METHOD FOR DETERMINING A FLOW RATE OF A FLUID FLOWING THROUGH AN IMPLANTED VASCULAR SUPPORT SYSTEM, AND IMPLANTABLE VASCULAR SUPPORT SYSTEM

Abstract

The invention relates to a method for determining a flow rate of a fluid flowing through an implanted vascular assist system (1), said method comprising the following steps: a) carrying out a first pulsed Doppler measurement at a first pulse repetition rate by means of an ultrasonic sensor (2) of the assist system (1); b) carrying out a second pulsed Doppler measurement at a second pulse repetition rate by means of the ultrasonic sensor (2) of the assist system (1), wherein the second pulse repetition rate differs from the first pulse repetition rate; c) determining the flow rate using measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement.

Claims

1.-11. (canceled)

12. A method for determining a flow rate v of blood flowing through a cardiac assist system, comprising: performing a first pulsed Doppler measurement at a first pulse repetition rate PRF.sub.1 using an ultrasonic sensor of the cardiac assist system; performing a second pulsed Doppler measurement at a second pulse repetition rate PRF.sub.2 using the ultrasonic sensor of the assist system, wherein the second pulse repetition rate PRF.sub.2 is greater than the first pulse repetition rate PRF.sub.1; and determining the flow rate using measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement by solving for an integer n.sub.1, an integer n.sub.2, and a main component f.sub.1 of the first pulsed Doppler measurement and a main component f.sub.2 of the second pulsed Doppler measurement using a linear Diophantine equation, wherein the linear Diophantine equation comprises:
n.sub.1.Math.PRF.sub.1=n.sub.2.Math.PRF.sub.2=f.sub.1−f.sub.2 for a boundary condition of: $v<\frac{a.Math.{\mathrm{PRF}}_1.Math.c_0}{2f_0}$ assuming
−|α|≤n.sub.1≤|α| and −|b|≤n.sub.2≤|b|, wherein $a:=\frac{PR{F}_2}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)}\mathrm{and}b:=\frac{PR{F}_1}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)},$ and wherein f.sub.0 comprises the ultrasonic transmission frequency of the ultrasonic sensor and c.sub.0 comprises a speed of sound in the fluid.

13. The method of claim 12, wherein at least one of performing a first pulsed doppler measurement and performing a second pulsed doppler measurement comprises emitting a new ultrasonic pulse using an ultrasonic element only after an echo of an immediately previously emitted ultrasonic pulse has been received by the ultrasonic sensor.

14. The method of claim 12, wherein PRF.sub.1 or PRF.sub.2 is smaller than twice a maximum occurring Doppler shift.

15. The method of claim 12, wherein determining the flow rate comprises using a correlation between a detected main frequency component of the Doppler frequency spectrum of the first pulsed Doppler measurement and the first pulse repetition rate PRF.sub.1 or the second pulsed Doppler measurement and the second pulse repetition rate PRF.sub.2.

16. The method of claim 12, wherein determining the flow rate comprises solving the linear Diophantine equation using Bezout coefficients or an exhaustion method.

17. The method of claim 12, wherein an observation window of the ultrasonic sensor is in a range of 25 mm to 55 mm from an ultrasonic element of the ultrasonic sensor.

18. A method for determining a fluid flow through a cardiac assist system comprising: determining a flow rate v of blood flowing through the cardiac assist system, wherein determining the flow rate v comprises: performing a first pulsed Doppler measurement at a first pulse repetition rate PRF.sub.1 using an ultrasonic sensor of the cardiac assist system; performing a second pulsed Doppler measurement at a second pulse repetition rate PRF.sub.2 using the ultrasonic sensor of the assist system, wherein the second pulse repetition rate PRF.sub.2 is greater than the first pulse repetition rate PRF.sub.1; and determining the flow rate using measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement by solving for an integer n.sub.1, an integer n.sub.2, and a main component f.sub.1 of the first pulsed Doppler measurement and a main component f.sub.2 of the second pulsed Doppler measurement using a linear Diophantine equation, wherein the linear Diophantine equation comprises:
n.sub.1.Math.PRF.sub.1=n.sub.2.Math.PRF.sub.2=f.sub.1−f.sub.2 for a boundary condition of: $v<\frac{a.Math.{\mathrm{PRF}}_1.Math.c_0}{2f_0}$ assuming
−|α|≤n.sub.1≤|α| and −|b|≤n.sub.2≤|b|, wherein $a:=\frac{PR{F}_2}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)}\mathrm{and}b:=\frac{PR{F}_1}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)},$ and wherein f.sub.0 comprises the ultrasonic transmission frequency of the ultrasonic sensor and c.sub.0 comprises a speed of sound in the fluid; and determining the fluid flow in the cardiac assist system based on the flow rate v.

