HK1139341B - Apparatus for monitoring and optimizing blood circulation generated by a pump - Google Patents

Description

Device for monitoring and optimizing blood circulation caused by a pump

Technical Field

The invention relates to a method for automatically monitoring and optimizing the blood flow of a pump. In this case, the patient condition and possible problems in the blood circulation caused by the pump can be inferred from the control parameters. In this case, a sufficient blood flow is ensured at all times.

The invention further relates to a medical device for controlling a pump according to the method.

Background

Blood pumps and associated control devices have long been known from the prior art. The blood pump in conjunction with its associated control device has the task of ensuring a blood circulation in vivo or in vitro and thus supporting or replacing the pump function of the heart of the patient. Such blood pumps are used for a defined period of time to replace the heart function during cardiac surgery or in extracorporeal circulation to support the heart function as a weak cardiac recovery solution. All customary products for extracorporeal circulation, such as cannulas, tubes, hoses and connectors, storage tanks, oxygenators, heat exchangers, hemoconcentrators and dialysers, bubble traps and filters, can be used outside the blood pump as required.

Furthermore, the use of cardiac support indefinitely, as well as an artificial heart, in which case the implantation of the pump is performed, is disclosed.

In all these applications of blood pumps, it is of utmost importance that a sufficient blood flow is reliably achieved by the pump. In particular, when using blood pumps for long periods of time, it is not possible to monitor the blood flow and, if necessary, to correct parameters associated therewith, such as the venous or arterial blood pressure or the gas flow and gas mixture when using an oxygenator at the same time.

Therefore, a reliable automatic regulation of the blood pump is required especially for the use of ECMO (extracorporeal membrane oxygenation) or ELS (extracorporeal life support), since these are applied outside the operating room and are performed over a long period of time from days to weeks without intensive monitoring of personnel. However, even when applied in cardiac surgery, automatic pump regulation provides greater comfort in operation and increases reliability in the event of disturbed flow. Furthermore, such regulation systems can also be used in implantable support systems (ventricular assist devices, VAD's) or in implantable artificial hearts.

Factors are discussed below that affect the desired blood flow rate, necessitating changes in pump settings and possibly additional parameters.

From practice and from Baloa et al Vollkron et al it is disclosed that in the event of an insufficient venous return of blood, the large venous vessels and/or the atrium of the heart (depending on which atrium is drained from, the left or right atrium) collapse and subsequently either no blood flow at all or only an insufficient blood flow is generated. In addition, the venous return depends on the filling state and position and structure of the venous line. When the filling state is too small or when the catheter opening comes into contact with the vessel wall, the vessel can collapse and prevent venous and complete blood flow. This phenomenon is disclosed in connection with centrifugal pumps, however, it may also occur independently of the pump principle. Furthermore, a hindrance to venous return caused by a thrombus or similar phenomenon or a break of the flexible tube in the venous line may be considered. If this condition of insufficient venous return occurs, damage to the blood caused by cavitation and bubble formation may occur in addition to the insufficient blood flow. Such life-threatening and patient-endangering conditions must therefore be reliably avoided.

Monitoring the pressure on the venous line is not sufficient to avoid collapse, since negative pressure can occur not only in the event of collapse, but also in the case of reliable operation. Furthermore, the pressure on the venous side of the blood pump depends to a large extent on its position relative to the patient (hydrostatic pressure) and can thus influence this pressure.

Furthermore, in this way, it is possible to identify, at best, the collapse that has occurred, but it is difficult to identify the tendency for collapse to occur.

In arterial lines, only obstacles or occlusions caused by, for example, thrombi or emboli or tube fractures can be considered as damage to the blood flow. Due to the pressure of the artery, there is no problem of collapse of the vessel wall. The magnitude of the arterial pressure alone is unsuitable as a measure for detecting flow disturbances, since high arterial flow resistance can be caused by so-called flow disturbances in the arterial line or high vascular resistance.

Methods for mathematically characterizing this problem are disclosed in the prior art. Baloa et al disclose a method of forming a differential DRI (regressive Return Index) DRI ═ dQ/d ω, for example, by stepping up the pump speed, measuring the corresponding flow rate, and differentiating according to the pump speed. In the ideal case of an unimpeded flow, this characteristic variable is a constant. With the onset of collapse, DRI decreases to reach a value of 0 at the time of collapse. However, there is no disclosure of how to avoid a possible collapse or how to eliminate an already occurring collapse.

