Systems and methods for determining cardiac performance

Abstract

Determines cardiac performance. [Solution] The systems and methods described herein determine indicators of cardiac performance via a mechanical circulatory support device and use cardiac performance to calibrate, control, and provide mechanical circulatory support for the heart. The system includes a controller configured to operate the device, receive inputs representing device operating conditions and hemodynamic parameters, and determine vascular performance, including vascular resistance and compliance, as well as intrinsic cardiac output. The systems and methods operate to introduce controlled disturbance in the vascular system by using a mechanical circulatory support device (e.g., a cardiac pump) and, in response, determine cardiac parameters such as stroke volume, vascular resistance and compliance, left ventricular end-diastolic pressure, and ultimately determine intrinsic cardiac output.

Description

[Technical Field]

[0001] Cross-reference of related applications This application claims priority and benefits of U.S. Patent Provisional Application No. 62 / 687,133, entitled “METHODS AND SYSTEMS FOR IMPROVED ASSESSMENT OF VASCULAR AND CARDIAC STATE,” filed on 19 June 2018; U.S. Patent Provisional Application No. 62 / 863,136, entitled “SYSTEMS AND METHODS FOR SYSTEM IDENTIFICATION,” filed on 18 June 2019; and U.S. Patent Provisional Application No. 62 / 863,146, entitled “SYSTEMS AND METHODS FOR DETERMINING CARDIAC PERFORMANCE,” filed on 18 June 2019. The entire contents of these applications are incorporated herein by reference. [Background technology]

[0002] background Cardiovascular disease is a leading cause of illness, death, and the burden on healthcare worldwide. A wide range of treatments have been developed for heart health, from pharmaceuticals to mechanical devices and implants. Temporary cardiac support devices, such as cardiac pump systems, provide hemodynamic support and promote cardiac recovery. Some cardiac pump systems are inserted percutaneously into the heart and can operate in parallel with the natural heart to supplement cardiac output. An example of such a device is the IMPELLA® family of devices (Abiomed, Inc., Danvers, Massachusetts). Such cardiac pump systems have sensors that detect blood pressure (or assess differential pressure across a membrane) and can monitor motor current, using sensor and motor current readings to help locate the pump.

[0003] The cardiac support a given patient needs can vary from patient to patient. Cardiac output (CO) is the volumetric flow rate of blood pumped by the heart. Normal cardiac output is approximately 5 L / min in a healthy adult, but can vary depending on various factors, including the physical composition of a given patient. Quantitatively determining, using known techniques, how much cardiac output a given heart produces, how much additional support a device should provide, when and for how long, is difficult for clinicians. This determination can be particularly difficult for patients recovering from intervention or other cardiac treatments. Therefore, clinicians tend to rely on assessments and indirect estimates of cardiac function, such as measuring intracardiac or intravascular pressure using fluid-filled catheters. In particular, quantifying cardiac output (CO) is difficult. One approach is to use a pulmonary artery catheter (PAC) to measure central venous and pulmonary artery pressure in real time. The PAC relies on estimating CO using Fick's law or bolus thermodilution by measuring systemic oxygen consumption. However, due to the assumptions that must be made to arrive at CO metrics, and the corresponding lack of precision, PAC may have limitations in its use in complex interventions and high-risk situations such as cardiogenic shock. Measurements by PAC do not take into account the dynamic changes in cardiac function and are not continuous, while at the same time, they may not be able to address the nonlinear aspects of systemic ventricular vascular connections. [Overview of the project]

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

[0005] The system and method consist of a time-variant nonlinear model of the vascular system, using a device-arterial correlation to continuously determine systemic vascular resistance and compliance, thereby quantifying cardiac stroke volume. In some implementations, the system and method improve upon traditional linear approximations using a Windkessel model of the vascular system, providing dynamic variations in vascular response. In some embodiments, the system and method are configured as a cardiac output sensor capable of directly determining the intrinsic cardiac output of a patient's heart.

[0006] In various applications, the system and method are configured to "ping" the vascular system during a heartbeat using a mechanical circulatory assist system and then detect a response by the heart at one or more subsequent periods or time points, e.g., at a subsequent heartbeat. A "ping" involves increasing or decreasing the output of the mechanical circulatory assist system over a short period, e.g., during a single heartbeat (e.g., increasing or decreasing the pump speed of a heart pump system), thereby creating a spike in blood pressure and blood flow (e.g., aortic pressure and blood flow out of the left ventricle). The ping causes a change in hemodynamic parameters (e.g., aortic pressure) from a baseline, this change is detected, and compared to hemodynamic parameters at another time point (e.g., aortic pressure when the heart is not being pinged) to determine heart performance. The ping may involve changing the pump speed during a period (or time point) within a part of a single heartbeat (e.g., a certain phase of the heartbeat) and comparing the hemodynamic parameters during this "change" time to hemodynamic parameters during a period or time point of "normal" operation when the ping is not being applied (e.g., during a subsequent heartbeat).

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

[0008] Intracardiac pinging (i.e., adjusting the operation of the pump within a single heartbeat) enables comparison of hemodynamic parameters that occur between multiple heartbeats (adjacent to each other in sequence or separated by other beats), and also minimizes noise (e.g., sympathetic responses) that can occur when the pump speed is changed over a relatively long period. As described above, in some embodiments, comparison of hemodynamic parameters is accomplished by a control system processor programmed with a model of the vascular system, such as a two-element Windkessel model, that models and accounts for the changing non-linear interaction between the flow of the pump and the operation of the heart. The control system uses known terms (received as inputs) to approximate values of the pump flow and the operation of the heart (e.g., aortic pressure), thereby enabling such a model to be easily verified and utilized in clinical applications. The system and method provide metrics representative of the health of the patient's heart, such as the resistance and compliance of the systemic vascular system, which enables determination of CO and other aspects of cardiac performance. In some applications, the system and method are deployed without the need for additional measurements or diagnostic catheters. The ability to continuously and accurately track changes in systemic vascular performance (e.g., resistance or compliance) and estimate the cardiac stroke volume results in a significant advance beyond traditional measurements obtained from PAC or other diagnostics currently deployed in clinical practice.

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

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

[0011] The systems and methods described herein vary the operation of a mechanical circulatory device within a heartbeat to compare one or more monitored hemodynamic parameters in that phase with the same parameters in different heartbeats, thereby calculating an index representing the patient's vascular performance. For example, a pump may be inserted into a blood vessel and operated at a first pump velocity (or other power level) for a series of baseline heartbeats, including a first heartbeat, and then the heart may be "pinged" by increasing or decreasing the pump velocity for a very short period in a second, i.e., target heartbeat or in a specific phase of the target heartbeat (e.g., located at or after the overlapping notch of the target heartbeat). Aortic pressure or other hemodynamic parameters are measured in both the series of baseline heartbeats and the target heartbeat. In some adaptations, the hemodynamic parameters are measured in the same portion of the first and second heartbeats (e.g., during the systole of both heartbeats, or at or after the overlapping notch of both heartbeats). The system then calculates or characterizes vascular resistance or compliance (which can be used to determine CO or modify pump action for better patient treatment) by comparing hemodynamic parameters, such as aortic pressure, identified during a series of baseline heartbeats (e.g., the first heartbeat), with such hemodynamic parameters identified during a period of increased pump speed (e.g., the second heartbeat), for example, by using the same sensor. The series of baseline heartbeats (e.g., including the first heartbeat) may occur after or before the ping. This method can be performed using the cardiac pump system 100 shown in Figure 1, described later, or any other suitable pump.

[0012] To implement the system and method, a pump or other mechanical circulatory support device is placed within the patient's vascular system (e.g., within the patient's heart) and is capable of altering the patient's hemodynamics. For example, the device's operation may increase the patient's aortic pressure by reducing the load on the left ventricle or by other means. In some implementations, the pump is an intravascular blood pump device placed within the patient's heart by percutaneous insertion. The pump may be a surgically implanted device, a left ventricular support device, a counterpulsation device, a deployable cardiac pump, an extracorporeal device, or any other suitable device. The pump may be appropriate because the patient is in a state of cardiogenic shock or otherwise facing impaired vascular health. The pump may be positioned across the aortic valve such that the blood inlet to the pump is located in the left ventricle and the outlet from the pump is located in the aorta. CO=i h +i p (1) This contributes to the normal function of the heart, where CO is the total cardiac output, i h This is the actual cardiac output, i p This is the flow rate due to the pump's contribution. Such pumps can provide life-saving benefits to patients with cardiogenic shock by increasing oxygenated blood flow from the heart to the coronary arteries and other areas of the vascular system.

