpulsation device, an expandable heart pump, an extracorporeal membrane oxygenation device, or any other suitable device. The pump may be appropriate because the patient is in cardiogenic shock or is otherwise experiencing a decline in vascular health. The pump can be positioned across the aortic valve so that the blood inlet of the pump is in the left ventricle and the outlet of the pump is in the aorta. The pump contributes to native heart operation so that: CO=i h +i p (1) where CO is total cardiac output, i h is the native cardiac output, and i p is the flow contributed by the pump. Such pumps can provide a lifesaving advantage to patients in cardiogenic shock by increasing the flow of oxygenated blood from the heart and into the coronary arteries and other areas of the vasculature.

[0014] A hemodynamic parameter is monitored while operating a pump or other mechanical circulatory support system (e.g., the parameter may be monitored continuously during cardiac performance, and relevant data indicative of the parameter at selected times during selected cardiac beats may be identified for use). Suitable hemodynamic parameters include parameters relating to blood flow within organs and tissues of the body. Example hemodynamic parameters include: heart rate, blood pressure, arterial oxygen saturation, mixed venous oxygen 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 gradient, left ventricular end pressure, stroke volume, stroke volume index, stroke volume variability, systemic vascular resistance, systemic vascular resistance index, pulmonary vascular resistance, pulmonary vascular resistance index, pulmonary vascular resistance, pulmonary vascular The heart rate, pressure gradient (P) and cardiac output (CPO) may be used to measure the pressure difference between the heart and the lungs. The heart rate, pressure gradient (P) and cardiac output (CPO) may be used to measure the pressure difference between the heart and the lungs. In some embodiments, the pressure gradient (P) may be used during diastole. diff ) can be used in the calculation instead of aortic pressure (e.g., if the pressure difference is known but the aortic pressure is unknown). diff= Aortic pressure minus left ventricular pressure (LVP). In many cases, LVP is much less than aortic pressure during diastole and does not change much during diastole compared to aortic pressure. For these cases, P diff is close enough to aortic pressure (ie, LVP is negligible) that it can be used as a surrogate if AoP is not available. diff If aortic pressure is used instead of aortic pressure, the accuracy of the results in certain calculations described herein may be affected.

[0015] The pump speed is the operating speed of the pump and corresponds to the amount of blood flow provided by the operation of the pump. In some embodiments, the pump speed corresponds to the rotational speed of the rotor. For example, the 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. The pump speed can correspond to a power level or P level, as described above with respect to Figure 1 As described herein. For example, the pump speed is P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8 or P-9 or any other suitable value. In some embodiments, the pump speed instead corresponds to the rate at which the chamber of the pump fills and releases blood. By monitoring the hemodynamic parameters, the systems and methods described herein can detect changes in the hemodynamic parameters over time. Such changes can be used to quantify cardiac performance.

[0016] In some embodiments, a method of detecting a heart beat is provided by measuring hemodynamic parameters during cardiac performance (e.g., during multiple heart beats), identifying and predicting various cardiac phases and their alignment in a timely manner, and then determining when to adjust the pump speed or otherwise check the heart based on the prediction of when the subsequent cardiac phase will occur. The first (baseline) phase of the heart beat in the cardiac beat cycle can be identified as the heart's systole, its diastole, or any other suitable phase or combination of phases, which occurs within a first time period during the cardiac performance. For example, for a given cardiac beat or other cycle, the first time period can be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length of time.

[0017] In the second step, the second phase of the cardiac cycle is predicted or otherwise identified as a target phase for use as when the heart should receive the "check". For example, the second phase can be a second heart beat, a different systolic phase, a different diastolic phase, or any other suitable phase or phase combination, provided that the second phase is selected as the time when the effect of the "check" should be delivered to the heart. The system and method can predict when the second phase of the cardiac cycle begins based on previously monitored hemodynamic parameters (e.g., aortic pressure measured during the first phase of the cardiac cycle). The second phase of the cardiac cycle is predicted to occur within a second time period. For example, the second time period can be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable time length. The second time period can be set to the length of the second phase of the cardiac cycle corresponding to a specific patient. For example, if the second phase of the cardiac cycle is a diastolic phase, the second time period can be set to the average diastolic time period of the specific patient, or can be set to the predicted length of the next diastolic phase that occurs. In some examples, the second time period can be preset to a time period that is less than the time period of the cardiac beat. The second phase is predicted based on the monitored hemodynamic parameters and the identified first phase of the cardiac cycle. The cardiac signal can be monitored, and based on the monitored signal, the system and method predict when the next diastole or systole will occur. By timing the expected cardiac cycle phase, the system and method can then time the increase or decrease ("check") of the pump speed to accurately start (or enable its impact to be delivered) at the beginning of the second phase of the cardiac cycle. For example, the system and method can be configured to check the heart when the pump output is temporarily increased (for example, by unloading the heart with a higher pump speed) so that the resulting increase in blood flow occurs simultaneously with the start of a preferred point or period in the target cardiac beat, such as at the beginning of the dicrotic notch in the subsequent cardiac beat or a moment later, or in the middle of the contraction of the subsequent cardiac beat, or during the entire diastole, or during a predetermined portion of the subsequent cardiac beat.

[0018] In certain embodiments, the first phase of the cardiac cycle is diastole of a first heart beat, and the second phase of the cardiac cycle is diastole of a second immediately subsequent heart beat. The first phase of the cardiac cycle may be systole of a first heart beat, and the second phase of the cardiac cycle may be systole of a second heart beat. In some embodiments, the second phase is during a heart beat that has a number of beats removed from the first heart beat, while in other embodiments, the second phase is during a heart beat that is adjacent to the first heart beat.

[0019] After establishing the baseline and target heart beat phase, the pump speed is then changed to check the heart, for example by operating the pump at a second pump speed different from the first pump speed during a second heart beat. Pump speed regulation can be an increase or decrease in pump speed for checking the heart beat. For example, the pump can be adjusted so that the pump speed is temporarily increased during a period of heart beat (for example, during diastole of the heart beat). The pump can also be adjusted so that the speed returns to the baseline or is otherwise reduced during the same period of the heart beat or at some other point during the same heart beat. The check can also be configured to occur in the opposite direction - by temporarily reducing the pump speed from the baseline.

[0020] Briefly changing the pump speed and returning to baseline causes the heart to be temporarily "pinned," such as at a higher pump speed. In an embodiment, the heart pump is operated at a first baseline pump speed, then temporarily changed to a higher second pump speed during the systole or diastole phase (or other target time period) of a subsequent heart beat, and then quickly returned to the first pump speed. In some embodiments, the change in pump speed lasts for less than the length of a heart beat, so that the pump returns to its baseline during the same heart beat in which it was pinned. For example, the entire pinning may occur within a target heart beat, such that the duration of the pinning is shorter than the target heart beat. Changing the pump speed within a single heart beat reduces the effects of noise on the collection of hemodynamic data between the first and second phases to improve accuracy.

[0021] The pump speed is adjusted so that the change in pump speed is delivered during the desired portion. For example, the change in speed can be delivered during contraction, diastole, or both within the heart beat. In an embodiment, in order to adjust the pump speed, the controller sends a signal to the pump to change the pump speed in time before the target phase begins to account for any time delay between sending the control signal and the change in pump speed. The pump should deliver the actual speed increase or decrease during the desired portion of the target heart beat (e.g., diastole, at or after the venous notch). The check is timed so that the increased pump speed appears in time during the process of a known period of the heart beat. For example, the start of the speed check can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, the peak systolic pressure, or any other suitable time. The end of the speed check can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, the peak systolic pressure, or any other suitable time. In some embodiments, the check is achieved by increasing or decreasing the pump speed during a set time period. For example, the check can be synchronized with the start of diastole so that the check occurs during diastole. Alternatively, the check may be synchronized with the end of diastole so that the check occurs during the systole of the following heart beat. In other adaptations, the check is synchronized with the start of systole, the end of systole, the peak systolic pressure, or any other suitable time. The check is configured to last for a set time period. In some adaptations, the check is set to last for a time period corresponding to the length of the phase of the heart beat. For example, the check may be set to last for approximately 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable length of time.

[0022] The heart or other vascular system is checked by temporarily adjusting the mechanical circulatory support (e.g., pump speed) to apply a disturbance to the heart or other vascular system, which allows determination of systemic resistance, compliance, and additional measures of cardiac performance, including cardiac output. Such determination can be completed without introducing additional hardware (except for the pump that provides hemodynamic support) into the patient's body (although such additional hardware can still be used if necessary). In order to make a determination, the hemodynamic waveform (e.g., aortic (or ventricular) pressure waveform) during a regular heart beat and a checked heart beat is compared via a nonlinear model (such as the Windkessel model) described below. The change in the pressure waveform (or other hemodynamic parameter) between a normal heart beat and a checked heart beat is reflected in different values ​​within the model during two time periods (one time period for a baseline (normal) heart beat, and one time period for a target (check) heart beat), thereby creating two model equations. Knowing the pressure waveform for two time periods allows the number of unknown variables between the two model equations to be reduced, so that resistance and compliance can be calculated. The total cardiac flow can then be calculated using the calculated resistance and compliance values ​​and the aortic pressure waveform by applying equation (2): Where C is compliance, R is resistance, P is the aortic pressure waveform, i h is the native heart flow, i p is the flow rate contributed by the pump, and i h +i p is the total cardiac flow. This can be obtained by taking the total cardiac flow i generated by equation (1) over a certain time period (e.g., 5 seconds, 10 seconds, or 30 seconds) h +i p CO can be calculated by averaging the CO values ​​of the heart. For example, the time period can be the length of a single heart beat, and the CO calculation can represent the CO during that heart beat. In at least this aspect, the systems and methods feature a CO sensor for determining the native cardiac output of the heart based on the interrogation techniques disclosed herein.