The methods implemented in some VAD's present on the market become full of danger due to a sudden complete stop of the pump function and thus of the circulatory support, in the case of which the pump is simply switched off for a few seconds. In such cases, the heart of a weakened patient must therefore suddenly assume all the loads of the blood circulation without adaptation processes, which can lead to its overloading and thus to additional damage. In addition, when the rotational speed is set back to the original value, the collapse can occur again at any time with the described consequences. Alternatively, if a smaller pump speed is preset for safety reasons, the resulting smaller flow rate may remain stable. However, this is a reduced flow relative to the basic setting, which does not lead to a completely intentional loop support. Further, the application described by Baloa et al is applied in VAD, where not blood vessels but ventricles are collapsed.

Vollkron et al ascertains venous return from pulses in the flow signal, which is only given when the heart is beating. In the case of a too high pump speed, the pump speed is slowly reduced to a minimum value, and then the speed is slowly accelerated again to a safe or desired maximum flow. However, this slow reduction and acceleration requires a certain time span in which the pump searches for the optimum flow rate, but in which optimum support is not ensured.

Disclosure of Invention

Against this background, the object of the invention is to provide a method for automatically regulating and optimizing the blood flow of a blood pressure pump in the range of desired target values while avoiding life-threatening states, wherein the method should be universally applicable, i.e., reliably effective even when the heart is stopped, and should be able to identify the position of the lesion in the circulation in the vein or artery with the lesion of the blood flow being controlled, so that rapid measures can be taken in a targeted manner.

This object is achieved completely by the method according to the invention and the device for carrying out the method.

The new method makes it possible to automatically intervene periodically on the pump speed, wherein differentiated parameters are determined by recording measured values of pump parameters, such as speed and blood flow rate per unit time, and their ratio to one another, in order to determine the influence on the flow, said parameters characterizing the flow state and being suitable for detecting and eliminating back flow disturbances and for analyzing the heart condition.

The blood flow required according to the patient's requirements is determined in a known manner by the treating physician or cardiac perfusion specialist, and the quantity required for the method according to the invention "3/4 for the blood flow required in the circulation" is also disclosed. The required blood flow may be calculated, for example, by the du Bois formula for the body surface of the patient, where it is between 2.5 and 41min^-1m^-2Cardiac indices within the range were normal.

Contrary to the solution of Baloa et al, the pump speed can be adjusted up and down and a varying blood flow can be measured. In this way, when a tendency toward collapse is detected, sufficient venous return can be achieved again by reducing the pump speed, while the beginning collapse is eliminated. The optimization of the rotational speed and the flow can then be achieved stepwise by increasing the rotational speed and measuring the changing flow (blood flow), wherein stable values above the initial value at which the collapse starts to occur can be obtained.

The method according to the invention not only makes it possible to recognize the beginning of a collapse, but also to quickly eliminate it and subsequently optimize the flow, in contrast to the prior art (Baloa et al, Vollkron et al), which, although also determining a differential value, does not disclose a solution for eliminating the beginning or existing collapse. The methods implemented in some VAD's present on the market become full of danger due to a sudden complete stop of the pump function and thus of the circulatory support, in the case of which the pump is simply switched off for a few seconds. In such cases, the heart of a weakened patient must therefore suddenly assume the full load of the blood circulation without adaptation processes, which can lead to its overloading and thus to additional damage.

On the other hand, the elimination of the collapse disclosed by Vollkron et al requires a certain time before the pump speed is once reduced and then increased again. The principle according to the invention is distinguished in particular here by a rapid pump throttling and a subsequent increase in the pump speed, wherein this function can be achieved even without a pulsating blood flow, for example when the heart is stopped.

Drawings

The method according to the invention and the device with which the method according to the invention can be carried out are described with reference to the following figures.

Wherein:

FIG. 1 is a schematic of a minimized extracorporeal bypass with pump, oxygenator and hose lines;

FIG. 2 is a flow chart with an illustration of the main steps of the algorithm;

FIG. 3 is a diagram of a control loop according to the invention, in which the setpoint value is a range of permissible values of the differential flow rate ratio (DFSR), which enables control of the Pump Speed (PS) for the regulation and thus generates a flow rate, wherein the maximum flow rate also influences the DFSR regulation, that is to say prevents over-filling;

FIG. 4 is a graphical relationship between varying pump power and a derived DFSR value; and

fig. 5 is a diagrammatic illustration of the adjustment sequence, as it can be realized in accordance with the invention.

Detailed Description

In fig. 1, a patient is indicated generally at 10, who is connected to an extracorporeal blood circuit comprising a pump 12, an oxygenator 14, a venous hose line 16 and an arterial hose line 18.