[0013] Hemodynamic parameters are monitored during the operation of a pump or other mechanical circulatory support system (for example, parameters can be continuously monitored during cardiac performance, and relevant data representing the parameters at a selected point in time during a selected beat can be identified for use). Appropriate hemodynamic parameters include those related to blood flow in the organs and tissues of the body. Exemplary hemodynamic parameters include heart rate, blood pressure, arterial oxygen saturation, mixed venous saturation, central venous oxygen saturation, arterial blood pressure, mean arterial pressure, right arterial pressure, central venous pressure, right ventricular pressure, pulmonary artery pressure, mean pulmonary artery pressure, pulmonary occlusive pressure, left atrial pressure, aortic pressure, differential pressure, left ventricular end-stage 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 resistance index, left heart Includes 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, arterial-venous oxygen content difference, oxygen transport, oxygen transport index, oxygen consumption, oxygen consumption index, oxygen uptake rate, oxygen uptake index, total peripheral resistance, CO, cardiac index, and cardiac output (CPO). In some implementations, differential pressure (P) is used instead of aortic pressure in the diastolic calculation. diff ) can be used (for example, when the differential pressure is known and the aortic pressure is unknown). P diff P is equal to aortic pressure minus left ventricular pressure (LVP). In many cases, LVP is much smaller than aortic pressure during diastole and does not change as much during diastole compared to aortic pressure. In these cases, P diff Since it is sufficiently close to aortic pressure (i.e., LVP can be ignored), it can serve as a substitute when AoP is unavailable. diff If used, it may affect the accuracy of the results of certain calculations described herein.

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

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

[0016] In the second step, the second phase of the heartbeat cycle, which is the target phase in which the heart receives the “ping,” is predicted or identified. For example, the second phase may be a second heartbeat, a different systole, a different diastolic phase, or any other suitable phase, or a combination of phases, provided that the second phase is selected such that the “ping” effect is on the heart for a certain period of time. The system and method may predict when the second phase of the heartbeat cycle will begin based on previously monitored hemodynamic parameters (e.g., aortic pressure measured in the first phase of the heartbeat cycle). The second phase of the heartbeat cycle is predicted to occur over a second period. For example, the second period may be 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any suitable length. The second period may be set to correspond to the length of the second phase of the heartbeat cycle in a particular patient. For example, if the second phase of the heartbeat cycle is diastolic, the second period may be set to the average diastolic period of a particular patient, or to the predicted length of the next diastolic period to occur. In some examples, the second period may be preset to be shorter than the duration of the heartbeat. The second phase is predicted based on monitored hemodynamic parameters and the identified first phase of the heartbeat cycle. The heartbeat signal can be monitored, and based on that monitored signal, the system and method predict when the next diastolic or systolic phase will occur. By predicting the timing of the phases of the heartbeat cycle, the system and method can determine the timing of an increase or decrease in pump speed ("ping") so that it begins (or can influence) precisely when the second phase of the heartbeat cycle begins. For example, the system and method can be configured to ping the heart with a short-term increase in pump output (e.g., by reducing the cardiac load with a higher pump speed) so that the resulting increase in blood flow coincides with the start of a preferred time or period within a target heartbeat, such as at or immediately after the start of a overlapping notch in a subsequent heartbeat, during the systole of a subsequent heartbeat, throughout the diastolic phase, or a predetermined portion of a subsequent heartbeat.

[0017] In certain implementations, the first phase of the heartbeat cycle is the diastolic phase of the first heartbeat, and the second phase of the heartbeat cycle is the diastolic phase of the immediately following second heartbeat. The first phase of the heartbeat cycle may be the systolic phase of the first heartbeat, and the second phase of the heartbeat cycle is the systolic phase of the second heartbeat. In some implementations, the second phase is during a heartbeat that consists of multiple beats removed from the first heartbeat, and in other implementations, the second phase is during a heartbeat adjacent to the first heartbeat.

[0018] After the baseline and target heart rate phases are established, the pump rate is modified to ping the heart, for example, by operating the pump at a second pump rate different from the first pump rate in the second heartbeat. The adjustment of the pump rate may be an increase or decrease in the pump rate to ping the heartbeat. For example, the pump may be adjusted so that the pump rate is temporarily higher during a period of the heartbeat, such as the diastole. Furthermore, the pump may be adjusted so that the rate returns to baseline or is otherwise decreased during the same period of the heartbeat or at some other point in the same heartbeat. It is also possible to configure ping to occur in the opposite way, by temporarily decreasing the pump rate from baseline.

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

[0020] Pump speed is adjusted to produce a change in pump speed at a desired point in the heartbeat. For example, the change in speed can occur during systole, diastole, or both within a heartbeat. In some implementations, to adjust the pump speed, the controller sends a signal to the pump and changes the pump speed before the start of the target phase, allowing time to account for any time delay between the transmission of the control signal and the change in pump speed. The pump must produce an actual increase or decrease in speed at a desired point in the target heartbeat (e.g., diastole, during overlapping notches, or thereafter). Ping is timed to cause a temporary increase in pump speed during a known period of the heartbeat. For example, the start of speed ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other appropriate time. The end of speed ping can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other appropriate time. In some implementations, ping is achieved by increasing or decreasing the pump speed over a set period of time. For example, ping can be synchronized to the start of diastole so that ping occurs during diastole. Alternatively, ping can be synchronized to the end of diastole so that ping occurs during the systole of the next heartbeat. In other adaptations, ping is synchronized to the start of systole, the end of systole, the peak of systolic blood pressure, or any other appropriate time. Ping is configured to last for a set duration. In some adaptations, ping is set to last for a duration corresponding to the length of the heartbeat phase. For example, ping can be set to last for approximately 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other appropriate time.

[0021] Pinging the heart or other vasculature by instantaneously adjusting mechanical circulatory assistance (e.g., pump speed) disturbs the heart or other vasculature, thereby enabling determination of additional indices of cardiac performance, including systemic resistance, compliance, and cardiac output. Such determination can be made without introducing additional hardware into the patient's body (other than the pump providing hemodynamic assistance) (although such additional hardware can still be used if desired). To make the determination, hemodynamic waveforms (e.g., aortic pressure (or intraventricular pressure) waveforms) at normal heartbeats and pinged heartbeats are compared via a non-linear model such as the Windkessel model described below. The change in the pressure waveform (or other hemodynamic parameter) between the normal heartbeat and the pinged heartbeat is reflected in different values within the model during two periods (the period of the baseline (normal) heartbeat and the period of the target (pinged) heartbeat), which generates two model equations. Knowing the pressure waveforms for the two periods allows reduction of the number of unknown variables between the two model equations and calculation of resistance and compliance. Then, using the calculated resistance and compliance values, and the aortic pressure waveform, Equation (2):

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[0022] The above-described technologies can be applied in various ways. In some implementations, hemodynamic parameters are monitored during the second phase of the second heartbeat. For example, the cardiac pump system can continuously monitor aortic pressure or any other hemodynamic parameters. In some implementations, hemodynamic parameters monitored in the first phase are compared with hemodynamic parameters monitored in the second phase. For example, the first blood volume pumped by the heart in the first phase and the second blood volume pumped by the heart in the second phase can be calculated. By calculating the numerical difference between the first and second blood volumes, the hemodynamic parameters in the first phase can be quantitatively compared with those in the second phase. For example, the area under the curve (AUC) of a flow curve can represent blood volume. The difference in AUC between the first and second phases may indicate the difference in blood volume pumped at the first and second pump speeds. Furthermore, comparing hemodynamic parameters between the first and second phases may include evaluating the linearity of changes in hemodynamic parameters. For example, aortic pressure may not change linearly between pump speeds; that is, the change in aortic pressure from one pump speed to the next may not progress linearly. Depending on how the change in aortic pressure progresses between pump speeds, it may be possible to predict the aortic pressure at different pump speeds.

[0023] Indicators representing cardiac performance can be calculated based on changes in hemodynamic parameters between the first and second phases. For example, indicators of cardiac performance can be determined from the different pressure waveforms of the cardiac cycle in the first and second heartbeats, and such indicators may include systemic resistance, systemic compliance, CO, CPO, stroke volume, stroke work, ejection fraction, myocardial contractility, ventricular elastance, cardiac index, and prediction of patient survival. Numerous indicators representing cardiac performance are interrelated. For example, CO is determined based on the flow rate of blood through the intravascular pump located within the patient's heart. Stroke volume is an indicator 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 amount of work done by the ventricles to pump blood, and can be calculated from stroke volume according to the formula SW = SV * MAP, where SW is stroke work, SV is stroke volume, and MAP is mean arterial pressure. Cardiac work is calculated by the product of stroke volume and heart rate. CPO is a measure of cardiac function that represents the heart's pumping capacity, and its unit is watts. CPO is calculated using the following formula: CPO = MAP * CO / 451(3) The cardiac output is calculated using the formula expressed by , where CPO is cardiac output, MAP is mean arterial pressure, CO is cardiac output, and 451 is a constant used to convert mmHg x L / min to watts. Ejection fraction can be calculated by dividing stroke volume by ventricular blood volume. 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 through these parameters. In some implementations, indices representing cardiac performance are calculated via a two-element Windkessel model of the vascular system (e.g., the Windkessel model in Figure 5) to model the dynamic and nonlinear cardiac-vascular interactions. Therefore, this process employs a time-varying nonlinear model of the vascular system and uses the correlation between an intravascular blood pump device, which is a well-controlled analogue of the ventricular-vascular interaction, and the patient's hemodynamic function to continuously determine systemic vascular resistance and vascular compliance and quantify cardiac stroke volume without requiring additional external measurements.