[0023] Adaptation of the techniques described above may be applied in various ways. In some embodiments, a hemodynamic parameter is monitored during the second phase of the second heart beat. For example, the heart pump system may continuously monitor aortic pressure or any other hemodynamic parameter. In some embodiments, the hemodynamic parameter monitored during the first phase is compared with the hemodynamic parameter 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. The numerical difference between the first blood volume and the second blood volume may be calculated to quantifiably compare the hemodynamic parameter during the first phase with the hemodynamic parameter during the second phase. For example, the area under the curve (AUC) of the flow curve may represent blood volume. The difference in AUC during the first phase and during the second phase may indicate the difference in blood volume pumped at the first pump speed and the second pump speed. Comparing the hemodynamic parameters between the first phase and the second phase may also include evaluating the linearity of the change in the hemodynamic parameter. For example, the aortic pressure may not scale linearly between pump speeds, which means that the change in aortic pressure from one pump speed to the next may not be a linear progression. Depending on how changes in aortic pressure scale between pump speeds, aortic pressure at different pump speeds can be predicted.

[0024] A metric indicative of cardiac performance of the heart may be calculated based on changes in hemodynamic parameters between the first phase and the second phase. For example, a metric indicative of cardiac performance may be determined from different pressure waveforms of the cardiac cycle during the first heart beat and the second heart beat; the metric may be a prediction of systemic resistance, systemic compliance, CO, CPO, stroke volume, stroke work, ejection fraction, cardiac contractility, ventricular elastance, cardiac index, patient survivability. Many metrics indicative of cardiac performance are interrelated. For example, CO is determined based on the flow rate of blood pumped through a blood vessel placed in the patient's heart. Stroke volume is an index of left ventricular function, and its formula is SV=CO / HR, where SV is stroke volume, CO is cardiac output, and HR is heart rate. Stroke work is the work done by the ventricle to eject a certain volume of blood, and can be calculated from stroke volume according to the equation 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 and heart rate. CPO is a measure of cardiac function that represents the pumping capacity of the heart in watts. CPO is calculated using the cardiac power output equation, which is represented below by the following equation: CPO=MAP*CO / 451 (3) Where CPO is cardiac power output, MAP is mean aortic 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 blood volume in the ventricle. Other parameters such as chamber pressure, preload state, afterload state, cardiac recovery, flow load state, variable volume load state, and / or cardiac cycle flow state can be calculated from these values ​​or determined via these parameters. In some embodiments, via a two-element Windkessel model of the vascular system (e.g., Figure 5 The process uses a time-varying, nonlinear model of the vascular system and uses the coupling between the intravascular blood pump device and the patient's hemodynamic function (a well-controlled simulation of the ventricle-vascular coupling) to continuously determine systemic vascular resistance and compliance and quantify cardiac stroke volume without the need for additional external measurements.

[0025] The operation of the pump may be adjusted based on the metric indicative of cardiac performance. Adjusting the pump operation may include increasing the pump speed, decreasing the pump speed, adjusting pump placement, shutting off the pump, or any other suitable adjustment. For example, where the metric indicative of cardiac performance is stroke volume, if the stroke volume is below a threshold, the pump speed may be increased, and if the stroke volume is above the threshold, the pump speed may be decreased.

[0026] In some embodiments, a CO sensor for determining the cardiac output of a patient's native heart is provided. The CO sensor may include one or more hardware, software, and firmware elements configured to perform the methods described herein. In some embodiments, the CO sensor includes a mechanical circulatory support device (e.g., an intravascular blood pump) having a pressure sensor and a processor configured to receive measurements from the pressure sensor and use intra-beat verification to determine the native heart output, as described herein. The mechanical circulatory support device may be configured to be at least partially placed in the patient's heart. In some adaptations, the intravascular blood pump includes: a cannula; a rotor configured to pump blood through the cannula; and a drive mechanism configured to impart power to rotate the rotor. In some embodiments, the cannula is configured to extend across the aortic valve so that the distal end of the cannula is within the left ventricle and the proximal end of the cannula is within the aorta. For example, when the cannula is placed across the aortic valve so that the blood inlet of the pump is in the left ventricle and the outlet of the pump is in the aorta, the heart pump system may be considered to be "in place". The drive mechanism may include an onboard motor, a drive cable, a drive shaft, or any other suitable element or combination thereof.

[0027] 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 embodiments, the pump includes a pump housing and a motor housing coupled to the cannula at a distal end of the motor housing. The rotor may rotate within the pump housing to induce blood to flow into the cannula.

[0028] The CO sensor may include a hemodynamic parameter sensor that is operatively positioned at (or proximal or distal to) the blood pump and configured to detect pressure within the blood vessel that is at least partially caused by the pumping of blood within the blood vessel. For example, the pressure sensor may be an optical sensor 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, wherein one side or surface of the differential pressure sensor may be exposed to aortic pressure, a second side or surface of the differential pressure sensor may be exposed to ventricular pressure, and the differential pressure sensor may measure the difference between the aortic pressure and the ventricular pressure.

[0029] The CO sensor includes a controller electrically coupled to the pressure sensor and configured to detect a signal indicative of blood pressure from the sensor. All or a portion of the controller may be in a controller unit separate from / remote from the intravascular blood pump. In some embodiments, the control system is internal to the intravascular blood pump.

[0030] The controller may be configured to calculate CO based on a nonlinear model that relates CO to vascular resistance and compliance, which are based on changes in hemodynamic values ​​due to examination of the heart. For example, the nonlinear model may be a Windkessel model, or a simplified Windkessel model used in a manner associated with a heart pump system positioned across the aortic valve of a patient. The governing equation for this model is: Where C is compliance, P is aortic pressure, R is resistance, i h is the flow from native heart operation, and i p is the flow from the pump. During diastole, the aortic valve is closed, so the only flow through the left ventricle comes from the pump located across the valve. By neglecting the cardiac current source and assuming a constant pump flow, the model can be simplified as follows: Resistance and compliance can then be determined via the following two equations, where P1 and P2 are pressure waveforms measured at different pump speeds (e.g., one at the time of the test and the other before or after the test): At low pump speed, the pump flow rate i p1 And therefore, i in equation (5) p1 The R term can be approximated to zero, resulting in a simple exponential form of equation (5): The pump flow rate is determined by a number of factors, including pump speed, the pressure difference between aortic and ventricular pressures, and the pump model. For example, for a particular pump model, the pump flow rate is close to zero when the pressure difference between aortic and ventricular pressures is approximately 40 to 50 mmHg and the pump speed is approximately 23,000 RPM. At this pressure level (40 to 50 mmHg), the flow rate should not be approximated to zero when the speed is above 30,000 RPM. After using equation (6) and the above simplification to determine R, i p1 The R term is added back to equation (5) to then accurately determine C.

[0031] In some embodiments, P0 is assumed to be proportional to the inverse of the corresponding pump speed, such that Therefore, at t = 0, equations (5) and (6) become: P2(t=0)=P 02 +i p2 R (6a) From equations (5a) and (6a), R can be calculated as follows: Where P1(t=0) and P2(t=0) are the initial aortic pressures measured at the beginning of diastole of the first heart beat and the second heart beat, respectively. For a desired time period (e.g., 5 seconds, 10 seconds, 1 minute, or any given time period as desired), the total cardiac flow can then be calculated from equation (3). Specifically, the CO or average cardiac flow over the desired time can be calculated as: where t end is equal to the end of the desired time period, and t start = ... end )-P(t=t start )) can be close to zero or much smaller than CO, so that it can be approximated as zero in equation (8). Therefore, CO can be calculated as: CO then equals: Where MAP is the value from t start to end The mean arterial pressure within the desired time window.

[0032] In some aspects, according to the systems and methods described herein, a blood pump can be used to provide mechanical circulatory support to a patient. Providing mechanical circulatory support can include actuating a blood pump within the patient's vasculature, determining the CO of the patient's heart using any of the systems and methods described herein, and adjusting the pumping speed of the blood pump based on the determined CO.

[0033] In some aspects, the mechanical circulatory support system may include an intracardiac blood pump having a cannula configured to extend within the left ventricle of the heart and a pressure sensor configured to detect left ventricular end-diastolic pressure. The system may be configured to determine CO according to any of the methods described herein.

[0034] In some embodiments, the pump is placed in the patient's heart. The pump may be introduced to the patient because the patient is in cardiogenic shock or, is undergoing coronary intervention, or has a heart attack, or is otherwise experiencing a decline in heart health. The pump may be positioned across the aortic valve so that the blood inlet of the pump is in the left ventricle and the outlet of the pump is in the aorta. The pump contributes to the non-native heart operation so that the CO from the heart is equal to the native CO plus the pump output.

[0035] A first aortic pressure wave can be detected. The first aortic pressure wave reflects multiple beats of the heart, each reflected beat including a dicrotic notch. The pressure waveform can be measured via a pressure sensor. In some embodiments, the pressure sensor can be located on-board the pump. In some embodiments, the pressure sensor is located outside the pump and receives a fluid or electrical signal. The pressure sensor can communicate with a controller that is configured to control the operation of the pump.

[0036] Hemodynamic support may 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 may be a first rotor speed, such as the P level described above. During a second beat of the plurality of beats, hemodynamic support to the heart is adjusted by providing a second pumping rate to the heart during the second beat (e.g., during contraction following its dicrotic notch). The first pumping rate will be set to be different from the second pumping rate.