19. A cardiac assist system comprising: an ultrasonic sensor configured to perform a first pulsed Doppler measurement at a first pulse repetition rate PRF.sub.1 and a second pulsed Doppler measurement at a second pulse repetition rate PRF.sub.2, wherein PRF.sub.2 is greater than PRF.sub.1; and a processing unit configured to: determine a flow rate of a fluid flowing through the cardiac assist system using measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement, wherein determining the flow rate using the measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement comprises solving for an integer n.sub.1, an integer n.sub.2 and a main component f.sub.1 of the first pulsed Doppler measurement and a main component f.sub.2 of the second pulsed Doppler measurement using a linear Diophantine equation comprising:
n.sub.1.Math.PRF.sub.1=n.sub.2.Math.PRF.sub.2=f.sub.1−f.sub.2 for the boundary condition: $v<\frac{a.Math.{\mathrm{PRF}}_1.Math.c_0}{2f_0}$ assuming:
−|α|≤n.sub.1≤|α| and −|b|≤n.sub.2≤|b|, wherein $a:=\frac{PR{F}_2}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)}\mathrm{and}b:=\frac{PR{F}_1}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,\mathrm{PR}{F}_2\right)},$ and wherein f.sub.0 is the ultrasonic transmission frequency of the ultrasonic sensor and c.sub.0 is a speed of sound in the fluid.

20. The cardiac assist system of claim 19, wherein the processing unit is configured to solve the linear Diophantine equation using Bezout coefficients or an exhaustion method to determine the flow rate.

21. The cardiac assist system of claim 19, wherein the processing unit is configured to calculate a fluid flow based on the flow rate.

22. The system of claim 19, wherein the ultrasonic sensor comprises an observation window in a range of 25 mm to 55 mm from an ultrasonic element of the ultrasonic sensor.

23. The system of claim 19, further comprising a cannula, wherein the ultrasonic sensor is configured to perform the first pulsed Doppler measurement and the second pulsed Doppler measurement within the cannula.

24. The system of claim 23, wherein the ultrasonic sensor is integrated into the tip of the cannula.

Description

[0039] The solution presented herein and the technical environment thereof are explained in greater detail below with reference to the figures. It should be noted that the invention is not to be limited by the exemplary embodiments shown. In particular, unless explicitly stated otherwise, it is also possible to extract partial aspects of the situation explained in the figures and to combine them with other components and/or findings from other figures and/or the present description. The drawings show the following in schematic form:

[0040] FIG. 1 an implanted vascular assist system in a heart,

[0041] FIG. 2 the assist system from FIG. 1,

[0042] FIG. 3 a sequence of a method presented here in a normal operating mode,

[0043] FIG. 4 an exemplary Doppler frequency spectrum; and

[0044] FIG. 5 another exemplary Doppler frequency spectrum.

[0045] FIG. 1 shows in schematic form an implanted vascular (here: ventricular) assist system 1 in a heart 6. The assist system 1 assists the heart 6 by helping to convey blood out of the (left) ventricle 7 into the aorta 8. For this purpose, the assist system 1 is anchored in the aortic valve 9, as illustrated by way of example in FIG. 1. At a level of assist of 100%, the assist system 1 (LVAD) conveys the complete blood volume flow. The level of assist describes the proportion of the volume flow conveyed through a delivery means such as a pump of the assist system 1 or through the assist system 1 to the total volume flow of blood from the ventricle 7 to the aorta 8.