[0024] The pump's operation can be adjusted based on an indicator of cardiac performance. Adjustments to the pump's operation may include increasing or decreasing the pump speed, adjusting the pump placement, turning the pump off, or any other appropriate adjustments. For example, if the indicator of cardiac performance is stroke volume, the pump speed can be increased when the stroke volume falls below a threshold, while it can be decreased when the stroke volume exceeds a threshold.

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

[0026] The CO sensor may include an elongated catheter body coupled to a cannula. The elongated catheter may include a drive cable, electrical wiring connecting the blood pump to a control system, any appropriate elements, or any combination thereof. In some implementations, the pump includes a pump housing and a motor housing, the motor housing being coupled to the cannula at its distal end. A rotor can be rotated within the pump housing to generate blood flow to the cannula.

[0027] CO sensors may include hemodynamic parameter sensors operably positioned on the blood pump (or proximal or distal to the blood pump) and configured to detect intravascular pressure resulting at least partially from the pumping of blood within the vessel. For example, the pressure sensor may be an optical sensor located on or near the pump housing or cannula. In another example, the pressure sensor may comprise a pressure measuring lumen configured to measure aortic pressure. A differential pressure sensor may also be used, with one side or surface of the differential pressure sensor exposed to aortic pressure and a second side or surface of the differential pressure sensor exposed to intraventricular pressure, and the differential pressure sensor may measure the difference between the aortic pressure and the intraventricular pressure.

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

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

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[0030] In some implementations, P0 is,

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[0031] In some cases, mechanical circulatory support can be provided to a patient using a blood pump in accordance with the systems and methods described herein. Providing mechanical circulatory support may include operating a blood pump within the patient's vascular system, 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.

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

[0033] In several implementations, the pump is placed within the patient's heart. The pump may be introduced to a patient if they are in cardiogenic shock, undergoing coronary intervention, experiencing a heart attack, or otherwise facing impaired heart health. The pump can be positioned across the aortic valve, with the blood inlet to the pump located in the left ventricle and the outlet from the pump located in the aorta. The pump contributes to the heart's natural function, ensuring that CO from the heart equals the original CO+ pump output.

[0034] The first aortic pressure wave can be detected. The first aortic pressure wave reflects multiple heartbeats, each reflected beat containing overlapping notches. The pressure waveform can be measured by a pressure sensor. In some implementations, the pressure sensor may be built into the pump. In some implementations, the pressure sensor is located outside the pump and receives fluid or electrical signals. The pressure sensor can communicate with a controller configured to control the operation of the pump.

[0035] Hemodynamic assistance can be applied to the heart at a first pumping rate during the first of several heartbeats. For example, the first pumping rate may be a first rotor rate, such as the P level mentioned above. Hemodynamic assistance to the heart is adjusted in the second heartbeat (e.g., during systole or after the overlapping notch) by providing the heart with a second pumping rate during the second of several heartbeats. The first pumping rate is thought to be set differently from the second pumping rate.

[0036] The second aortic pressure wave of the heart can be detected during the second beating. Changes in the second aortic pressure wave can be detected by comparing it with a portion of the first aortic pressure wave corresponding to the second beating. In some examples, the second aortic pressure wave can be compared with the first aortic pressure wave by comparing the area under the curve (AUC) for a time period represented by a portion of the second aortic pressure wave with the same time period represented by a portion of the first aortic pressure wave. In some examples, the overall maximum and minimum values ​​of the first and second aortic pressure waves can be compared. The shape or slope of the first and second aortic pressure waves can be compared, i.e., through the change in the derivative of the wave over time. In some implementations, the first aortic pressure wave is compared with the second aortic pressure wave via a nonlinear model (e.g., the Windkessel model described below). Different waveforms provide the model with two sets of values, resulting in two different equations, one for each aortic pressure waveform. The changes between the first and second aortic pressure waves can be used to identify the resistance and compliance of the systemic vascular system. Furthermore, comparing hemodynamic parameters between the first and second aortic pressure waves may include evaluating the linearity of changes in the aortic pressure wave between the first and second pump speeds. For example, aortic pressure may not change linearly between pump speeds; that is, the change in aortic pressure from one pump speed to the next may not progress linearly. Depending on how the change in aortic pressure progresses between pump speeds, it may be possible to predict the aortic pressure at different pump speeds.

[0037] In some implementations, CO is determined based on a nonlinear transfer function that relates CO to whole-body resistance and compliance. In some implementations, the nonlinear transfer function includes the Windkessel model. In some implementations, the transfer function is further related to the aortic pressure waveform.

[0038] The system and method can calculate the change in hemodynamic parameters between a first and second heartbeat by comparing the hemodynamic parameters in the first heartbeat with those in the second heartbeat. This change is at least partially caused by the difference between the first and second output levels of the mechanical circulatory support device. For example, if the hemodynamic parameter is aortic pressure, increasing the output level of the device will result in a higher measured aortic pressure, and decreasing the output level will result in a lower measured aortic pressure. This change in aortic pressure from the first pump output level to the second pump output level demonstrates the contribution of the mechanical circulatory support device to the change in aortic pressure.

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

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

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

[0042] Blood flow in the aorta is due to the pumping action (i p) the original contribution of the heart (i h Since this is equivalent to adding (i), we compare the first set of sine curves and the second set of sine curves to determine the patient's cardiac contribution to blood flow in the aorta (i h ) can be determined. For example, if the hemodynamic parameter is aortic pressure, it can be expressed as the sum of sinusoidal curves obtained from the Fourier transform.

number

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

number

[0044] The heart rate, representing a patient's cardiac function, can be simulated based on a comparison of sinusoidal curves representing the gradual changes in hemodynamic parameters resulting from changes in the pump's operating parameters. For example, a blood pump can be operated at a range of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), each corresponding to the rotational speed of the rotor within the pump and similar frequencies (e.g., 100Hz, 200Hz, 300Hz, 400Hz, 500Hz, 1000Hz, 2000Hz, 3000Hz, etc.). Changing the pump speed (or frequency) alters the blood flow in the vascular system brought about by the pump's operation, thus changing the values ​​of hemodynamic parameters. By gradually passing through multiple pump speeds (or operating parameters such as blood flow brought about by the pump) to identify changes corresponding to one or more hemodynamic parameters, forming hemodynamic waveforms, and decomposing the hemodynamic waveforms from each pump speed, the relationship between pressure and flow rate in diastole is established. Next, the patient's overall cardiac function can be mapped as a mathematical representation (as a function of measured hemodynamic parameters) that can be used to simulate future cardiac function and inform the patient about the handover and control of mechanical circulatory assistance. For example, the measured aortic pressure waveform of a recorded heartbeat can be constructed using the method described below, and the CO of that heartbeat can be calculated.

[0045] As described above, in some implementations, a short-term change in pump speed can be applied to the pump within a single heartbeat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heartbeat can be compared to the aortic pressure of a heartbeat without this short-term speed change or impulse stimulus. The difference between the two (aortic pressure of the modified heartbeat and aortic pressure of the "normal" heartbeat) is the impulse response of aortic pressure: Δp(t)=p1(t)-p2(t) This can be considered as follows, where P1(t) is the pressure waveform measured during impulse stimulation, P2(t) is the pressure waveform without impulse stimulation, and ΔP(t) is the impulse response of the aortic pressure. If this impulse stimulation is applied only during diastole, the difference in total cardiac blood flow between the two heartbeats is... Δi(t)=i1(t)-i2(t) This can be expressed as follows, where i1(t) and i2(t) are the pump flow rates for impulsed and unimpulsed heartbeats, respectively, and Δi(t) is the impulse response of cardiac blood flow. Next, the relationship between aortic pressure and pump flow rate in the frequency domain,

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[0046] [Figure 1] This shows an exemplary cardiac pump system inserted into a patient's blood vessel. [Figure 2] This document describes a process for calculating an index that represents the cardiac performance of a heart in a specific implementation configuration. [Figure 3] This shows plot 300 of the pressure of a cardiac pump system against time in a specific implementation configuration. [Figure 4] The graph shows plots of pressure, motor speed, and flow rate against time for specific implementation configurations. [Figure 5] This shows the Windkessel model in a specific implementation. [Figure 6] This shows a CO sensor combined with a patient in a specific implementation configuration. [Figure 7] This outlines the process for determining the CO (Concept of Occasion) based on a specific implementation. [Figure 8] This describes a process for determining the changes in hemodynamic parameters between two heartbeats under specific implementation configurations. [Figure 9] This outlines the process for determining the CO (Concept of Occasion) based on a specific implementation. [Modes for carrying out the invention]

[0047] Detailed explanation To provide an overall understanding of the systems, methods, and devices described herein, specific exemplary embodiments are described. While the embodiments and features described herein are specifically described in relation to percutaneous cardiac pump systems, it will be understood that the components and other features outlined below can be combined with each other in any suitable way, and that they can be adapted and applied to other types of cardiac treatments and cardiac pump systems, such as cardiac pump systems implanted using surgical incisions and intra-aortic pumps.