[0037] A second aortic pressure wave of the heart may be detected during a second beat. The second aortic pressure wave may be compared to a portion of the first aortic pressure wave corresponding to the second beat to detect changes in the second aortic pressure wave. In some examples, the second aortic pressure wave may be compared to the first aortic pressure wave by comparing the area under the curve (AUC) within the time length represented by a portion of the second aortic pressure wave with the area under the curve (AUC) within the same time length represented by a portion of the first aortic pressure wave. In some examples, the global maximum and minimum values ​​of the first aortic pressure wave and the second aortic pressure wave may be compared. The shapes or slopes of the first aortic pressure wave and the second aortic pressure wave may be compared, that is, the changes over time through the derivatives of these waves. In some embodiments, the first aortic pressure wave and the second aortic pressure wave are compared 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. Changes between the first aortic pressure wave and the second aortic pressure wave can be used to identify the resistance and compliance of the systemic vascular system. Comparing the hemodynamic parameter between the first aortic pressure wave and the second aortic pressure wave may also include evaluating the linearity of the change in the aortic pressure wave between the first pump speed and the second pump speed. For example, the aortic pressure may not scale linearly between pump speeds, meaning that the change in aortic pressure from one pump speed to the next may not progress linearly. Depending on how the change in aortic pressure scales between pump speeds, the aortic pressure at different pump speeds may be predicted.

[0038] In some embodiments, CO is determined based on a nonlinear transfer function that relates CO to systemic resistance and compliance. In some embodiments, the nonlinear transfer function comprises a Windkessel model. In some embodiments, the transfer function further relates to the aortic pressure waveform.

[0039] Systems and methods may compare a hemodynamic parameter during a first heart beat to a hemodynamic parameter during a second heart beat to calculate a change in the hemodynamic parameter between the first heart beat and the second heart beat. The change is caused at least in part by a difference between a first output level and a second output level of a mechanical circulatory support 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 a first pump output level to a second pump output level indicates a contribution of the mechanical circulatory support device to the change in aortic pressure.

[0040] A metric indicating cardiac performance of a blood vessel and / or heart may be calculated based on a change in a hemodynamic parameter between a first heart beat and a second heart beat. For example, the hemodynamic parameter during the first heart beat and the second heart beat may be calculated via a nonlinear model (such as a Windkessel model) described below. The change in the hemodynamic parameter between a normal heart beat and a checked heart beat is reflected in different values ​​within the model during two time periods (one time period for the first heart beat and one time period for the second heart beat), and thus two model equations may be used to determine cardiac performance. In some embodiments, the metric indicating cardiac performance is cardiac output. In order to calculate cardiac output, vascular resistance and compliance may be determined based on a change in the hemodynamic parameter between the first heart beat and the second heart beat, as described above. Knowing the hemodynamic parameter waveforms for the two time periods allows the number of unknown variables between the two model equations to be reduced, so that resistance and compliance (and ultimately cardiac output) may be calculated.

[0041] In some embodiments, the systems and methods described herein include modeling the patient's heart beats to represent the heart beats as a series of sinusoids, and the processor can use these sinusoids to construct one or more heart beats representing the patient's cardiac function. Then, the processor uses the constructed (one or more) heart beats to adjust the pump speed. As described above, the blood pump operates at a first pump speed (or other operating parameters), and is then adjusted to a second pump speed (or other operating parameters) to check the heart, and then quickly reduced to a baseline first speed or parameter. During the pump operation (including during the check period), monitor hemodynamic parameters (e.g., aortic pressure). The processor calculates a metric indicating the cardiac performance of the heart based on the following: (i) a first operating parameter (pump speed), (ii) a second operating parameter (e.g., pump speed during the check period), and (iii) a hemodynamic parameter during a first time period and a second time period (e.g., during a first diastole and a second diastole). The metric is used in a transfer function or a system of equations (such as, the transfer function or system of equations described above for the Windkessel model). A mathematical representation of the hemodynamic parameter is determined by the controller processor for the first diastolic period and the second diastolic period. For example, the mathematical representation may be a sum of sinusoidal curves or other waveform functions indicative of the hemodynamic parameter at a given pump speed.

[0042] The cardiac performance is then calculated by the processor from the sum of sinusoids or other waveforms. The calculations may include: deconstructing a first waveform representing hemodynamic parameters for a first diastole (where the pump is operating at a first pump speed) to determine a first set of sinusoids; and deconstructing a second waveform representing hemodynamic parameters for a second diastole (where the pump is operating at a second pump speed) to determine a second set of sinusoids. The deconstructions include applying a Fourier transform to the first waveform, the second waveform, or both. A set of sinusoids may include one or more sinusoids added together.

[0043] Since the blood flow in the aorta is equal to the pump contribution (i p ) plus native heart contribution (i h ), so the first set of sinusoids and the second set of sinusoids can be compared to determine the contribution of the patient's heart to the blood flow in the aorta (i h ). For example, if the hemodynamic parameter is aortic pressure, it can be expressed as the sum of sinusoids generated by Fourier transformation, such as: Where P is the aortic pressure, f n is the frequency associated with the pump speed or other operating parameter, and A n and θ nare coefficients of the operating parameters. Changing the operating parameters will change the sinusoids. Since the change in pressure between operating parameters will be proportional to the change in flow, the difference in each set of sinusoids between the operating parameters can be used to calculate the difference in flow from the changed operating parameters. In some embodiments, the Fourier transform can be calculated for each pump speed in a series of pump speeds. In some embodiments, due to the limitation of speed changes within a short period of time (i.e., the time it takes to ramp up the pump to an increased speed or to decelerate the pump to a decreased speed), the patient's response to the "tested" pump speed may be minimal.

[0044] Decomposing the hemodynamic parameter over time into its component frequencies (resulting from changes in the operating parameters of the device) allows the hemodynamic parameter to be characterized using a complex mathematical equation or set of equations. In some embodiments, the mathematical representation is an exponential equation based on a comparison of sinusoids. After the hemodynamic parameter waveform has been characterized by a mathematical equation, cardiac parameters (such as vascular resistance and compliance) can be determined from the equation. For example, if the hemodynamic parameter waveform is characterized as (where B is equal to R*C, and D is equal to i p *C, P is pressure, R is systemic resistance, and C is systemic compliance), then systemic resistance and compliance values ​​can be calculated by solving a system of equations with these coefficients for at least three time points (i.e., with three known pressure measurements corresponding to three known pump operating parameters).

[0045] The heart beats representing the patient's cardiac function can be simulated based on the comparison of sinusoids indicating incremental changes in hemodynamic parameters caused by changes in pump operating parameters. For example, a blood pump can be operated at a series of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), each of which corresponds to the rotation rate of the rotor within the pump and a similar 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) will change the value of the hemodynamic parameter because it will change the blood flow in the vascular system provided by the operation of the pump. By gradually stepping through multiple pump speeds (or operating parameters, such as the blood flow provided by the pump) to identify corresponding changes in one or more hemodynamic parameters, forming a hemodynamic waveform, and deconstructing the hemodynamic waveform generated by each pump speed, a relationship between pressure and flow during diastole is established. The patient's overall cardiac function can then be mapped into a mathematical representation (as a function of measured hemodynamic parameters) that can be used to simulate future cardiac function and inform the delivery and control of mechanical circulatory support for the patient. For example, the method described below can be used to construct a measured aortic pressure waveform for any recorded heart beat - allowing CO to be calculated for that heart beat.

[0046] As described above, in some embodiments, a brief change in pump speed can be applied to the pump within one heart beat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heart beat can be compared to the aortic pressure of a heart beat without this brief speed change or impulse stimulus. The difference between the two (the aortic pressure of the altered heart beat and the aortic pressure of the "normal" heart beat) can be considered the impulse response of the aortic pressure: Δp(t)=p1(t)-p2(t) Wherein P1(t) is the pressure waveform measured with pulse stimulation, P2(t) is the pressure waveform without pulse stimulation, and ΔP(t) is the pulse response of aortic pressure.

[0047] If this pulse stimulus is applied only during diastole, the difference in total cardiac flow for the two heart beats can be expressed as: Δi(t)=i1(t)-i2(t) where i1(t) and i2(t) are the pump flows for heart beats with and without pulse stimulation, respectively, and Δi(t) is the impulse response of the heart flow.

[0048] The relationship of aortic pressure to pump flow can then be estimated in the frequency domain as: Where ΔP(f) is the frequency domain representation (eg, Fast Fourier Transform or FFT) of Δp(t), ΔI(f) is the frequency domain representation of Δi(t), and H(f) is the frequency domain transfer function of aortic pressure versus pump flow.

[0049] Once this relationship H(f) is established as outlined above, the total cardiac flow for any heart beat with the aortic pressure measured as p(t) can be calculated as: where P(f) is the frequency domain representation of p(t) and IFFT is the Inverse Fast Fourier Transform. BRIEF DESCRIPTION OF THE DRAWINGS

[0050] Figure 1 An illustrative heart pump system is shown inserted into a blood vessel of a patient; Figure 2 illustrates a process for calculating a metric indicative of cardiac performance of a heart according to certain embodiments; Figure 3 shows a graph 300 of pressure versus time for a heart pump system according to certain embodiments; Figure 4 shows a graph of pressure, motor speed, and flow rate versus time according to certain embodiments; Figure 5 shows a Windkessel model according to certain embodiments; Figure 6 shows a CO sensor coupled to a patient according to certain embodiments; Figure 7 illustrates a process for determining CO according to certain embodiments; Figure 8 illustrates a process for determining changes in hemodynamic parameters between heart beats in accordance with certain embodiments; and Fig. 9 illustrates a process for determining CO according to certain embodiments; Fig.10 Shown is a graph of aortic pressure and a graph of cardiac flow for the same ten second period in accordance with certain embodiments. DETAILED DESCRIPTION

[0051] In order to provide a comprehensive understanding of the systems, methods, and devices described herein, certain illustrative embodiments will be described. Although the embodiments and features described herein are specifically described as being used in conjunction with a percutaneous heart pump system, it will be understood that the components and other features outlined below can be combined with each other in any suitable manner and can be adapted and applied to other types of cardiac therapy and heart pump systems, including heart pump systems implanted using surgical incisions, aortic pumping, etc.