In the initial phase of the minimized closed extracorporeal circulation system (fig. 1), the Pump Speed (PS) is increased stepwise α U/min. In this case, Δ is continuously passed_{Flow rate}/Δ_{Rotational speed}(. DELTA.FLOW/. DELTA.PS [ l/min).][ rotation/min]) The value of the "differential flow rate ratio" (DFSR) is calculated until 3/4 of the required flow rate is reached. The median of the calculated differential is an approximation of the flow to rotational speed ratio for each individual patient with circulatory support. This value is divided by two and used as (defined as) a threshold T. In the bypass, the pump speed (pump speed, PS) decreases by β U/min during ε seconds. The flow rate change (dF) is calculated from the pump speed reduction (dPS), and a differential value differential flow rate ratio (DFSR), i.e. delta flow rate/delta PS, is determined therefrom.

DFSR ≦ 1T indicates excessive drain, i.e., too little backflow, and a sudden pump speed PS downshift to 50% of the original pump speed is required. After γ seconds, PS is stepped up δ U/min until 95% of the initial pump speed PS. DFSR > 2T indicates too little discharge, which requires an increase in the pump speed PS by ε U/min. Calculating a new DFSR indicates that reflow optimization is achieved.

The value of 1.5T is specified as the upper threshold. DFSR > 1.5T indicates that higher pump speeds and higher flow rates are achieved. DFSR > 1.5T allows the pump to increase the speed by ζ U/min when the option is activated. Calculating a new DFSR indicates that reflow optimization is achieved. If a possible over-filling is to be avoided, a maximum flow rate may be input, which has priority over the rotational speed increase.

A schematic representation of the above described control unit is shown in fig. 2 and 3. Fig. 4 shows the results of the practical application.

Fig. 2 shows a flow chart of the process (algorithm) according to the invention for achieving an optimal discharge. After the extracorporeal blood circulation (EKZ) has been initiated, the necessary limit values for the patient are determined, as is illustrated in an exemplary manner in fig. 1. The venous return of blood is then quantified by varying the pump power, and the pump power for optimal drainage is determined and set therefrom.

Fig. 3 shows a control loop according to the invention, which continuously determines a new actual DFSR value from the instantaneous flow of the pump speed by means of the desired DFSR value, corrects the actual DFSR value with the desired DFSR value and thus continuously readjusts the pump speed, so that excessive and too low discharge is prevented.

Fig. 4 shows how the median value (instantaneous DFSR) is formed from the acceleration of the pump and the simultaneous measurement of the DFSR. Half of this value is taken as the limit value. Venous return was reduced (shown by arrows) after the DFSR was determined. The following measurement for determining the DFSR determines the reduction in the DFSR, which is then effectively reduced to 95% by means of a 2-step reduction. Subsequent measurements of DFSR clearly show an increase in this value and optimization of drainage, indicating that the pump speed is matched to the venous flow for use.

Furthermore, method steps a and b in method claim 1 are depicted in fig. 4, wherein method step d is illustrated by a further arrow.

The backflow barrier increases the flow resistance on the pump inlet side, which reduces the pump preload and negatively affects the pump flow. In the case of oxygenator coagulation, the resistance is also increased, but this additional resistance occurs behind the pump. Both of these conditions affect the DFSR, but additionally measuring the pressure drop across the oxygenator (pressure at inlet-pressure at outlet) can provide the required information in order to distinguish between coagulation of the oxygenator and obstruction of the pump by venous return. If the DFSR decreases, the magnitude of the pressure differential across the oxygenator indicates the cause of the increased resistance, i.e., whether it is before or after the pump. If the pressure difference is small, there is an obstacle to venous return, and if the pressure difference is large, coagulation of the oxygenator occurs.

Furthermore, measuring the DFSR provides information about the pump capacity of the heart, in particular the right ventricle. The DFSR is not changed in case of cardiac recovery and thus a reduction of the average pump speed. The restored heart responds to the increased preload according to the Frank-Starling mechanism and attempts to achieve the required ejection power. The venous as well as arterial pressure and thus the DFSR is kept constant. Reducing the average pump speed results in an increase in DFSR when the heart is not fully or only partially restored. An incompetent heart is not able to cope with increased preload if pump support is reduced. Thus, in this case the venous pressure increases and the arterial pressure decreases. The body again responds to the reduced arterial pressure with vasoconstriction, balancing the further drop in arterial pressure. Thereby, the pump preload is increased, while the afterload remains almost constant: the DFSR is increased.