[0048] The systems, devices, and methods described herein enable the assessment of organ function by assistive devices located entirely or partially within the organ. In particular, the systems, devices, and methods enable the assessment of cardiac function using cardiac pump systems, such as percutaneous ventricular assist devices. For example, such devices can be used in the treatment of cardiogenic shock.

[0049] By using a cardiac pump system to assess cardiac function, changes in cardiac function can be alerted to healthcare professionals, allowing them to adjust the degree / level of assistance provided by the assistive device (i.e., the blood flow rate pumped by the device) based on the specific patient's needs. For example, the degree of assistance can be increased when the patient's cardiac function deteriorates, or decreased when the patient's cardiac function recovers and returns to a baseline of normal cardiac function. This allows the device to dynamically respond to changes in cardiac function to promote cardiac recovery, enabling the patient to be gradually weaned off treatment. Furthermore, assessment of cardiac function can indicate when it is appropriate to discontinue the use of the cardiac pump system. While some embodiments presented herein concern cardiac pump systems implanted across the aortic valve and partially located in the left ventricle, the concept is applicable to devices in the heart, cardiovascular system, or other areas of the body.

[0050] Assessment of cardiac function may include determining cardiac parameters by utilizing the interaction between the heart and a device. To improve upon traditional linear approximations and provide dynamic variability of vascular responses, using the Windkessel model of the vascular system, the systems and methods described herein introduce controlled turbulence in the vascular system by a cardiac pump system and, in response, calculate cardiac parameters such as stroke volume, vascular resistance and compliance, CO, and left ventricular end-diastolic pressure. In particular, the systems, devices, and methods described herein “ping” the heartbeat using a mechanical circulatory support system. “Ping” includes, for example, increasing the pumping speed of the cardiac pump system over a period of time in a single heartbeat to produce a spike in aortic pressure and flow. Hemodynamic parameters change during ping, and it is possible to detect these changes and compare them to hemodynamic parameters at another point in time (i.e., when the cardiac pump system is not pinged) to calculate other hemodynamic parameters or measure cardiac performance.

[0051] By continuously measuring vascular and cardiac performance using the action of a cardiac pump system, additional clinical data can be provided to aid in the titration of appropriate device assistance. The system and method also provide the use of device-arterial coupling for determining the state of the heart and vascular system, including the determination of intrinsic cardiac output. The mechanical circulatory assistance systems presented herein are located within the heart and operate in parallel with intrinsic ventricular function. This allows the system to have sufficient sensitivity to detect intrinsic ventricular function, unlike some more invasive devices. Therefore, the system, device, and method enable the mechanical circulatory assistance system to be used not only as an assistive device but also as a diagnostic and predictive tool. The cardiac pump system can function as a sensor that extracts information about cardiac function by hydrodynamically coupling with the heart. In some implementations, the power supplied to the assistive device is measured while the cardiac pump system operates at a constant level (e.g., a constant rotor rotation speed). In certain implementations, the rotor speed of the cardiac pump system can be changed (e.g., as a delta, step, or ramp function) to further investigate intrinsic cardiac function.

[0052] Figure 1 shows an exemplary cardiac pump system inserted into a patient's blood vessel. As an example, a cardiac pump system compatible with this disclosure is disclosed in U.S. Patent Application Publication 2018-0078159-A1, the entirety of which is incorporated herein by reference. In general, any other cardiac pump system or other mechanical circulatory support system (as well as sensors for obtaining physiological data from the patient) can be used in conjunction with this disclosure. In some implementations, the systems and methods described herein can be used with deployable pumps (e.g., devices from the Heartmate PHP® family (Thoratec Corporation)) or bypass pumps from the left atrium to the femoral artery (e.g., devices from the TandemHeart family (LivaNova, PLC)). In some implementations, the systems and methods described herein can be used with devices from the IMPELLA® family (Abiomed, Inc., Danvers, Massachusetts).

[0053] The cardiac pump system 100 may operate within the heart, partially within the heart, outside the heart, partially outside the heart, partially outside the vascular system, or at any other suitable location in the patient's vascular system. The cardiac pump system can be considered "in the right place" when the cannula 173 is positioned across the aortic valve such that the blood inlet to the pump (e.g., blood inlet 172) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170) is located in the aorta. The cardiac pump system 100 includes a cardiac pump 106 and a control system 104. All or part of the control system 104 may be located in a controller unit separate from the cardiac pump 106 / remote. In some implementations, the control system 104 is located inside the cardiac pump 106. The control system 104 and cardiac 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 on which the pump elements are formed. The pump 100 includes a pump housing 134 and a motor housing 102, the motor housing 102 which rotates an impeller blade on a drive shaft, coupled to the cannula 173 at the distal end 111 of the motor housing 102, within the pump housing 134 to generate blood flow to the cannula 173 at the suction head 174. The suction head 174 brings a blood inlet 172 to the distal end 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.

[0054] The rotation of the drive shaft within the pump housing 134 rotates the pump elements within the bearing gap. Blood-compatible fluid is delivered through the elongated catheter 105 and motor housing 102 to the proximal end of the cannula 173, where the fluid lubricates the pump. The flow of blood-compatible fluid has a second direction 122 through the pump's bearing gap. After leaving the bearing gap, the blood-compatible fluid, following the flow direction 123, mixes with the blood flow and flows into the aorta with the blood.

[0055] The cardiac pump 100 is inserted into the patient's blood vessels through a sheath 175. The pump housing 134 encloses the rotor and internal bearings and may be sized to be inserted percutaneously into the patient's blood vessels. In some implementations, the pump is advanced through the vascular system and past the aortic arch 164. Although the pump is illustrated in the left ventricle, the pump may instead be located in the right heart to pump blood from the patient's inferior vena cava or right atrium through the right ventricle to the pulmonary artery.

[0056] To stabilize the cardiac pump 100 in a blood vessel or ventricle, a flexible projection 176 is included at the distal end portion 171 of the cannula 173, distal to the suction head 174. The flexible projection 176 is non-traumatic and helps prevent the suction head 174 from approaching the blood vessel wall and sticking to the wall due to suction. Since the flexible projection 176 does not have suction, it mechanically extends the pump 100 but not with respect to fluid. In some implementations, the flexible projection can be formed as a pigtail. In some aspects, the pump does not necessarily have to include a flexible projection.

[0057] The elongated catheter 105 houses a connection 126 equipped with a fluid supply line and an electrical connection cable. Furthermore, the connection 126 supplies blood-compatible fluid from a fluid reservoir to a pump and is housed within a control system 104.

[0058] The control system 104 includes a controller 182 that controls the pump 106 by supplying power to the motor and controlling the motor speed. The control system 104 includes circuits for monitoring motor current for a drop in current indicating air in the line, monitoring changes in the differential pressure signal, monitoring the flow position, monitoring suction, or any other appropriate measurement. In some implementations, the control system 104 includes a display screen for displaying measured values ​​such as the differential pressure signal and motor current. The control system 104 may include warning sounds, warning lights, or indicators to alert the operator to sensor failure, disconnection or damage to connection 126, or a sudden change in the patient's health.

[0059] Motor 108 is configured to operate at the speed necessary to maintain the rotor at a set speed. As a result, the motor current drawn by the motor to maintain the rotor speed can be monitored and used to understand the underlying cardiac state, as will be further described below. Control system 104 is configured to change the pump speed within the heartbeat cycle of the assisted heart to bring about a change in blood flow through the pump, and the change in pump speed is synchronized with the heartbeat by at least one event per heartbeat cycle relating to a predetermined event within the heartbeat cycle, i.e., the systems, devices, and methods described herein “ping” the heartbeat using the cardiac pump system. When “ping” occurs, the pump speed of the cardiac pump system (or other mechanical circulatory assistance device) is increased or decreased for a relatively short period of time, for example, during one phase of a heartbeat cycle, and then changed to a baseline or other speed. The pump speed can be increased for a period in a single heartbeat or for a period spanning multiple heartbeats.

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

[0061] In some implementations, the pump velocity is increased during systole, diastolic, or both of the same heartbeat. The ping is timed so that the increase in pump velocity occurs over a known period of the heartbeat. For example, the start of velocity ping can be synchronized to the start of diastole, the end of diastole, the start of systole, the end of systolic, peak systolic pressure, or any other appropriate time. The end of velocity ping can be synchronized to the start of diastole, the end of diastole, the start of systole, the end of systolic, peak systolic pressure, or any other appropriate time. In some implementations, the pump velocity is increased or decreased over a set period of time. For example, the start of ping can be synchronized to the start of diastole, the overlapping notch or after the overlapping notch, the end of diastole, the start of systole, the end of systolic, peak systolic pressure, or any other appropriate time. The ping may continue for a set period of time. For example, the ping may continue for 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other appropriate time.