[0052] The systems, devices and methods described herein enable a support device that resides completely or partially within an organ to assess the function of the organ. In particular, these systems, devices and methods enable a heart pump system (such as a percutaneous ventricular assist device) to be used to assess the function of the heart. For example, such a device can be used to treat cardiogenic shock.

[0053] Using a heart pump system to assess the function of the heart can alert health care professionals to changes in heart function and allow professionals to customize the degree / level of support provided by the assist device (i.e., the flow rate of blood pumped by the device) based on the needs of a specific patient. For example, the degree of support can be increased when the patient's heart function deteriorates, or the degree of support can be reduced when the patient's heart function recovers and returns to a baseline of normal heart function. This can allow the device to dynamically respond to changes in heart function to promote heart recovery, and can allow patients to gradually wean themselves off treatment. In addition, an assessment of heart function can indicate when it is appropriate to terminate the use of the heart pump system. Although some of the embodiments presented herein relate to a heart pump system that is implanted across the aortic valve and resides partially in the left ventricle, these concepts can be applied to devices in the heart, cardiovascular system, or other areas of the body.

[0054] Assessment of cardiac function may include utilizing cardiac device interactions to determine cardiac parameters. Using the Windkessel model of the vascular system to improve traditional linear approximations and provide dynamic changes in vascular response, the systems and methods described herein introduce controlled perturbations of the vascular system through the 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 use a mechanical circulatory support system to "check" the heart beat. "Checking" includes increasing the pump speed of the heart pump system over a period of time (e.g., within a single heart beat), thereby generating spikes in aortic pressure and flow. During the check, a hemodynamic parameter is altered and can be detected and compared with the hemodynamic parameter at another time (i.e., when the heart pump system is not checked) to calculate other hemodynamic parameters or otherwise measure cardiac performance.

[0055] Continuously measuring vascular and cardiac performance by using the action of the heart pump system can provide additional clinical data to help adjust (titration) appropriate device support. The systems and methods described herein also provide the use of device-arterial connection to determine the state of the heart and blood vessels (including determining the native heart output). The mechanical circulatory support system presented herein resides in the heart and works in parallel with the native ventricular function. Unlike some more invasive devices, this allows the system to be sensitive enough to detect native ventricular function. Therefore, these systems, devices and methods enable the mechanical circulatory support system to be used not only as a support device, but also as a diagnostic and prognostic tool. The heart pump system can function as a sensor, which extracts information about cardiac function by being hydraulically connected to the heart. In some embodiments, the heart pump system operates at a constant level (e.g., a constant rotational speed of the rotor) while measuring the power delivered to the auxiliary device. In certain embodiments, the speed of the rotor of the heart pump system can be varied (e.g., as a delta function, a step or a ramp function) to further detect native heart function. Figure 1 An illustrative heart pump system is shown 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 support system (and sensors for obtaining physiological data from a patient) can be used with the present disclosure. In some embodiments, the systems and methods described herein can use an expandable pump (e.g., Heartmate PHP TM family of devices (Thoratec Corporation)) or a left atrium-femoral artery bypass pump (e.g., TandemHeart family of devices (LivaNova, PLC)). In some embodiments, the systems and methods described herein may be used Family Devices (Abiomed, Inc., Danvers, Mass.) The heart pump system 100 can be operated in the heart, partially in the heart, outside the heart, partially outside the heart, partially outside the vascular system, or in any other suitable location in the patient's vascular system. The heart pump system can be considered to be "in place" when the cannula 173 is placed across the aortic valve so that the blood inlet of the pump (e.g., blood inlet 172) is in the left ventricle and the outlet of the pump (e.g., outlet opening 170) is 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 can be in a controller unit that is separate from / away from the heart pump 106. In some embodiments, the control system 104 is inside 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 in which a pump element is formed. The pump 100 includes a pump housing 134 and a motor housing 102 coupled to a cannula 173 at a distal end 111 of the motor housing 102. Impeller blades on a drive shaft can rotate within the pump housing 134 to induce blood to flow into the cannula 173 at a suction head 174. The suction head 174 provides a blood inlet 172 at a distal portion 171 of the cannula 173. The 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.

[0056] Rotation of the drive shaft within the pump housing 134 causes the pump elements to rotate within the bearing gaps. A hemocompatible fluid is delivered through the motor housing 102 by the elongated conduit 105 to the proximal 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 gaps of the pump. After leaving the bearing gaps, the hemocompatible fluid follows a flow direction 123 and becomes entrained in the blood stream and flows with the blood into the aorta.

[0057] Heart pump 100 is inserted into the patient's blood vessel through sheath 175. Pump housing 134 encloses the rotor and internal bearings and can be sized for percutaneous insertion into the patient's blood vessel. In some embodiments, the pump is advanced through the vasculature and over the aortic arch 164. Although the pump is shown in the left ventricle, the pump can alternatively be placed in the right heart so that blood is pumped from the patient's inferior vena cava or right atrium through the right ventricle into the pulmonary artery.

[0058] A flexible protrusion 176 is included at the distal portion 171 of the cannula 173, distal to the suction head 174, to stabilize the heart pump 100 in a vessel or chamber of the heart. The flexible protrusion 176 is non-traumatic and helps prevent the suction head 174 from approaching the vessel wall, where the suction head 174 may become stuck due to suction. The flexible protrusion 176 extends the pump 100 mechanically rather than hydraulically because the flexible protrusion 176 is non-suctionable. In some embodiments, the flexible protrusion may be formed as a pigtail. In some aspects, the pump need not include a flexible protrusion.

[0059] The elongated conduit 105 houses a connector 126 having a fluid supply line and electrical connection cables. The connector 126 also supplies hemocompatible fluid from a fluid reservoir contained within the control system 104 to the pump.

[0060] The control system 104 includes a controller 182 that controls the pump 106 by delivering power to the motor and controlling the motor speed. The control system 104 includes circuitry for monitoring the motor current for a drop in current (indicating air in the line), a change in the differential pressure signal, flow position, aspiration, or any other suitable measurement. In some embodiments, the control system 104 includes a display screen to show measurements such as the differential pressure signal and motor current. The control system 104 may include an alarm sound, light, or indicator to alert the operator to a sensor failure, a disconnection or break in the connection 126, or a sudden change in the patient's health.

[0061] The motor 108 is configured to operate at a speed required to maintain the rotor at a set speed. As a result, and as further described below, the motor current drawn by the motor to maintain the rotor speed can be monitored and used to understand the underlying heart state. The control system 104 is configured to change the speed of the pump during the cardiac cycle of the assisted heart, thereby causing a change in the blood flow through the pump, and the speed change of the pump is synchronized with the heart beat by means of at least one event per cardiac cycle, and the event is related to a predetermined event in the cardiac cycle (i.e., the systems, devices and methods described herein use the heart pump system to "check" the heart beat). "Checking" occurs when the pump speed of the heart pump system (or other mechanical circulatory support device) increases or decreases within a relatively short period of time (e.g., during a phase of the cardiac cycle) and then changes to a baseline or another speed. The pump speed can be increased within a certain period of time within a single heart beat or across multiple heart beats.

[0062] The heart pump can operate at a variety of pump speeds or P levels. The P level is the performance level of the heart pump system and is related to the flow control of the system. As the P level increases, the flow rate, motor current, and revolutions per minute associated with the heart pump system increase; therefore, higher P levels correspond to higher flow rates and revolutions per minute associated with the heart pump system. For example, power level P-1 may correspond to a first revolutions per minute (RPM) of the rotor, while power level P-2 corresponds to a second RPM number. In some examples, the pump operates at ten different power levels ranging from P-0 to P-9. These P levels may correspond to 0 RPM to 100,000 RPM or any suitable number. Changing the speed of the rotor changes the CO of the heart, such as Figure 3 As shown in and described below.

[0063] In some embodiments, the pump speed is increased during contraction, diastole, or both within a single heart beat. The check is timed so that the increased pump speed occurs within a known period of the heart beat. For example, the start of the speed check can be synchronized with the start of diastole, the end of diastole, the start of contraction, the end of contraction, the peak systolic pressure, or any other suitable time. The end of the speed check can be synchronized with the start of diastole, the end of diastole, the start of contraction, the end of contraction, the peak systolic pressure, or any other suitable time. In some embodiments, the pump speed is increased or decreased within a set time period. For example, the start of the check can be synchronized with the start of diastole, at or after the dicrotic notch, with the end of diastole, the start of contraction, the end of contraction, the peak systolic pressure, or any other suitable time. The check can be continuous for a set time period. For example, the check can be continuous for 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable time length.