[0046] Therefore, at an assist level of 100%, the total fluid volume flow 10 from the ventricle 7, the volume flow 11 from the heart valve into the ventricle 7, and the fluid volume flow 5 through the assist system 1 are identical. In this case, it follows that the aortic valve volume flow or bypass volume flow 12 (formula symbol: Q.sub.a) is zero. The total fluid volume flow 10 can also be described as the (total) heart-time volume (HTV, formula symbol: Q.sub.HTV). The fluid volume flow 5 can also be referred to as a so-called pump volume flow (formula symbol: Q.sub.p), which quantifies only the flow through the assist system 1 itself. As a result, the level of assist can be calculated from the ratio Q.sub.P/Q.sub.HTV.

[0047] In the case of lower levels of assist and healthier hearts with a strong ventricular contraction, the heart 6 continues to fulfill its function to a certain extent, so that during systole (heart muscle contracts and pushes the blood out into the aorta 8 as a result of the decrease in volume of the ventricle 7), a pulsatile volume flow component 12 (bypass) is produced by the heart valve or the aortic valve 9. At the same time, the pressure difference in the assist system 1, in particular in the pump typically provided (not shown here) in the assist system 1, drops, so that correspondingly the assist system 1 also conveys an increased amount of fluid volume flow 5 during systole.

[0048] FIG. 2 shows in schematic form the assist system 1 from FIG. 1. The assist system 1 comprises an ultrasonic sensor 2, which is designed to carry out pulsed Doppler measurements at different pulse repetition rates, and a processing unit 3 that is designed to determine a flow rate of a fluid (here: blood) flowing through the assist system 1, using the measurement results of the pulsed Doppler measurements at different pulse repetition rates.

[0049] In addition, FIG. 2 also shows by way of example that the ultrasonic sensor 2 can be integrated in the tip of a cannula 13 of the assist system 1. The ultrasonic sensor 2 helps to determine the flow rate (amount and at least one direction) of a fluid or fluid volume flow 5, which flows through the assist system 1 and which is also referred to as a pump volume flow (Q.sub.P). For this purpose, the ultrasonic sensor 2 is designed to carry out pulsed Doppler measurements in the fluid inside the cannula 13. The fluid can enter the interior of the cannula 13 through one or more inlet openings 15 (from the ventricle 7) and exit through one or more outlet openings 16 (into the aorta 8). In order to assist the fluid flow through the assist system 1, in particular through the cannula 13, the assist system 1 includes here a continuous flow machine 17. The continuous flow machine 17 is generally formed in the manner of a pump. Furthermore, an observation window or a range of measurement 18 of the ultrasonic sensor 2 is also shown by way of example in FIG. 2.

[0050] FIG. 3 shows in schematic form a sequence of a method presented here in a normal operating mode. The method is used to determine a flow rate of a fluid flowing through an implanted vascular assist system 1 (see FIGS. 1, 2). The illustrated sequence of process steps a), b), and c) with the blocks 110, 120, and 130 is shown merely for illustrative purposes. In block 110, a first pulsed Doppler measurement is carried out at a first pulse repetition rate by means of an ultrasonic sensor 2 of the assist system 1. In block 120, a second pulsed Doppler measurement is carried out at a second pulse repetition rate by means of the ultrasonic sensor 2 of the assist system 1, where, in this case, the second pulse repetition rate differs from the first pulse repetition rate. In block 130, the flow rate is determined using the measurement results of the first pulsed Doppler measurement and the second pulsed Doppler measurement.