[0062] The control system 104 includes a current sensor (not shown). The controller 182 supplies current to the motor 108 via a connection 126 through one or more wires. The current supplied to the motor 108 via the connection 126 is measured by the current sensor. The load on the motor of the mechanical pump corresponds to the force of the pressure head or to the difference between aortic pressure and left ventricular pressure. The cardiac pump 106 is subjected to a nominal load in steady-state operation at a given pressure head, and variations from this nominal load are the result of changes in 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 operate at the 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 the underlying cardiac state. The cardiac state can be accurately quantified and understood by simultaneously monitoring the pressure head during the heartbeat cycle using the pressure sensor 112. The cardiac parameter estimator 185 receives current signals from the current sensor and pressure signals from the pressure sensor 112. The cardiac parameter estimator 185 uses these current and pressure signals to reveal cardiac function. The cardiac parameter estimator 185 can access a stored lookup table to obtain additional information to reveal cardiac function based on the pressure and current signals. For example, the cardiac parameter estimator 185 can receive aortic pressure from the pressure sensor 112 and use the lookup table to determine delta pressure using the aortic pressure. The cardiac parameter estimator 185 may be software programmed into the controller 182, or it may be separate hardware connected to the controller 182 by a wired or wireless connection. The cardiac parameter estimator 185 is configured to perform the algorithms described herein. For example, the cardiac parameter estimator 185 may be configured to estimate pump flow rate based on the current supplied to the pump, and may be configured to reveal intrinsic cardiac output according to the method described herein.

[0063] Various implementations of pressure sensors 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 external and internal pressures of the cannula, which may be displayed by the cardiac pump system. When the cardiac pump system is positioned correctly across the aortic valve, the upper (external) side of the sensor is exposed to the aortic pressure, and the lower (internal) side of the sensor is exposed to the intraventricular pressure. Therefore, the differential pressure signal is approximately equal to the difference between the aortic pressure and the intraventricular pressure. Other sensors, such as optical sensors or fluid-filled columns, can also be used.

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

[0065] In step 202, the pump (e.g., pump 102 in Figure 1) is placed in the patient's heart. In some implementations, the pump is an intravascular blood pump device placed in the patient's heart by percutaneous insertion. In some implementations, the pump may be a surgically implanted device, a left ventricular assist device, a counterpulsation device, a deployable cardiac pump, or any other suitable device. The pump may be introduced to the patient because the patient is in a state of cardiogenic shock or otherwise facing a decline in health. The pump may be positioned across the aortic valve such that the blood inlet to the pump (e.g., blood inlet 172 in Figure 1) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 in Figure 1) is located in the aorta. The pump is CO=i h +i p (1) This contributes to the normal function of the heart, where CO is the total cardiac output, i h This is the actual cardiac output, i p This represents the flow rate due to the pump's contribution.

[0066] In step 204, hemodynamic parameters are monitored while the pump operates at a first pump speed. The hemodynamic parameters may be any parameters related to blood flow in the body. For example, the hemodynamic parameters may include heart rate, blood pressure, arterial oxygen saturation, mixed venous saturation, central venous oxygen saturation, arterial blood pressure, mean arterial pressure, right arterial pressure, central venous pressure, right ventricular pressure, pulmonary artery pressure, mean pulmonary artery pressure, pulmonary occlusion pressure, left atrial pressure, aortic pressure, differential pressure, left ventricular end-stage 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 resistance index, and left ventricular stroke work. This may include at least one of the following: 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, arterial-venous oxygen content difference, oxygen transport, oxygen transport index, oxygen consumption, oxygen consumption index, oxygen uptake rate, oxygen uptake index, total peripheral resistance, CO, cardiac index, and CPO. Pump speed is the speed of the pump's operation and corresponds to the amount of blood flow brought about by the pump's operation. In some implementations, pump speed may correspond to the rotor's rotational speed. For example, the pump speed may be 10,000 RPM, 20,000 RPM, 30,000 RPM, 40,000 RPM, 50,000 RPM, 60,000 RPM, 70,000 RPM, 80,000 RPM, 90,000 RPM, 100,000 RPM, or any other appropriate speed. The pump speed may correspond to a power level or P level, as described above in relation to Figure 1. For example, the pump speed may be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, P-9, or any other appropriate value. In some implementations, the pump speed may instead correspond to the rate at which the pump chamber is filled with blood and the rate at which blood is released.

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

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

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

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

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

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

[0073] In step 216, an index representing the cardiac performance of the heart is calculated based on the change in hemodynamic parameters between the first and second phases. For example, an index representing cardiac performance may be systemic resistance, cardiac compliance, CO, CPO, stroke volume, stroke work, ejection fraction, myocardial contractility, ventricular elastance, cardiac index, and prediction of patient survival. For example, the numerical difference between the first blood volume pumped by the heart in the first phase of the heartbeat cycle and the second blood volume pumped by the heart in the second phase of the heartbeat cycle can be calculated. Using the numerical difference in blood volume, the stroke volume of an individual heartbeat or the average cardiac output (CO) over a desired period can be determined. Many of the indicators representing cardiac performance are interrelated. For example, CO is determined based on the flow rate of blood passing through the pump. Stroke volume is an indicator 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 performed by the ventricles to pump blood and can be calculated from stroke volume according to the formula SW = SV * MAP, where SW is stroke work, SV is stroke volume, and MAP is mean arterial pressure. Cardiac work is calculated by the product of stroke work and heart rate. CPO is a measure of cardiac function representing the pumping capacity of the heart, and its unit is watts. CPO is calculated using the formula CPO = mAoP * CO / 451, where CPO is cardiac power, mAoP is mean arterial pressure, CO is cardiac output, and 451 is a constant used to convert mmHg x L / min to watts. Ejection fraction can be calculated by dividing stroke volume by the ventricular blood volume. Other parameters such as chamber pressure, preload state, afterload state, cardiac recovery, flow load state, variable volume load state, and / or heart rate cycle flow state can be calculated from these values ​​or determined through these parameters. In some implementations, indices representing cardiac performance are calculated via a two-element Windkessel model of the vascular system (e.g., the Windkessel model in Figure 5) to model the dynamic and nonlinear interactions between the heart and blood vessels.Therefore, this process employs a time-varying nonlinear model of the vascular system and leverages a well-controlled analogue of the ventricular-vascular interface with a device-arterial interface to continuously determine systemic vascular resistance and compliance and quantify cardiac stroke volume without requiring additional external measurements.

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

[0075] Figure 3 shows a pressure-versus-time plot 300 of a cardiac pump system in a specific implementation configuration. The y-axis of plot 300 represents aortic pressure (mmHg), while the x-axis represents time as a percentage of heartbeat length. In particular, plot 300 shows the effect of ping on aortic pressure. t1 represents the time of the first heartbeat, and t2 represents the time of the second heartbeat following the first. Times t1 and t2 occur when the cardiac pump system is at least partially located within the patient's heart. Point 310 represents the systolic peak pressure in the first heartbeat, and point 320 represents the systolic peak pressure in the second heartbeat. Point 312 represents the overlapping notch in the first heartbeat, and point 322 represents the overlapping notch in the second heartbeat. Diastolic times t3 and t4 represent the diastolic phases of the first and second heartbeats, respectively. At time t1, the pump operates at the first pump speed. At time t4, the pump operates at a second pump speed that is faster than the first pump speed.

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

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

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

[0079] Figure 5 shows the Windkessel model 500. The Windkessel model 500 includes a current source 510, a current source 520, a resistor 530, and a compliance 540. The governing equations for this model are:

number

number

number

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

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

[0082] In some implementations, the CO sensor 610 includes an elongated catheter body coupled to a cannula. The elongated catheter may include a drive cable, electrical wiring connecting the blood pump to a control system, any appropriate elements, or any combination thereof. In some implementations, the blood pump includes a pump housing and a motor housing, the motor housing being coupled to the cannula at its distal end. A rotor can be rotated within the pump housing to generate blood flow to the cannula.

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

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

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

[0086] Figure 7 shows the process 700 for determining CO. Process 700 can be performed using the cardiac pump system 100 in Figure 1 or any other suitable pump. In some implementations, the pump is an intravascular blood pump device placed inside the patient's heart by percutaneous insertion. The pump may be introduced to a patient because the patient is in a state of cardiogenic shock or facing other forms of health deterioration. The pump can be positioned across the aortic valve such that the blood inlet to the pump (e.g., blood inlet 172 in Figure 1) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 in Figure 1) is located inside the aorta. CO=i h +i p (1) In this way, it contributes to the normal function of the heart, where CO is the total cardiac output, i h This is the actual cardiac output, i p This represents the flow rate due to the pump's contribution.

[0087] In step 702, a first aortic pressure wave is detected. The first aortic pressure wave reflects multiple heartbeats, each reflected beat containing overlapping notches. The pressure waveform can be measured by a pressure sensor. In some implementations, the pressure sensor may be built into the pump. In some implementations, the pressure sensor may be located outside the pump. The pressure sensor can communicate with a controller configured to control the operation of the pump.