[0064] The control system 104 includes a current sensor (not shown). The controller 182 supplies current to the motor 108 via the connector 126 (such as, through one or more wires). The current supplied to the motor 108 via the connector 126 is measured by the current sensor. The load experienced by the motor of the mechanical pump corresponds to the force of the pressure head or the difference between the aortic pressure and the left ventricular pressure. The heart pump 106 experiences a nominal load during steady-state operation at a given pressure head, and changes from this nominal load are the result of changing external load conditions, such as the dynamics of left ventricular contraction. Changes in dynamic load conditions change the motor current required to operate the pump rotor at a constant or substantially constant speed. As described above, the motor can be operated at a speed required to maintain the rotor at a set speed, and the motor current drawn by the motor to maintain the rotor speed can be monitored and used to detect potential heart states. By using the pressure sensor 112 to monitor the pressure head simultaneously during the heart beat cycle, the heart state can be accurately quantified and understood. The heart 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 characterize cardiac function. The cardiac parameter estimator 185 can access the stored lookup table to obtain additional information to characterize cardiac function based on the pressure signal and the current signal. For example, the cardiac parameter estimator 185 can receive the aortic pressure from the pressure sensor 112, and in the case of using the lookup table, the aortic pressure can be used to determine the pressure difference (delta pressure). The cardiac parameter estimator 185 can be software programmed in the controller 182, or it can be independent hardware connected to the controller 182 by a wired or wireless connection. The cardiac parameter estimator 185 is configured to execute the algorithm described herein. For example, the cardiac parameter estimator 185 can be configured to estimate the pump flow based on the current delivered to the pump, and can be configured to determine the native cardiac output according to the method described herein.

[0065] Various embodiments 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 the blood pressure outside the cannula, and the other side is exposed to the blood pressure inside the cannula. The sensor generates an electrical signal (differential pressure signal) proportional to the difference between the pressure outside the cannula and the internal pressure, which can be displayed by the heart pump system. When the heart pump system is placed in the correct position across the aortic valve, the top (outer surface) of the sensor is exposed to the aortic pressure, and the bottom (inner surface) of the sensor is exposed to the ventricular pressure. Therefore, the differential pressure signal is approximately equal to the difference between the aortic pressure and the ventricular pressure. Other sensors can be used, such as optical sensors or columns filled with fluid.

[0066] Figure 2A process 200 for determining the cardiac performance of a heart is illustrated. The process includes a series of steps that involve modifying the operation of a pump within the patient's heart (e.g., checking) in order to compare monitored hemodynamic parameters and thereby calculate a metric indicative of the cardiac performance of the heart (e.g., CO). For example, the process described below may increase the speed of the pump for a short period of time and then compare the aortic pressure during the time of increased pump speed to the aortic pressure during normal pump operation to calculate or characterize vascular resistance or compliance, which may be used to determine CO and / or modify the operation of the pump to better treat the patient. By constructing a system of two equations (one for normal operation of the pump and operation of the pump with increased speed), the CO can be calculated using a Windkessel model (as described below in Figure 5 ) or other nonlinear time-dependent models to determine vascular resistance or compliance, the two equations can be solved using measured or estimated pressure and flow values ​​to calculate resistance and compliance values ​​for the systemic vasculature. Figure 1 Process 200 may be performed using cardiac pump system 100 or any other suitable mechanical circulatory support system.

[0067] In step 202, a pump (e.g., Figure 1 The pump 102 is positioned within the patient's heart. In some embodiments, the pump is an intravascular blood pump device that is placed within the patient's heart via percutaneous insertion. In some embodiments, the pump can be a surgically implanted device, a left ventricular assist device, a counterpulsation device, an expandable heart pump, or any other suitable device. The pump can be introduced to the patient because the patient is in cardiogenic shock or is otherwise experiencing a decline in health. The pump can be positioned across the aortic valve so that the blood inlet of the pump (e.g., Figure 1 The blood inlet 172 of the pump is in the left ventricle and the pump outlet (e.g., Figure 1 The outlet opening 170) is within the aorta.

[0068] The pump contributes to native heart operations by: CO=i h +i p (1) where CO is total cardiac output, i h is the native cardiac output, and i p is the flow contributed by the pump.

[0069] In step 204, while operating the pump at the first pump speed, a hemodynamic parameter is monitored. A hemodynamic parameter may be any parameter related to blood flow within the body. For example, the hemodynamic parameter may include at least one of the following: heart rate, blood pressure, arterial oxygen saturation, mixed venous oxygen 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 gradient, left ventricular end 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 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, arteriovenous oxygen content difference, oxygen delivery, oxygen delivery index, oxygen consumption, oxygen consumption index, oxygen uptake quantification, oxygen uptake index, total peripheral resistance, CO, cardiac index, and CPO. The pump speed is the operating speed of the pump and corresponds to the amount of blood flow provided by the operation of the pump. In some embodiments, the pump speed may correspond to the rotational speed of the rotor. For example, the 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. The pump speed can correspond to a power level or P level, as described above with respect to Figure 1 For example, the pump speed can be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, or P-9 or any other suitable value. In some embodiments, the pump speed can be modified to correspond to the rate at which the chamber of the pump fills and releases blood.

[0070] In step 206, a first phase of a first heart beat of the heart is identified. For example, the first phase can be systole, diastole, or any other suitable phase. The first phase of the first heart beat is identified from the shape of a hemodynamic parameter waveform. For example, the hemodynamic parameter can be aortic pressure. Process 200 includes: identifying local minima in the aortic pressure waveform; and determining a dicrotic notch from these local minima, the onset of the dicrotic notch indicating the onset of diastole. The first phase of the first heart beat occurs during a first time period. For example, the first time period can be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length of time.

[0071] In step 208, a second phase of the second heart beat is predicted based on the monitored hemodynamic parameters. For example, the second phase can be systole, diastole, or any other suitable phase; the second phase can be the same phase as the first phase (e.g., diastole). The second phase is predicted by the following steps: monitoring the hemodynamic parameters over time; and determining a pattern in the hemodynamic parameters to anticipate when the second phase of the heart beat cycle will begin. In some embodiments, the second phase prediction can be further based on the identified first phase of the heart beat cycle. For example, if the first phase is diastole of the first heart beat and the second phase is diastole of the second heart beat immediately after the first heart beat, the second phase can be anticipated by the following steps: determining the average length of the heart beat; and calculating the start of the second phase by adding the length of the heart beat to the start time of the first phase. The second phase of the second heart beat occurs within a second time period. For example, the second time period can be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length of time. By estimating when the second heart beat (and subsequent heart beats) will begin, the system can time changes in pump speed so that their effects (e.g., increasing flow from the left ventricle, increasing aortic pressure) occur during the desired second phase of the second heart beat.

[0072] In an example, the first phase is diastole of the first heart beat, and the second phase is diastole of the second heart beat. In another example, the first phase is systole of the first heart beat, and the second phase is systole of the second heart beat. In an example, the first phase is diastole of the first heart beat, and the second phase is systole of the first heart beat.

[0073] In step 210, the pump speed is changed so that the pump operates at a second pump speed during the second phase of the cardiac cycle to "check" the cardiac beat during the second phase. The pump speed may be increased or decreased. Figure 3 As shown in and described below, the pump speed may be increased during the diastolic phase—that is, the first phase may be the contraction of the first heart beat, and the second phase may be the diastole of the first heart beat. To achieve the change in pump speed, a controller (e.g., Figure 1 The controller 104 may send a signal to the pump to change the pump speed before the start of the second phase so that the pump speed changes during the second phase to account for any time delay between sending the signal and the actual change in pump speed. Changing the pump speed within a single heart beat ensures that there is little noise or external factors affecting the collection of hemodynamic data between the first phase and the second phase.

[0074] After the "check", the pump speed is changed. In some embodiments, after the second phase of the second heart beat, the pump speed is changed back to the first pump speed. For example, the heart pump may be operated at the second pump speed only during the systolic or diastolic phase, and then restored to the first pump speed at or during that phase.

[0075] In step 212, while the ping occurs, a hemodynamic parameter is monitored during the second phase of the second heart beat. For example, the heart pump system may continuously monitor aortic pressure or any other hemodynamic parameter. In step 214, the monitored hemodynamic parameter during the first phase is compared to the monitored hemodynamic parameter 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 blood volume and the second blood volume may be calculated to quantifiably compare the hemodynamic parameter during the first phase of the first heart beat and the hemodynamic parameter during the second phase of the second heart beat.

[0076] In step 216, a metric indicative of cardiac performance of the heart is calculated based on the change in the hemodynamic parameters between the first phase and the second phase. For example, the metric indicative of cardiac performance may be systemic resistance, cardiac compliance, CO, CPO, stroke volume, stroke work, ejection fraction, cardiac contractility, ventricular elastance, cardiac index, prediction of patient survivability. For example, the numerical difference between a first blood volume pumped by the heart during the first phase of the cardiac cycle and a second blood volume pumped by the heart during 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 each heart beat or the average cardiac flow (CO) over a desired time period. Many metrics indicative of cardiac performance are interrelated. For example, CO is determined based on the flow rate of blood passing through and through the pump. Stroke volume is an index of left ventricular function, and its formula is SV=CO / HR, where SV is stroke volume, CO is cardiac output, and HR is heart rate. Stroke work is the work done by the ventricle to eject a certain volume of blood and can be calculated from the stroke volume according to the equation SW=SV*MAP, where SW is the stroke work, SV is the stroke volume, and MAP is the mean arterial pressure. Cardiac work is calculated by multiplying the stroke work and the heart rate. CPO is a measure of cardiac function that represents the heart's pumping ability in watts. CPO is calculated using the equation CPO=mAoP*CO / 451, where CPO is the cardiac power output, mAoP is the mean aortic pressure, CO is the 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 blood volume in the ventricle. Other parameters such as chamber pressure, preload state, afterload state, cardiac recovery, flow load state, variable volume load state, and / or cardiac cycle flow state can be calculated from these values ​​or determined via these parameters. In some embodiments, via a two-element Windkessel model of the vascular system (e.g., Figure 5 The process uses a time-varying, nonlinear model of the vascular system and utilizes the device-arterial connection (a well-controlled analog of the ventricle-vascular connection) to continuously determine systemic vascular resistance and compliance and quantify cardiac stroke volume without the need for additional external measurements. In optional step 218, the operation of the pump is adjusted based on the metric indicative of cardiac performance. In some embodiments, the pump speed is increased or decreased based on the metric indicative of cardiac performance.