[0051] The following parameters are assumed for an exemplary representation of the method: [0052] diameter inlet region or range of measurement, e.g. 5 mm, [0053] maximum blood flow to be measured, e.g. Q=9 l/min., [0054] resulting maximum blood flow rate: v.sub.blood,max=7.64 m/s, [0055] speed of sound in the blood, e.g. c.sub.blood=1,540 m/s, [0056] ultrasonic frequency, e.g. f.sub.0=6 MHz, [0057] distance of the ultrasonic element from the start of the observation window, e.g. 25 mm, [0058] number of ultrasonic vibration cycles per emitted ultrasonic PWD pulse, e.g. 10, [0059] resulting burst length (in distance): I.sub.burst=c.sub.0×10/f.sub.0=2.57 mm, [0060] resulting maximum propagation distance ultrasonic burst: d=55.13 mm.

[0061] The above data gives the following (expected) maximum Doppler shift for a measurement directly in the emission direction (flow direction corresponds to the main emission direction; α=0):

[00001]$\begin{array}{cc}\mathrm{df}=\frac{2.Math.v_{\mathrm{blood},\max }.Math.f_0}{c_0}=\frac{2.Math.7.64\frac{m}{s}.Math.6\mathrm{MHz}}{1540\frac{m}{s}}=59.53\mathrm{kHz}& \left(1\right)\end{array}$

[0062] The measurement should be carried out as a pulsed Doppler measurement, in which a new ultrasonic pulse is not emitted until an echo of a just previously emitted ultrasonic pulse has decayed. The choice of the pulse repetition rate (PRF) to be used for this purpose is explained below.

[0063] Taking into account the (Nyquist) sampling theorem (which, however, does not have to be and is not taken into account in the solution presented here), a maximum Doppler frequency of 59.53 kHz would mean that a minimum pulse repetition rate or a minimum pulse repetition frequency of

PRF.sub.min=2.Math.df=119.06 kHz,  (2)

[0064] would have to be observed.

[0065] However, in the implanted vascular assist systems that are the focus of the present invention, the following maximum pulse repetition rate PRF-.sub.max, is calculated from the geometric consideration (maximum propagation distance of the ultrasonic pulse) or the geometric boundary conditions in the assist system and the resulting transit time of all of the relevant signal components:

[00002]$\begin{array}{cc}{\mathrm{PRF}}_{\max }=\frac{c_{\mathrm{blood}}}{d}=27.93\mathrm{kHz}& \left(3\right)\end{array}$

[0066] Therefore, the maximum pulse repetition rate of the pulsed Doppler measurements here (i.e. for the assist systems that are the focus of the present invention) is smaller than twice the maximum occurring Doppler shift.

[0067] These boundary conditions lead to a violation of the sampling theorem and, consequently, to an ambiguity of the measurement results that can be corrected by an evaluation, as described in the following sections.

[0068] However, in order to illustrate the problems arising from these boundary conditions, the ambiguity occurring in this case is illustrated in FIGS. 4 and 5 (said ambiguity can be resolved with the solution presented here). FIG. 4 shows in schematic form an exemplary Doppler frequency spectrum 4. FIG. 4 shows a Doppler shift at a pulse repetition rate of approximately 25 kHz. The main frequency component 19 (peak) is below the carrier frequency at approximately 0 Hz. FIG. 5 shows in schematic form another exemplary Doppler shift frequency spectrum 4. FIG. 5 shows a Doppler shift at a pulse repetition rate of approximately 20 kHz. The main frequency component 19 (peak) is at approximately +8 kHz.

[0069] In the following sections, an exemplary evaluation of the ambiguous measurement results is described in the context of the solution proposed herein.

[0070] Two measurement cycles (sequence of a defined number of ultrasonic pulses emitted in succession) at different PRFs (with respect to the same fluid flow, for example in the same observation window) are included. The actual Doppler shift can be shown as a function of the resulting heart frequency components 19, here peaks f.sub.1 and f.sub.2:

df=f.sub.1+n.sub.1.Math.PRF.sub.1  (4)

df=f.sub.2+n.sub.2.Math.PRF.sub.2  (5)

[0071] This illustrates by way of example how a linear equation system can be set up, in which the Doppler shift df is shown as a function of main frequency components 19, here peaks f.sub.1 and f.sub.2, of the first pulsed Doppler measurement and the second pulsed Doppler measurement. In addition, this illustrates by way of example a correlation between a detected main frequency component 19 of the Doppler frequency spectrum, here a peak in the Doppler frequency spectrum, a pulsed Doppler measurement, and the pulse repetition rate applied for this Doppler measurement.