The response to the rotational speed-reducing regulation step can be used in conjunction with flow pattern analysis to ascertain the heart condition. The beating heart influences the flow pattern of the centrifugal pump, which is supportive, i.e. the arterial pressure increases when the heart contracts, and the flow through the pump decreases in the process. If the average pump speed is reduced and a subsequent regulation step of reducing the speed is performed, the following pulses of the flow signal give valuable information about the heart condition and its pump reserve. After reduced support, the slope and amplitude of the flow pulses give direct information about contractility and cardiac recovery.

The above method can be used to determine the heart condition and the reserve of pump capacity and can be used very significantly to let the patient abandon the habit of heart support. The method can be very usefully used as a tool for cardiac monitoring, which obtains a greater preload by periodically reducing the rotational speed and in the process gets exercised.

A further approach to ascertaining the heart state and heart recovery is to determine the time required by the heart to balance sudden pump support and provide the total flow required by the body, including pump flow and heart drain power.

The Frank-Starling mechanism is used directly in both methods. Unlike non-invasive determination of DFSR, however, this application requires invasive measurement of arterial and venous blood pressure.

In this way, monitoring and optimization of the blood flow can be automated, which is necessary for long-term use for cardiac support and cardiac recovery. Furthermore, signs of cardiac recovery can be ascertained by the heart's reaction to a decrease in blood flow through the pump. The higher the pulsation in the arterial blood pressure when the pump power is reduced, the greater the power of the heart and thus the restitution of the heart. The required support can be given automatically during the time of the heart support, and in the event of a reduction of the pump power of the heart, the pump can automatically assume a greater part of the pump power. In this way, a risk-free recovery of the heart is achieved, and the habit of using the pump is then automatically and steplessly discarded.

Fig. 5 shows an example of how the regulation according to the invention can be implemented. After the required blood flow has been determined, the regulation method is started and the DFSR is determined, which is conventional to the person skilled in the art. The fraction of DFSR, here half of DFSR in the example, is defined as the threshold T as described.

In fig. 5, a straight line with DFSR equal to 1T is denoted by 22. A curve with DFSR 1.5T is denoted by reference numeral 24, and a curve with DFSR 2T is denoted by reference numeral 26. Between the lines 22 and 24, the operating region 28 of the method according to the invention is illustrated. Below the straight line 22(DFSR ═ IT), the undesired operating region is marked as a gray region 30(DFSR < 1T), in which a partial or complete collapse can occur. This area is critical for patient safety and must therefore be reliably avoided.

The pump speed is increased at DFSR ≧ 1.5T, and reduced (and subsequently increased) at DFSR ≦ 1T. Thereby achieving stable operation in the operation region between 1T and 1.5T.

The starting points of the straight line DFSR-1. 1T, DFSR-1.5T and DFSR-2T do not start from the origin in the above example diagram. It can be understood that, depending on the pressure conditions prevailing during use, a positive flow > 0 is only caused by the pump in the centrifugal pump illustrated by way of example at a given rotational speed PS, and thus the determination of DFSR can only be started with this rotational speed and ended when 3/4 of the desired flow is reached. To determine DFSR and thus T, the rotational speed value (PS) occurring in the pump cycle and the (positive) flow rate caused by the pump are therefore evaluated > 0 until 3/4 of the desired flow rate is reached.

For general understanding, it is again pointed out that the DFSR is formed by a sequence of values of Δ flow/Δ rotation speed for different rotation speeds. However, this value is understood as a differential value. The DFSR is therefore the first derivative of the corresponding function (slope Δ flow/Δ rotation speed of the function at each point).

For the unchanged backflow situation, i.e. for an almost ideal unimpeded blood flow, it is expected that the function is a straight line and a constant value DFSR is obtained therefrom as a derivative. This relationship is used to determine the instantaneous DFSR and thus the threshold T.

A change in this slope over the range of consideration of the pump speed also means that a change in the value of the DFSR indicates a change in the flow conditions.

In the method according to the invention, a periodic rotational speed intervention is carried out starting from a given pump rotational speed, and the differential value DFSR of the intervention region is formed from the Δ flow/Δ rotational speed of the upper and lower rotational speeds considered for the intervention.

By comparison with a threshold value I (for example a fraction of the instantaneous DFSR, 1/2 of the instantaneous DFSR in special cases), the pump speed is set in such a way that the flow conditions in a predetermined region of the DFSR are observed. In this exemplary embodiment, this predefined region of T to 1.5T characterizes the pump speed which is adapted to the available return volume, wherein on the one hand an excessive drainage is avoided and on the other hand a predefined desired blood flow is achieved as well as possible under varying drainage conditions.