[0077] Figure 3A graph 300 of pressure versus time for a heart pump system according to certain embodiments is shown. The y-axis of graph 300 represents aortic pressure in mmHg, while the x-axis represents time as a percentage of the length of a heart beat. In particular, graph 300 shows the effect of a check on aortic pressure. t1 represents the time of a first heart beat, and t2 represents the time of a second heart beat after the first heart beat. When the heart pump system is at least partially placed in the patient's heart, time periods t1 and t2 occur. Point 310 represents the peak systolic pressure during the first heart beat, and point 320 represents the peak systolic pressure during the second heart beat. Point 312 represents the dicrotic notch during the first heart beat, and point 322 represents the dicrotic notch during the second heart beat. Diastolic time periods t3 and t4 represent the diastolic period of the first heart beat and the second heart beat, respectively. During time period t1, the pump operates at a first pump speed. During time period t4, the pump operates at a second pump speed greater than the first pump speed.

[0078] At higher pump speeds, the measured aortic pressure and total flow are higher than at lower pump speeds. Thus, during diastole t4 when the pump is operated at a second pump speed greater than the first pump speed, the aortic pressure is higher than during diastole t3 when the pump is operated at the first pump speed. The difference in aortic pressure between diastole t3 and t4 is shown by the shaded area 324. This difference is associated with an increase in flow and CO during the same time period t4.

[0079] Figure 4 Graphs of pressure, motor speed, and flow versus time are shown. The y-axis of the pressure graph 410 represents the aortic pressure in mmHg, the y-axis of the motor speed graph 420 represents the motor speed according to the P-level, and the y-axis of the flow graph 530 represents the flow in mL / s. The x-axis of the graphs 410, 420, 430 represents time as a percentage of the length of the heart beat. For all three graphs, t1 represents the time of the first heart beat, and t2 represents the time of the second heart beat after the first heart beat. Time periods t1 and t2 occur when the heart pump system is at least partially placed in the patient's heart. At time 440, the second heart beat begins. At time 450, the diastole t3 of the second heart beat begins. At time 460, the second heart beat ends.

[0080] Pressure graph 410 is similar to graph 300 described above. Point 410 represents the dicrotic notch of the first heart beat, point 414 represents the dicrotic notch of the second heart beat, and point 416 represents the beginning of the systolic upward beat of the second heart beat. At time 450, corresponding to point 410 of graph 410 (dicrotic notch of the first heart beat), the pump speed increases, as shown in motor speed graph 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 is increased. As can be seen in pressure graph 410 and flow graph 430, during time period t3, when the pump speed is increased to P-6, both flow and pressure increase.

[0081] Figure 5 A Windkessel model 500 is shown. The Windkessel model 500 includes a current source 510, a current source 520, a resistance 530, and a compliance 540. The governing equation for this model is: Where C is compliance, P is pressure, R is systemic resistance, i h is the flow from native heart operation, and i p is the flow from the pump. However, during diastole, the aortic valve is closed, so the only flow through the left ventricle comes from the pump positioned across the valve. By neglecting the cardiac current source and assuming a constant pump flow, the model can be simplified as follows: Where P0 is the initial aortic pressure during diastole. Resistance and compliance can then be determined via the following two equations, where P1 and P2 are the pressure waveforms measured at different pump speeds: At low pump speeds, i p1 R can be approximated to zero, resulting in a simple exponential of equation (5). After determining R using equation (6) and this simplification, i p1 The R term is added back to equation (5) to then accurately determine C.

[0082] Therefore, by measuring the difference in aortic pressure induced by changes in heart pump speed and potentially assuming that the vascular state remains stable over this interval, the vascular state can be determined by analyzing the aortic pressure waveform measured by the heart pump system. By using the above equation at two different Impella operating points and using the difference in the estimated Impella flow rate, the systemic vascular resistance is determined. The heart performance can then be determined by the following steps: using these vascular state values ​​in the above general equation and using the measured aortic pressure to calculate the flow from the heart. The pulsed jet component of the flow rate waveform is numerically integrated within the jet phase of the heart beat cycle to estimate the stroke volume or CO.

[0083] Figure 6 A CO sensor 610 is shown coupled to a patient 600, wherein the CO sensor is configured to determine native cardiac output. The CO sensor 610 may include a variety of hardware elements configured to perform the methods described herein. In some embodiments, the CO sensor includes an intravascular blood pump (e.g., Figure 1 The intravascular blood pump may be configured to be at least partially placed in the patient's heart. In some embodiments, the intravascular blood pump includes: a cannula; a rotor configured to rotate in a blood vessel and pump blood through the cannula; and a drive mechanism configured to impart power to rotate the rotor. In some embodiments, the cannula may be configured to extend across the aortic valve so that the distal end of the cannula is in the left ventricle and the proximal end of the cannula is in the aorta. For example, when the cannula is placed across the aortic valve so that the blood inlet of the pump is in the left ventricle and the outlet of the pump is in the aorta, the heart pump system may be considered to be "in place". The drive mechanism may include an onboard motor, a drive cable, a drive shaft, or any other suitable element or combination thereof.

[0084] In some embodiments, the CO sensor 610 includes 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 embodiments, the blood pump includes a pump housing and a 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 to flow into the cannula.

[0085] The CO sensor 610 includes a pressure sensor configured to detect a pressure within a blood vessel caused at least in part by the pumping of blood within the blood vessel. For example, the pressure sensor may be an optical pressure sensor as 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, a second side or surface of the differential pressure sensor may be exposed to ventricular pressure, and the differential pressure sensor may measure the difference between the aortic pressure and the ventricular pressure. As another example, the pressure sensor 612 may include a pressure measurement lumen configured to measure aortic pressure.

[0086] The CO sensor 610 includes a controller 614. The controller 614 is coupled to the pressure sensor 612. The controller 614 may be coupled to the pressure sensor 612 directly or indirectly. For example, the controller 614 may be connected to the pressure sensor 612 via electrical wiring, wireless signals, or any other suitable means. The controller 614 is configured to detect a signal indicative of blood pressure from the pressure sensor. All or a portion of the controller 614 may be in a controller unit separate from / remote from the intravascular blood pump. In some embodiments, the control system is internal to the intravascular blood pump.

[0087] In some embodiments, the controller 614 is configured to calculate CO based on a nonlinear model relating CO to vascular resistance and compliance. For example, the nonlinear model may be as described above with respect to Figure 5 Windkessel model described.

[0088] Figure 7 A process 700 for determining CO is shown. Figure 1 Process 700 is performed using the heart pump system 100 or any other suitable pump. In some embodiments, the pump is an intravascular blood pump device that is placed in the patient's heart via percutaneous insertion. The pump may be introduced to the patient because the patient is in cardiogenic shock or is otherwise experiencing a decline in health. The pump may be positioned across the aortic valve so that the blood inlet of the pump (e.g., Figure 1 The blood inlet 172 of the pump is in the left ventricle and the pump outlet (e.g., Figure 1 The outlet opening 170) is in the aorta. The pump contributes to native heart operation such that: CO=i h +i p (1) where CO is total cardiac output, i h is the native cardiac output, and i p is the flow contributed by the pump.

[0089] At step 702, a first aortic pressure wave is detected. The first aortic pressure wave reflects multiple beats of the heart, each reflected beat including a dicrotic notch. The pressure waveform may be measured via a pressure sensor. In some embodiments, the pressure sensor may be located on the pump. In some embodiments, the pressure sensor may be located external to the pump. The pressure sensor may communicate with a controller configured to control the operation of the pump.

[0090] At 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. At 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 its dicrotic notch during the second beat. The first pumping rate is different from the second pumping rate.

[0091] At step 708, a second aortic pressure wave of the heart is detected during the second beat. At step 710, the second aortic pressure wave is compared to a portion of the first aortic pressure wave corresponding to the second beat to detect a change in the second aortic pressure wave. The change between the first aortic pressure wave and the second aortic pressure wave can be used to identify resistance and compliance of the systemic vasculature.

[0092] At step 712, CO is determined based on a nonlinear transfer function that relates CO to systemic resistance and compliance. The transfer function may further relate to the aortic pressure waveform. In some embodiments, the nonlinear transfer function includes a Windkessel model, such as described above with respect to Figure 5 Windkessel model described.

[0093] Figure 8 A process 800 for determining changes in hemodynamic parameters between heart beats is illustrated. At step 802, a mechanical circulatory support device is positioned within the patient's vasculature. In some embodiments, the device is an intravascular blood pump device that is placed within the patient's heart via percutaneous insertion. The device may be introduced to the patient because the patient is in cardiogenic shock or is otherwise experiencing a decline in health. The device may be a left heart device or a right heart device. In some embodiments, the device is positioned across the aortic valve so that the blood inlet of the device (e.g., Figure 1 The blood inlet 172 of the device is in the left ventricle and the outlet of the device (e.g., Figure 1 The outlet opening 170) is within the aorta.

[0094] The device is operable to alter a hemodynamic parameter in a patient. For example, operation of the device may affect aortic pressure in the patient by pumping blood from the left ventricle into the aorta. The device is operated at a first output level and while the heart is beating. The first output level corresponds to a first blood flow rate of native blood flow contributed to the patient by the mechanical circulatory support device during operation of the device at the first output level. For example, the first output level may be associated with a first motor speed, such as the P level described above.