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

[0089] In step 708, the second aortic pressure wave of the heart is detected during the second beating. In step 710, the second aortic pressure wave is compared to the portion of the first aortic pressure wave corresponding to the second beating in order to detect changes in the second aortic pressure wave. The changes between the first and second aortic pressure waves can be used to identify the resistance and compliance of the systemic vascular system.

[0090] In 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 implementations, the nonlinear transfer function includes the Windkessel model described above in relation to Figure 5.

[0091] Figure 8 shows a process 800 for determining changes in hemodynamic parameters between two heartbeats. In step 802, a mechanical circulatory support device is placed within the patient's vascular system. In some implementations, the device is an intravascular blood pump device placed within the patient's heart by percutaneous insertion. The device may be introduced into a patient because the patient is in a state of cardiogenic shock or otherwise facing impaired vascular health. The device may be a left ventricular device or a right ventricular device. In some implementations, the device is positioned across the aortic valve such that the blood inlet to the device (e.g., blood inlet 172 in Figure 1) is located in the left ventricle and the outlet from the device (e.g., outlet opening 170 in Figure 1) is located within the aorta.

[0092] The device is capable of operating to alter hemodynamic parameters in the patient. For example, the device's operation can affect the patient's aortic pressure by pumping blood from the left ventricle into the aorta. The device operates at a first power level while the heart is beating. The first power level corresponds to a first blood flow velocity resulting from the contribution of the mechanical circulatory assistance device to the patient's intrinsic blood flow during device operation at the first power level. For example, the first power level may be associated with a first motor speed, such as the P level described above.

[0093] The device operates at a first power level over a period including the duration of a first heartbeat, and the patient's hemodynamic parameters are monitored during the device's operation. The results of this monitoring are determined as a function of time within each heartbeat and stored in the device's memory (or another data storage device). As described above, the hemodynamic parameters may be any parameters relating to blood flow in organs or tissues of the body. In step 804, the hemodynamic parameters are detected during the first heartbeat and are temporally coincidental to the first heartbeat. The device power level and hemodynamic parameter measurements during that heartbeat (or any other time at the first power level) coincide with events in the heartbeat cycle (systole, diastole, overlapping notch, etc.). As a result, the hemodynamic parameters and device power level can be associated with events in the heartbeat cycle at various points in time. For example, it can be easily detected that a pump operating at a first power level has a first measured hemodynamic parameter (e.g., aortic pressure) at or after the overlapping notch of the first heartbeat. In some implementations, 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 contained on a cannula that extends partially into the patient's left ventricle.

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

[0095] In step 808, hemodynamic parameters are detected during the second heartbeat (during the period of its second power level) at or near the same time as the first heartbeat in the cardiac phase of the second heartbeat. Hemodynamic parameters can be measured over the entire first or second heartbeat, or for a portion of each beat. For example, hemodynamic parameters can be measured during the systolic or diastolic phase of the second heartbeat, or at the overlapping notch.

[0096] In step 810, hemodynamic parameters measured in the first heartbeat are compared with hemodynamic parameters measured in the second heartbeat. These two measurements are taken at approximately the same point in the cardiac cycle, although in two different beats. The difference in hemodynamic measurements arises from the change in pump speed between the first and second heartbeats. For example, if the hemodynamic parameter is aortic pressure, increasing the power level will result in a higher measured aortic pressure, and decreasing the power level will result in a lower measured aortic pressure. This change in aortic pressure from the first power level to the second power level correlates with the contribution of the mechanical circulatory support device to the change in total cardiac output.

[0097] Figure 9 shows a process 900 for determining cardiac output by a heartbeat "ping" process. Process 900 can be performed using the cardiac pump system 100 of Figure 1 or any other suitable pump. The pump is placed in the patient's heart by percutaneous insertion. The patient may be in a state of cardiogenic shock or facing other forms of impaired vascular health. The pump may be a left ventricular device or a right ventricular device. The pump is positioned across the aortic valve such that the blood inlet to the pump (e.g., blood inlet 172 in Figure 1) is located in the left ventricle and the outlet from the pump (e.g., outlet opening 170 in Figure 1) is located in the aorta. CO=i h +i p (1) This contributes to the normal function of the heart, where CO is the total cardiac output, i h This is the actual cardiac output, i p This represents the flow rate due to the pump's contribution.

[0098] In step 902, the pump operates at a first pump speed for a first period, which includes the period of the first heartbeat. In step 904, hemodynamic parameters are monitored during the operation of the cardiac pump at the first pump speed during the first diastolic phase of the first heartbeat. Hemodynamic parameters relate to the flow of blood in the body. Pump speed is the speed of the pump's operation and corresponds to the amount of blood flow produced by the pump's operation. In some implementations, pump speed corresponds to the rotational speed of the pump's rotor. For example, pump speed can be 10,000 RPM, 20,000 RPM, 30,000 RPM, 40,000 RPM, 50,000 RPM, 60,000 RPM, 70,000 RPM, 80,000 RPM, 90,000 RPM, 100,000 RPM, or any appropriate speed or higher. Pump speed can correspond to power level or P level, as described above in relation to Figure 1. For example, the pump speed may be P-1, P-2, P-3, P-4, P-5, P-6, P-7, P-8, P-9, or any other appropriate value. In some implementations, the pump speed corresponds to the rate at which the pump chamber is filled with blood and released. By monitoring hemodynamic parameters, the systems and methods described herein can identify changes in those hemodynamic parameters over time, including the phases of the first and second heartbeats. Using such comparisons, cardiac performance can be quantified (e.g., via CO), as will be discussed more fully herein.

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

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

[0101] In step 908, the first pump speed is changed to a second pump speed so that the cardiac pump action produces a second power level during the second diastolic phase of the second heartbeat. The second pump speed may be greater than or less than the first pump speed. In some implementations, the timing of the increase in pump speed is determined so that the increase in pump speed occurs during a predicted period of the heartbeat. For example, the start of the speed increase can be synchronized with the start of diastole to account for the time delay between the transmission of the speed change instruction to the pump and the physical occurrence of the speed change. The end of the speed increase can be synchronized with the start of diastole, the end of diastole, the start of systole, the end of systole, peak systolic pressure, or any other appropriate time. In some implementations, the system is configured so that the pump speed is increased or decreased over a set period of time. For example, the speed change may last for approximately 0.05 seconds, 0.1 seconds, 0.2 seconds, 0.3 seconds, or any other appropriate time. The second heartbeat is distinct from the first heartbeat. Hemodynamic parameters are measured during the second heartbeat, and this can be done at the same time point (e.g., overlapping notch) as when the first hemodynamic parameters were measured during the first heartbeat. In some implementations, detection and measurement are applied to the second heartbeat that occurs after the first heartbeat.

[0102] In step 910, hemodynamic parameters are monitored during the second diastolic phase of the second heartbeat, for example, at the diastolic overlap notch. In step 912, second operating parameters of the intravascular blood pump during the second diastolic phase are determined. Determining the second operating parameters may include determining the second blood flow velocity (or second level of motor operating parameters) provided by the blood pump during the second diastolic phase.

[0103] In step 914, an index representing the cardiac performance of the heart is calculated. This index is based on (i) a first operating parameter, (ii) a second operating parameter, and (iii) hemodynamic parameters in the first and second diastolic periods (e.g., in overlapping notches during both periods). The index can be used in a set of transfer functions or equations as described above in relation to the Windkessel model. In some implementations, the mathematical representation of the hemodynamic parameters is determined for the first and second diastolic periods. For example, the mathematical representation may be the sum of sine curves.

[0104] The indicators are used to construct waveforms that can be used to determine cardiac output. Calculating cardiac performance may involve decomposing a first waveform representing hemodynamic parameters for the first diastolic phase to determine a first set of sinusoidal curves, and decomposing a second waveform representing hemodynamic parameters for the second diastolic phase to determine a second set of sinusoidal curves. These decompositions may involve applying a Fourier transform to the first waveform, the second waveform, or both. The set of sinusoidal curves may include one or more summed sinusoidal curves.

[0105] When the pump is operating in the patient's vascular system, blood flow in the aorta is due to the pump's contribution (i p ) the original contribution of the heart (i h This is equivalent to adding (i) to the first set of sine curves. By comparing the first set of sine curves with the second set of sine curves, the patient's heart (i) is used to determine the blood flow in the aorta. h The contribution of ) can be determined. For example, aortic pressure may be a hemodynamic parameter, and aortic pressure can be expressed as the sum of sinusoidal curves obtained from the Fourier transform,

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

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[0107] In some implementations, a model heart rate representing a patient's cardiac function can be simulated based on a comparison of sinusoidal curves and used to determine cardiac output, timing, and level of mechanical circulatory assistance. For example, a blood pump can be operated at a range of pump speeds (e.g., P-1, P-2, P-3, P-4, etc.), each corresponding to the rotational speed of the rotor within the pump and similar frequencies (e.g., 100Hz, 200Hz, 300Hz, 400Hz, 500Hz, 1000Hz, 2000Hz, 3000Hz, etc.). Changing the pump speed (or frequency) alters the blood flow in the vascular system brought about by the pump's operation, thus changing the values ​​of hemodynamic parameters. By gradually passing through multiple pump speeds (or operating parameters such as blood flow brought about by the pump) to identify changes corresponding to one or more hemodynamic parameters, forming hemodynamic waveforms, and decomposing the hemodynamic waveforms from each pump speed, the relationship between pressure and flow rate in diastole is established. Next, the patient's overall cardiac function can be mapped as a mathematical representation (as a function of measured hemodynamic parameters) that can be used to simulate future cardiac function and inform the patient about the handover and control of mechanical circulatory support. For example, the measured aortic pressure waveform of a recorded heartbeat can be constructed using the method described below, and the CO of that heartbeat can be calculated.