The pump speed is therefore adapted to the flow rate at a predetermined target value for the maximum blood flow rate, instead of the usual control method for the blood flow rate.

In a development of the method, the gas supply system of the oxygenator in the blood circuit can be integrated into the control loop. Since the blood flow is variably set by the pump regulating system, the gas flow and the oxygen content of the gas supply system must also be adapted automatically to the respective blood flow and to the patient's requirements. For this purpose, it is proposed to use the existing blood flow and the partial pressure pO of the blood gas measured on-line₂And pCO₂For regulating the gas input system. If these two partial pressures are normal, the gas mixture and the gas flow rate can be maintained. If pO₂If the oxygen content is too low, the oxygen content of the supplied gas mixture is automatically increased and, conversely, the oxygen content is reduced. Conversely, pCO, which is too high, is adjusted by increasing the gas flow₂Value, pCO adjusted too low by reducing gas flow₂The value is obtained. The algorithm may set a threshold or limit value similar to the algorithm proposed for blood flow.

In this case, it is particularly expedient to automatically detect the pressure difference across the oxygenator and to issue an alarm when a threshold value for the pressure difference or the partial gas pressure is reached. The risk of an initial short supply of oxygen to the heart and lung function of the patient is thus known and eliminated by appropriate measures, for example, the replacement of the oxygenator with the condensate.

Even though only the measured values of a centrifugal pump are disclosed in fig. 4, it is obvious to the person skilled in the art that in principle the problem of venous return and the risk of collapse of the blood vessels or ventricles can occur when using blood pumps of various operating principles and that said problem can be solved according to the invention.

It is understood that the specified flow rate of the pump and a characteristic variable of the pump proportional to this flow rate can be used as parameters in the non-invasive method described above, for example the power consumption or current consumption of a centrifugal pump or the current or power consumption of an active magnetic bearing of an axial pump can be used as a control variable without leaving the scope of the invention.

In the method for automatically regulating a blood pump, the blood flow is optimized by means of a periodic rotational speed intervention and the flow rate change occurring there, by means of the differential values formed and the regulating algorithm. It is additionally possible to ascertain the location of a possible flow resistance on the venous or arterial side.

The method for regulating the blood pump is characterized in that the change of the pump delivery power measured by the automatic periodic change of the rotation speed or the pump speed is based on delta_{Transmitting power}/Δ_{Rotational speed/pump speed}A differential characteristic parameter differential flow rate ratio (DFSR) is formed, the magnitude and variation of which is used in an algorithm that optimizes blood flow.

The method is characterized in that the automatic periodic speed change or pump speed change can be both an increase and a decrease.

Method, characterized in that the change of the delivered power produced by the change of the pump parameters is determined.

The method is characterized in that the change in the rotational speed or the change in the pump speed and the change in the pump delivery power measured for it are determined noninvasively.

The method is characterized in that at least one parameter is determined in an invasive manner.

Medical devices contain a pump with a regulating system for delivering blood.

Claims (4)

1. Device for operating a blood pump located inside an intracorporeal or extracorporeal blood circulation, characterized in that it comprises: by using the device

a) Determining a threshold value T which is a fraction of the median value of the measured characteristic variables DFSR, wherein the median value is an approximation of the flow/speed ratio for each individual patient with circulatory support, wherein the respective DFSR characteristic variable is calculated from the ratio of the differentiated blood supply power to the differentiated blood pump speed measured for this purpose, wherein the blood pump speed is increased stepwise to 3/4 of the blood flow required in the circulation, and the DFSR characteristic variable is generated for each step of increasing the blood pump speed;

b) reducing the rotation speed of the blood pump and gradually calculating the DFSR characteristic parameter matched with the rotation speed;

c) comparing the DFSR characteristic parameter in b) with the threshold value T in a);

d) varying the blood pump speed depending on whether the characteristic DFSR in b) is greater or less than a threshold value T,

wherein the device has a first means for detecting, calculating and comparing the determined DFSR characteristic variable, and a second means of the device is provided for adjusting the blood pump rotational speed as a function of the result of the comparison of the detected DFSR characteristic variable with a threshold value T.

2. The device according to claim 1, characterized in that the threshold value T is half the median value of the measured characteristic DFSR detected until 3/4 of the blood flow required in the circulation is reached.

3. Device according to claim 1 or 2, characterized in that the blood pump speed is reduced when DFSR < 1T and then increased again for a defined time.

4. The device of claim 3, wherein the blood pump speed is increased at a DSFR ≧ 1.5T.