[0095] The device operates at a first output level during a time period including the period of the first heart beat, and monitors the patient's hemodynamic parameters during operation of the device, so that the results of the monitoring are determined as a function of time within each heart beat, and the results are stored in the device memory (or other data storage device). As described above, a hemodynamic parameter is any parameter related to blood flow within organs and tissues of the body. At step 804, the hemodynamic parameter is detected during the first heart beat, and the hemodynamic parameter is consistent with the first heart beat in time. The device output level and hemodynamic parameter measurements during the heart beat (or any other time during the first output level) are consistent with events of the heart beat cycle (e.g., systole, diastole, dicrotic notch). As a result, the hemodynamic parameter and the device output level can be related to cardiac cycle events at various time points during the heart beat. For example, it can be easily detected that a pump operating at a first output level will have a first measured hemodynamic parameter (e.g., aortic pressure) at or after the dicrotic notch of the first heart beat. In some embodiments, the hemodynamic parameter is aortic pressure, and the mechanical circulatory support device includes a pressure sensor configured to detect aortic pressure. In some adaptations, the pressure sensor is included on a cannula that extends partially within the patient's left ventricle.

[0096] At step 806, the device is operated so that it outputs a second output level during a second time period (including during one or more time periods within the second heart beat). The second output level (delivered during the second heart beat) may be greater than or less than the first output level (delivered during the first heart beat). For example, the second output level may be associated with a second motor speed or P level that is greater than or less than the first motor speed, and the output level may be delivered during the second heart beat at the same phase point as the first output level (e.g., at or after the dicrotic notch).

[0097] At step 808, a hemodynamic parameter is detected during the second heart beat (during the period of the second output level) at the same or nearly the same point in the cardiac phase of the second heart beat as in the first heart beat. The hemodynamic parameter may be measured during the entire first heart beat or the second heart beat or within a portion of the respective beats. For example, the hemodynamic parameter may be measured during systole or diastole of the second heart beat or at the dicrotic notch.

[0098] At step 810, a hemodynamic parameter measured during a first heart beat is compared to a hemodynamic parameter measured during a second heart beat. The two measurements are obtained at approximately the same point in the cardiac cycle, albeit during two different beats. The difference in hemodynamic measurement is due to a change in pump speed between the first heart beat and the second heart beat. 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 a first output level to a second output level is related to the contribution of the mechanical circulatory support device to the change in total cardiac output.

[0099] Fig. 9 A process 900 is illustrated for determining cardiac output through a heart beat "check" process. Figure 1 Process 900 is performed using a cardiac pump system 100 or any other suitable pump. The pump is placed in the patient's heart via percutaneous insertion. The patient may be in cardiogenic shock or otherwise experiencing a decline in vascular health. The pump may be a left heart device or a right heart device. The pump is positioned across the aortic valve so that the blood inlet of the pump (e.g., Figure 1 The blood inlet 172 of the pump is in the left ventricle and the pump outlet (e.g., Figure 1 The outlet opening 170) is in the aorta. The pump contributes to native heart operation such that: CO=i h +i p (1) where CO is total cardiac output, i h is the native cardiac output, and i p is the flow contributed by the pump.

[0100] At step 902, a pump is operated at a first pump speed during a first time period including a period of a first heart beat. At step 904, during a first diastole of the first heart beat, a hemodynamic parameter is monitored during operation of the heart pump at the first pump speed. The hemodynamic parameter relates to blood flow within the body. The pump speed is the operating speed of the pump and corresponds to the amount of blood flow provided by the operation of the pump. In some embodiments, the pump speed corresponds to the rotational speed of the pump rotor. For example, the pump speed may be at or above 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. The pump speed may correspond to a power level or P level, as described above with respect to Figure 1 As described. For example, the pump speed can be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8 or P-9 or any other suitable value. In some embodiments, the pump speed corresponds to the rate at which the chamber of the pump fills and releases blood. By monitoring the hemodynamic parameters, the systems and methods described herein can identify changes in the hemodynamic parameters over time, including during the stages of the first heart beat and the second heart beat. Such comparisons can be used to quantify cardiac performance (e.g., by CO), as discussed more fully herein.

[0101] At step 906, a first operating parameter of the intravascular blood pump during diastole is determined. For example, the operating parameter can be an electric current supplied to the pump, a blood flow rate provided by the pump, or a placement of the pump within the patient's vasculature. Specifically, determining the first operating parameter can include determining a first blood flow rate provided by the blood pump during diastole. The first operating parameter and the measured hemodynamic parameter can be identified at a specific point in the cardiac cycle of a first heart beat. The flow from the pump is estimated based on a motor current supplied to a motor in the blood pump to maintain the pump speed.

[0102] For a given intravascular blood pump system, the flow output i can be determined by the speed of the pump (revolutions per minute or RPM) and the motor current supplied to the pump to maintain operation at that pump speed. p. This mathematical calculation from pump speed and motor current to flow rate can be implemented by setting up a lookup table, where the pump speed and motor current are indices into the table, and the flow values ​​in the table are pre-filled by bench testing. Another way is to determine the flow rate for a subset of 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 the 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.

[0103] At step 908, the first pump speed is changed to a second pump speed so that the operation of the heart pump delivers a second output level during the second diastolic period of the second heart beat. The second pump speed may be greater than or less than the first pump speed. In some embodiments, the pump speed increase is timed so that the increased pump speed occurs during the predicted period of the heart beat. For example, the start of the speed increase may be synchronized with the start of diastole to account for any delay between sending an instruction to the pump to change the speed and when the change in speed actually occurs. The end of the speed increase may be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, the peak systolic pressure, or any other suitable time. In some embodiments, the system is configured so that the pump speed is increased or decreased within a set time period. For example, the speed change may last for about 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other suitable length of time. The second heart beat is different from the first heart beat. The hemodynamic parameter is measured during the second heart beat, which may be done at the same point (such as the dicrotic notch) in the beat as in the first heart beat when the first hemodynamic parameter is measured. In some embodiments, the detection and measurement are applied to a second heart beat that occurs after the first heart beat.

[0104] At step 910, a hemodynamic parameter is monitored during a second diastolic period of a second heart beat (e.g., at the dicrotic notch). At step 912, a second operating parameter of the intravascular blood pump during the second diastolic period is determined. Determining the second operating parameter may include determining a second blood flow rate (or a second level of a motor operating parameter) provided by the blood pump during the second diastolic period.

[0105] At step 914, a metric indicative of cardiac performance of the heart is calculated. The metric is based on (i) a first operating parameter, (ii) a second operating parameter, and (iii) a hemodynamic parameter during a first diastole and a second diastole (e.g., at the dicrotic notch during both periods). The metric can be used in a transfer function or system of equations, such as the transfer function or system of equations described above with respect to the Windkessel model. In some embodiments, a mathematical representation of the hemodynamic parameter is determined for the first diastole and the second diastole. For example, the mathematical representation can be a sum of sinusoids.

[0106] The metrics are used to construct waveforms that can be used to determine cardiac output. Calculating cardiac performance may include: deconstructing a first waveform representing hemodynamic parameters for a first diastole to determine a first set of sinusoids; and deconstructing a second waveform representing hemodynamic parameters for a second diastole to determine a second set of sinusoids. The deconstructions may include applying a Fourier transform to the first waveform, the second waveform, or both. A set of sinusoids may include one or more sinusoids added together.

[0107] When the pump is operated within the patient's vasculature, the blood flow in the aorta is equal to the pump contribution (i p ) plus native heart contribution (i h The first set of sinusoids and the second set of sinusoids may be compared to determine the contribution of the patient's heart to blood flow in the aorta (i h ). For example, aortic pressure may be a hemodynamic parameter, and aortic pressure may be expressed as the sum of sinusoids generated by Fourier transform, such as: Where P is the aortic pressure, f n is the frequency associated with the operating parameter, and A n and θ n are coefficients of the operating parameter. Since the change in pressure between operating parameters will be proportional to the change in flow, the difference in each set of sinusoids between operating parameters can be used to calculate the difference in flow between operating parameters. In some embodiments, the Fourier transform can be calculated for each pump speed in a range of pump speeds. In some embodiments, the patient's response to the "tested" pump speed may be minimal due to the limitation of speed changes within a short period of time (i.e., the time it takes to ramp up the pump to an increased speed or to decelerate the pump to a decreased speed).

[0108] Decomposing the hemodynamic parameter over time into its component frequencies allows the hemodynamic parameter to be determined using a mathematical equation or set of equations. In some embodiments, the mathematical representation is an exponential equation based on a comparison of sinusoids. After the hemodynamic parameter waveform has been characterized by the mathematical equation, cardiac parameters such as vascular resistance and compliance can be determined from the equation. For example, if the hemodynamic parameter waveform is characterized as (where B is equal to R*C, and D is equal to i p *C, where P is pressure, R is systemic resistance, and C is systemic compliance), systemic resistance and compliance values ​​can be calculated by solving a system of equations with these coefficients at two points in time (i.e., with two known pressure measurements).