[0108] As described above, in some implementations, a short-term change in pump speed can be applied to the pump within a single heartbeat. This change in pump speed can be considered an impulse stimulus. The aortic pressure recorded for this heartbeat can be compared to the aortic pressure of a heartbeat without this short-term speed change or impulse stimulus. The difference between the two (aortic pressure of the modified heartbeat and aortic pressure of the "normal" heartbeat) is the impulse response of aortic pressure: Δp(t)=p1(t)-p2(t) It can be considered as such, Here, P1(t) is the pressure waveform measured during impulse stimulation, P2(t) is the pressure waveform without impulse stimulation, and ΔP(t) is the impulse response of aortic pressure. If this impulse stimulation is applied only during diastole, the difference in total cardiac blood flow for the two heartbeats is: Δi(t)=i1(t)-i2(t) It can be expressed as follows: Here, i1(t) and i2(t) are the pump flow rates for impulsed and unimpulsed heartbeats, respectively, and Δi(t) is the impulse response of cardiac blood flow. Next, the relationship between aortic pressure and pump flow rate in the frequency domain

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[0109] Figure 10 shows two plots for the same 10-second period, one for aortic pressure and the other for cardiac blood flow. The y-axis of the upper plot represents aortic pressure (mmHg), while the x-axis represents time (seconds). The y-axis of the lower plot represents calculated total cardiac blood flow (liters / minute), while the x-axis represents time (seconds). In this example, systemic vascular resistance R and compliance C are known. For example, R and C can be calculated using the measured aortic pressure obtained during the 10-second period shown in the figure, combined with the pump data as described above. Total cardiac blood flow i h +i p Equation (2):

number

[0110] The above is merely an example of the principles of this disclosure, and the device can be implemented in aspects other than those described above, which are presented for illustrative purposes only, not limitation. Although the device disclosed herein is shown for use in percutaneous insertion of a cardiac pump, it should be understood that it is applicable to devices for other applications requiring hemostasis.

[0111] Those skilled in the art will be able to conceive of changes and modifications by considering this disclosure. The disclosed features can be implemented in any combination and partial combination (including multiple dependent combinations and partial combinations) with one or more other features described herein. The various features described or illustrated above, including all their components, can be combined or integrated into other systems. Furthermore, certain features may be omitted or not implemented at all.

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

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

[0114] Computer programs (also known as programs, software, software applications, scripts, or code) can be written in any form of programming language, including compiled or interpreted languages, and can be deployed in any form, such as a standalone program or as modules, components, subroutines, or other units suitable for use in a computing environment. Computer programs can correspond to files in a file system. A program can be stored as part of a file that holds other programs or data (e.g., one or more scripts stored in a markup language document), as a single file dedicated to the program in question, or as part of multiple collaborative files (e.g., a file containing one or more modules, subprograms, or parts of code). Computer programs can be deployed to run on one computer, or on multiple computers located in one place or distributed across multiple locations and interconnected by a communication network.

[0115] The processes and logic flows described herein can be executed by one or more programmable processors that run one or more computer programs to perform their functions by manipulating input data and generating outputs. Alternatively, the processes and logic flows can be executed by dedicated logic circuits, such as FPGAs (Field-Programmable Gate Arrays) or ASICs (Application-Specific Integrated Circuits), and the device can also be implemented as such a dedicated logic circuit.

[0116] Processors suitable for executing computer programs include, for example, both general-purpose and dedicated microprocessors, as well as any one or more processors in any type of digital computer. Generally, a processor receives instructions and data from read-only memory or random-access memory or both. The main components of a computer are a processor for executing instructions, and one or more memory devices for storing instructions and data. Generally, a computer further includes one or more mass storage devices for storing data, such as magnetic disks, magneto-optical disks, or optical disks, or is operably coupled to such mass storage devices for receiving data, transferring data, or both. However, a computer does not necessarily need to have such devices.

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

Claims

1. A method for determining a patient's vascular performance, including the following steps: The process of placing a pump configured to operate at an adjustable pumping speed inside the heart, A step of pumping the aforementioned pump at a first pump speed, The process involves monitoring hemodynamic parameters during pumping at the first pump speed, A step of identifying the first phase of the first heartbeat of the heart over a first period, A step of predicting the second phase of the second heartbeat of the heart over a second period based on the monitored hemodynamic parameters, A step of changing the pump speed to a second pump speed in the second phase of the second heartbeat, The second phase includes the step of monitoring the hemodynamic parameters, A step of comparing the monitored hemodynamic parameters in the first phase with the monitored hemodynamic parameters in the second phase to calculate the change in the hemodynamic parameters between the first phase and the second phase, A step of calculating metrics representing vascular performance based on the changes in the hemodynamic parameters between the first phase and the second phase.

2. A step of changing the second pump speed to the first pump speed after the second phase. The method according to claim 1, further comprising:

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

4. The method according to any one of the claims, wherein the first phase is one of systole and diastole, and the second phase is one of systole and diastole.

5. The method according to claim 4, wherein the first phase is a first expansion phase, and the second phase is a second expansion phase.

6. The method according to claim 4, wherein the first phase is a first systolic period, and the second phase is a second systolic period.

7. The method according to any one of the claims, wherein the step of identifying the first phase of the first heartbeat is based on the change over time of the monitored hemodynamic parameters in the step of pumping at the first pump speed.

8. The step of predicting the second phase of the second heartbeat of the heart is the method according to any one of the claims, based on the identified first phase of the heart cycle.

9. Comparing the hemodynamic parameters in the first phase with those in the second phase is, The first phase involves calculating the first blood volume pumped by the heart, The calculation of the second blood volume pumped by the heart in the second phase, To determine the numerical difference between the first blood volume and the second blood volume. The method according to any one of the claims, including

10. A step of evaluating the linearity of the time-dependent changes in the hemodynamic parameters between the first phase and the second phase. The method according to claim 8, further comprising:

11. Determining the cardiac performance of the heart is Based on the changes in the hemodynamic parameters between the first phase and the second phase, the vascular compliance and vascular resistance of the systemic vascular system are calculated. Calculating cardiac output using the Windkessel model and The method according to any one of the claims, determined by...

12. The method according to any one of the claims, wherein changing the first pump speed to a second pump speed includes increasing the pump speed.

13. The method according to any one of the claims, wherein changing the first pump speed to a second pump speed includes reducing the pump speed.

14. A method for determining a patient's cardiac output using a processor-controlled intravascular device, including the following steps: A step of detecting the first aortic pressure wave of the heart, which reflects multiple heartbeats, wherein each reflected heartbeat includes overlapping notches. The steps include applying hemodynamic assistance to the heart at a first pumping speed during the first of the plurality of beats, The steps include adjusting hemodynamic assistance to the heart in the second of the plurality of beats by introducing a second pumping rate to the heart after the overlapping notch of the beat in the second beat, The steps include detecting the second aortic pressure wave of the heart during the second beating, The processor compares the second aortic pressure wave with the portion of the first aortic pressure wave corresponding to the second pulsation, detects changes in the second aortic pressure wave, and identifies the resistance and compliance of the vascular system. A step in which the intrinsic cardiac output of the heart is determined based on a nonlinear transfer function programmed in the software that correlates cardiac output with vascular resistance and compliance.

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

16. The method according to any one of the claims, wherein hemodynamic assistance is provided by an intracardiac blood pump having a cannula configured to be positioned across the aortic valve.

17. A step of adjusting hemodynamic support based on at least one of the determined cardiac output, resistance, or compliance. The method according to any one of the claims, further comprising:

18. An intravascular blood pump comprising a cannula, a rotor configured to rotate within a blood vessel and pump blood through the cannula, and a drive mechanism configured to supply power to rotate the rotor, A pressure sensor, operably positioned on the proximal or distal side of the blood pump and configured to detect intravascular pressure resulting at least partially from the pumping of blood within the blood vessel, and A controller electrically connected to the pressure sensor and configured to detect a signal from the pressure sensor representing blood pressure, and configured to calculate the intrinsic cardiac output based on a nonlinear model relating cardiac output to vascular resistance and compliance. A sensor equipped with a function to determine the actual cardiac output of a beating heart.