[0109] In some embodiments, a model heart beat representing the patient's cardiac function can be simulated based on a comparison of sinusoids, and the model heart beat is used to determine cardiac output, determine when to apply mechanical circulatory support, and what level. For example, a blood pump can operate at a series of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), where each pump speed corresponds to the rotation rate of the rotor in the pump and a similar 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) will change the value of the hemodynamic parameter because it will change the blood flow in the vascular system provided by the operation of the pump. By gradually single-stepping multiple pump speeds (or operating parameters, such as the blood flow provided by the pump) to identify the corresponding changes in one or more hemodynamic parameters, forming a hemodynamic waveform, and deconstructing the hemodynamic waveform generated by each pump speed, the relationship between pressure and flow during diastole is established. The patient's overall cardiac function can then be mapped into a mathematical representation (as a function of measured hemodynamic parameters) that can be used to simulate future cardiac function and inform the delivery and control of mechanical circulatory support to the patient.

[0110] For example, the method described below may be used to construct a measured aortic pressure waveform for any recorded heart beat - allowing CO to be calculated for that heart beat.

[0111] As described above, in some embodiments, a brief change in pump speed can be applied to the pump within one heart beat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heart beat can be compared to the aortic pressure of a heart beat without this brief speed change or impulse stimulus. The difference between the two (the aortic pressure of the altered heart beat and the aortic pressure of the "normal" heart beat) can be considered the impulse response of the aortic pressure: Δp(t)=p1(t)-p2(t) Wherein P1(t) is the pressure waveform measured with pulse stimulation, P2(t) is the pressure waveform without pulse stimulation, and ΔP(t) is the pulse response of aortic pressure.

[0112] If this pulse stimulus is applied only during diastole, the difference in total cardiac flow for the two heart beats can be expressed as: Δi(t)=i1(t)-i2(t) where i1(t) and i2(t) are the pump flows for heart beats with and without pulse stimulation, respectively, and Δi(t) is the impulse response of the heart flow.

[0113] The relationship of aortic pressure to pump flow can then be estimated in the frequency domain as: Where ΔP(f) is the frequency domain representation (eg, Fast Fourier Transform or FFT) of Δp(t), ΔI(f) is the frequency domain representation of Δi(t), and H(f) is the frequency domain transfer function of aortic pressure versus pump flow.

[0114] Once this relationship H(f) is established as outlined above, the total cardiac flow for any heart beat with the aortic pressure measured as p(t) can be calculated as: where P(f) is the frequency domain representation of p(t) and IFFT is the Inverse Fast Fourier Transform.

[0115] Fig.10 Two graphs are shown for the same ten second period, one graph of aortic pressure and one graph of cardiac flow. The y-axis of the upper graph represents aortic pressure in mmHg, while the x-axis represents time in seconds. The y-axis of the lower graph represents the calculated total cardiac flow in liters per minute, while the x-axis represents time in 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 obtained during the depicted ten second time period in combination with pump data as described above. The total cardiac flow i is calculated using R, C, and the aortic pressure waveform by applying equation (2) h +i p : The total cardiac flow i generated by equation (1) over a period of time (e.g., 5 seconds, 10 seconds, or 30 seconds) can be calculated by taking the total cardiac flow i h +i pThe average value of CO is used to calculate CO. Figure 7 In the example in , the time period is 10 seconds. The average R value for this time period is 0.6143 mmHg*sec / ml, and the average C value is 1.5 mL / mmHg, resulting in a calculated CO of 6.9 L / min.

[0116] The foregoing is merely illustrative of the principles of the present disclosure, and the device may be practiced in other aspects than those described, which are presented for purposes of illustration and not limitation. It will be understood that the device disclosed herein, although shown for percutaneous insertion of a heart pump, may be applied to devices in other applications requiring hemostasis.

[0117] After reviewing this disclosure, variations and modifications will occur to those skilled in the art. 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 components thereof) may be combined or integrated in other systems. In addition, certain features may be omitted or not implemented.

[0118] The described systems and methods may be implemented locally on a heart pump system or a controller of a 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 implemented remotely on an independent data processing device. The independent 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 independent data processing device in real time (or near real time).

[0119] In general, aspects of the subject matter and functional operations described in this specification may be implemented in digital electronic circuits or in computer software, firmware or hardware, including the structures disclosed in this specification and their structural equivalents, or in a combination of one or more of them. Aspects of the subject matter described in this specification may 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 a data processing device or to control the operation of a data processing device. The computer-readable medium may be a machine-readable storage device, a machine-readable storage substrate, a memory device, a material combination that affects a machine-readable propagation signal, or a combination of one or more of them. The term "data processing device" encompasses all devices, devices, and machines for processing data, including, as examples, a programmable processor, a computer, or multiple processors or computers. In addition to hardware, a device may also include code that creates an execution environment for the computer program in question, for example, code that constitutes a processor firmware, a protocol stack, a database management system, an operating system, or a combination of one or more of them. A propagated signal is an artificially generated signal, for example, a machine-generated electrical, optical, or electromagnetic signal that is generated to encode information for transmission to a suitable receiver device.

[0120] A computer program (also referred to as a program, software, software application, script, or code) may be written in any form of programming language, including compiled or interpreted languages, and it may be deployed in any form, including 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 may be stored in a portion 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 in question, or in multiple coordinated files (e.g., files that store multiple portions of one or more modules, subroutines, or code). A computer program may be deployed to execute on one computer, or on multiple computers located at one site or distributed across multiple sites and interconnected by a communication network.

[0121] The processes and logic flows described in this specification can be performed by one or more programmable processors executing one or more computer programs to perform functions by operating on input data and generating output. The processes and logic flows can also be performed by special purpose logic circuits, and the device can also be implemented as special purpose logic circuits, such as FPGAs (field programmable gate arrays) or ASICs (application specific integrated circuits).

[0122] As an example, processors suitable for executing computer programs include both general-purpose microprocessors and special-purpose microprocessors, and any one or more processors of any kind of digital computer. Typically, the processor will receive instructions and data from a read-only memory or a random access memory or both. The basic elements of a computer are a processor for executing instructions and one or more memory devices for storing instructions and data. Typically, a computer will also include one or more mass storage devices (e.g., magnetic disks, magneto-optical disks, or optical disks) for storing data, or be operatively connected to receive data from the mass storage device or transfer data to the mass storage device. However, a computer need not have such a device.

[0123] Examples of changes, substitutions, and alterations are ascertainable by those skilled in the art and may be made without departing from the scope of the information disclosed herein.All references cited herein are incorporated by reference in their entirety and constitute a part of this application.

Claims

1. A controller, configured to: operating a pump at a first pump speed, the pump being configured to be positioned within the heart and configured to operate at an adjustable pump speed; monitoring a hemodynamic parameter during pumping at the first pump speed; identifying a first phase of a first heart beat of the heart during a first time period; predicting a second phase of a second heart beat of the heart within a second time period based on the monitored hemodynamic parameters; changing the pump speed to a second pump speed during the second phase of the second heart beat; monitoring said hemodynamic parameters during said second stage; comparing a monitored hemodynamic parameter during the first phase with a monitored hemodynamic parameter during the second phase to calculate a change in the hemodynamic parameter between the first phase and the second phase; as well as A metric indicative of vascular performance is calculated based on a change in the hemodynamic parameter between the first phase and the second phase.

2. The controller according to claim 1, wherein: The controller is further configured to change the second pump speed to the first pump speed after the second stage.

3. A controller according to any one of the preceding claims, wherein: The hemodynamic parameter is aortic pressure.

4. A controller according to any one of the preceding claims, wherein: The first phase is one of systole and diastole, and the second phase is one of systole and diastole.

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

6. The controller according to claim 4, wherein: The first phase is a first contraction period, and the second phase is a second contraction period.

7. A controller according to any one of the preceding claims, wherein: Identification of the first phase of the first heart beat is based on changes in a monitored hemodynamic parameter over time while pumping at the first pump speed.

8. A controller according to any one of the preceding claims, wherein: The prediction of the second phase of the second heart beat of the heart is based on the identified first phase of the heart beat cycle.

9. A controller according to any one of the preceding claims, wherein: In comparing the hemodynamic parameter during the first phase with the hemodynamic parameter during the second phase, the controller is further configured to: calculating a first volume of blood pumped by the heart during the first stage; calculating a second volume of blood pumped by the heart during the second phase; as well as A numerical difference between the first blood volume and the second blood volume is determined.

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

11. A controller according to any one of the preceding claims, wherein: The cardiac performance of the heart is determined by: calculation of vascular compliance and vascular resistance of systemic vasculature based on changes in said hemodynamic parameters between said first phase and said second phase; and Calculation of cardiac output of the heart using the Windkessel model.

12. A controller according to any one of the preceding claims, wherein: The first pump speed is less than the second pump speed.

13. A controller according to any one of the preceding claims, wherein: The first pump speed is greater than the second pump speed.

14. A controller configured to: detecting a first aortic pressure wave of the heart reflecting a plurality of beats of the heart, each reflected beat including a dicrotic notch; applying hemodynamic support to the heart at a first pumping rate during a first beat of the plurality of beats; regulating the hemodynamic support to the heart during a second beat of the plurality of beats by providing a second pumping rate to the heart following a dicrotic notch thereof during the second beat; detecting a second aortic pressure wave of the heart during the second beat; comparing the second aortic pressure wave to a portion of the first aortic pressure wave corresponding to the second pulse to detect changes in the second aortic pressure wave and identify resistance and compliance of the vasculature; as well as The native cardiac output of the heart is determined based on a non-linear transfer function programmed within the software that relates cardiac output to vascular resistance and compliance.

15. The controller according to claim 14, wherein: The nonlinear transfer function includes a Windkessel model.

16. A controller according to any one of the preceding claims, wherein: Hemodynamic support is provided by an intracardiac blood pump having a cannula configured to be positioned across the aortic valve.

17. A controller according to any one of the preceding claims, wherein: The controller is further configured to adjust the hemodynamic support based on at least one of: the determined cardiac output, the resistance, or the compliance.

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