19. The aforementioned controller, The first aortic pressure wave of the heart is detected, reflecting multiple heartbeats (where each reflected heartbeat includes overlapping notches). In the first of the aforementioned multiple pulsations, hemodynamic assistance is applied to the heart at a first pumping speed. In the second of the aforementioned multiple pulsations, a second pumping rate is introduced to the heart after the overlapping notch of the pulsation, thereby adjusting the hemodynamic assistance to the heart in the second pulsation. The second aortic pressure wave of the heart is detected during the second beating. The second aortic pressure wave is compared with a portion of the first aortic pressure wave corresponding to the second pulsation to detect changes in the second aortic pressure wave, and to identify the aortic resistance and compliance, and The software determines the intrinsic cardiac output based on a nonlinear transfer function that correlates cardiac output with aortic resistance and compliance. The sensor according to claim 18, configured as follows.

20. The sensor according to claim 19, wherein the nonlinear transfer function includes a Windkessel model.

21. The sensor according to any one of claims 18 to 20, wherein the controller is further configured to adjust hemodynamic assistance provided based on at least one of the determined cardiac output, the vascular resistance, or the vascular compliance.

22. A method for providing mechanical circulatory support to a patient using a blood pump, including the following steps: The process of operating the blood pump within the patient's vascular system, A step of determining the cardiac output of the patient's heart using either the above or a sensor, A step of adjusting the pumping speed of the blood pump based on the determined cardiac output.

23. A blood pump having a cannula configured to extend within the patient's vascular system, A pressure sensor configured to detect intravascular pressure resulting at least partially from the pumping of blood within the vascular system, The pressure sensor signal and the operating parameters of the blood pump are received. The intrinsic cardiac output is determined according to any of the methods described herein, and The operating parameters of the blood pump are adjusted based on the determined original cardiac output. A controller configured in such a way A mechanical circulation support system equipped with the above.

24. The system according to claim 23, wherein the pressure sensor is operably positioned proximal or distal to the blood pump and is configured to detect aortic pressure.

25. The system according to any one of claims 23 to 24, wherein the blood pump is an intracardiac blood pump, and the cannula is configured to extend into the left ventricle of the heart.

26. The aforementioned blood pump is A rotor configured to rotate within a blood vessel and pump blood through the cannula, A drive mechanism configured to supply power to rotate the rotor, A system according to any one of claims 23 to 25, comprising:

27. The aforementioned controller, A command is sent to the pump to operate at a first pump speed. A first pressure sensor signal representing the pressure during pumping at the first pump speed is received. Identify the first phase of the first heartbeat over a first period, Based on the first pressure sensor signal, the second phase of the second heartbeat of the heart over a second period is predicted. An operation command is sent to the pump to operate at a second pump speed during the second phase of the second heartbeat. A second pressure sensor signal representing the pressure during pumping at the second pump speed is received. The first pressure sensor signal is compared with the second pressure sensor signal to calculate the change in pressure between the first phase and the second phase, and Based on the pressure change between the first phase and the second phase, the intrinsic cardiac output is calculated. The system according to any one of claims 23 to 26, configured as described above.

28. The system according to claim 27, further comprising changing the second pump speed to the first pump speed after the second phase.

29. The system according to any one of claims 27 to 28, wherein the first phase is one of systole and diastole, and the second phase is one of systole and diastole.

30. The system according to any one of claims 27 to 29, wherein changing the first pump speed to a second pump speed includes increasing the pump speed.

31. The system according to any one of claims 27 to 30, wherein changing the first pump speed to a second pump speed includes reducing the pump speed.

32. The system according to any one of claims 27 to 31, wherein determining the intrinsic cardiac output of the heart is based on a nonlinear transfer function programmed in the software that correlates cardiac output with aortic resistance and compliance.

33. The system according to claim 32, wherein the nonlinear transfer function includes a Windkessel model.

34. The system according to any one of claims 23 to 33, wherein adjusting the operating parameters of the blood pump based on the determined original cardiac output includes transmitting a command to the blood pump to increase or decrease the pump speed.

35. The system according to any one of claims 23 to 34, further comprising a display, wherein the controller is configured to display at least one of the pressure sensor signals, the operating parameters of the blood pump, or the intrinsic cardiac output.

36. The aforementioned controller, The current supplied to the pump is detected, and Based on the current supplied to the pump, the blood flow rate to which the pump's operation contributes is determined. The system according to any one of claims 23 to 35, configured as described above.

37. A method for determining the heart's inherent cardiac performance, including the following steps: A step of placing a mechanical circulatory support device capable of operating to alter hemodynamic parameters within the patient into the patient's vascular system, and operating the device at a first output level during a first heartbeat, A step of detecting hemodynamic parameters in the first heartbeat, A step of operating the device to output a second output level in the second heartbeat, A step of detecting the hemodynamic parameters in the second heartbeat, A step of comparing the hemodynamic parameters in the first heartbeat with the hemodynamic parameters in the second heartbeat, and calculating the change in the hemodynamic parameters between the first heartbeat and the second heartbeat.

38. A step of calculating an index representing the intrinsic cardiac performance of the heart based on the changes in the hemodynamic parameters between the first heartbeat and the second heartbeat. The method according to claim 37, further comprising:

39. The method according to claim 37 or 38, wherein the mechanical circulatory support device comprises an intracardiac blood pump having a cannula configured to extend into the left ventricle of the heart.

40. The method according to any one of claims 37 to 39, wherein the hemodynamic parameter is aortic pressure, and the mechanical circulatory support device comprises a pressure sensor configured to detect aortic pressure.

41. A step of calculating the vascular compliance and vascular resistance of the heart based on the changes in the hemodynamic parameters between the first heartbeat and the second heartbeat, The process of calculating the heart's natural cardiac output and The method according to any one of claims 37 to 40, further comprising:

42. A step of comparing the monitored hemodynamic parameters during the first diastolic phase of the first heartbeat with the monitored hemodynamic parameters during the second diastolic phase of the second heartbeat to calculate the change in the hemodynamic parameters between the first and second diastolic phases, A step of determining aortic resistance and compliance based on the changes in the hemodynamic parameters between the first diastolic phase and the second diastolic phase, A process for determining cardiac output based on a nonlinear transfer function that relates cardiac output to aortic resistance and compliance, and The method according to any one of claims 37 to 41, further comprising:

43. A method for determining the patient's intrinsic cardiac output, including the following steps: A step of operating the intravascular blood pump at a first pump speed, The process involves monitoring hemodynamic parameters during the operation of the intravascular blood pump at the first pump speed during the first diastolic phase of the first heartbeat, A step of determining the first operating parameters of the intravascular blood pump during the first diastolic phase, The process involves changing the first pump speed to a second pump speed so that the operation of the intravascular blood pump changes during the second diastolic phase of the second heartbeat. A step of monitoring the hemodynamic parameters during the second diastolic phase of the second heartbeat, The process involves determining a second operating parameter of the intravascular blood pump during the second diastolic phase, (i) a step of calculating an index representing the cardiac performance of the heart based on the first operating parameter, (ii) the second operating parameter, and (iii) the hemodynamic parameters during the first and second diastolic phases.

44. The method according to claim 43, wherein the step of determining the first operating parameter includes determining a first blood flow rate provided by the blood pump during the first diastolic phase, and the step of determining the second operating parameter includes determining a second blood flow rate provided by the blood pump during the second diastolic phase.

45. The process involves determining the mathematical representation of the hemodynamic parameters for the first and second diastolic phases, (i) a step of calculating the representations of the first and second diastolic cardiac performance based on the mathematical representation, (ii) the first operating parameter, and (iii) the second operating parameter. The method according to any one of claims 43 to 44, further comprising:

46. Calculating cardiac performance is The first waveform representing the hemodynamic parameters during the first diastolic phase is decomposed using a Fourier transform to determine a first set of sinusoidal curves, The second waveform representing the hemodynamic parameters during the second diastolic phase is decomposed using a Fourier transform to determine a second set of sinusoidal curves, The first set of sinusoidal curves and the second set of sinusoidal curves are compared to determine the contribution of the patient's heart to blood flow in the aorta. The method according to claim 45, including the method described in claim 45.

47. The method according to claim 46, wherein the mathematical expression is an exponential equation based on the comparison of the sine curves.

48. The method according to claim 47, further comprising the step of simulating an ideal heart rate representing the patient's cardiac function based on the comparison of the sine curves.

49. The process of calculating the aforementioned index representing cardiac performance is as follows: Based on the mathematical representation of the hemodynamic parameters over time, the vascular compliance and vascular resistance of the heart are determined, Using the determined vascular compliance and vascular resistance, the cardiac output is calculated. The method according to any one of claims 37 to 48, including the method described in any one of claims 37 to 48.

50. The method according to any one of claims 37 to 49, wherein the hemodynamic parameter is aortic pressure.

51. The method according to any one of claims 37 to 50, wherein the step of changing the first pump speed to a second pump speed includes increasing the pump speed.

52. The method according to any one of claims 37 to 50, wherein the step of changing the first pump speed to a second pump speed includes reducing the pump speed.

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