[0072] By resolving both equations after the Doppler shift df and subsequently equating, the following Diophantine equation is obtained:

n.sub.1.Math.PRF.sub.1=n.sub.2.Math.PRF.sub.2=f.sub.2−f.sub.1  (6)

[0073] This illustrates by way of example how a linear Diophantine equation can be set up on the basis of the linear equation system.

[0074] At the speed of sound in blood c.sub.blood, the ultrasonic transmission frequency f.sub.0, the flow rate of the blood v.sub.blood, and an integer pulse repetition rate, an unambiguous solution can be found for this equation.

[00003]$\begin{array}{cc}-.Math.a.Math.\le n_1\le .Math.a.Math.\mathrm{where}a=\frac{{\mathrm{PRF}}_2}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,{\mathrm{PRF}}_2\right)}& \left(7\right)\\ -.Math.b.Math.\le n_2\le .Math.b.Math.\mathrm{where}b=\frac{{\mathrm{PRF}}_1}{2.Math.\mathrm{ggT}\left({\mathrm{PRF}}_1,{\mathrm{PRF}}_2\right)}& \left(8\right)\\ \text{?}<\frac{a.Math.{\mathrm{PRF}}_1.Math.c_0}{2.Math.f_0}\mathrm{where}{\mathrm{PRF}}_1<{\mathrm{PRF}}_2\text{?}\text{indicates text missing or illegible when filed}& \left(9\right)\end{array}$

[0075] In this case, the operator ggT represents the largest common divisor. Within these ranges, this equation can be obtained, for example, with the aid of Bezout coefficients or with an exhaustion (brute force) approach.

[0076] This procedure is illustrated by way of example with the aid of the values v=8 m/s, f.sub.0=4 MHz, PRF.sub.1=11 kHz and PRF.sub.2=19 kHz. At these values, the resulting peaks were detected at f.sub.1=−2442 Hz and f.sub.2=3558 Hz. The resulting Diophantine equation is:

n.sub.1−11−n.sub.2.Math.19=6  (10)

[0077] The largest common divisor for this case is 1, and the Bezout coefficients are 7 and 4. This gives the following possible solutions:

n.sub.1=6.Math.7+m.Math.19=4+m.Math.19  (11)

n.sub.2=6.Math.4+m.Math.11=2+m.Math.11  (12)

[0078] This illustrates by way of example how the linear Diophantine equation can be solved using Bezout coefficients.

[0079] Because a unique solution can be determined only for m=0, only this is considered. It can be used to determine the frequency of the Doppler shift that is now no longer ambiguous. Both equations give the same result.

df=f.sub.1+n.sub.1.Math.PRF.sub.1  (13)

=−2442 Hz+4.Math.11000 Hz=41558 Hz  (14)

df=f.sub.2+n.sub.2.Math.PRF.sub.2  (15)

=−3558 Hz+2.Math.19000 Hz=41558 Hz  (16)

[0080] The flow rate of the fluid flowing through the assist system 1 (here: blood) can be calculated on this basis by means of the frequency shift due to the Doppler effect:

[00004]$\mathrm{df}=f_0.Math.\frac{2v}{c}.Math.\cos \left(\alpha \right)$

[0081] where df is the resulting (unique) Doppler frequency shift; f.sub.0 the frequency of the emitted ultrasonic pulse; v the flow rate of the medium (sought here); c the speed of sound in the medium; and a the angle between the ultrasonic sound path and the main flow direction.

[0082] In a (ventricular) assist system, v is sought, and α, f.sub.0 and c are generally known (at least approximately). It is possible to compensate for the ambiguity usually occurring in such (ventricular) assist systems as discussed above in a particularly advantageous way by means of the solution proposed herein. On the basis of the determined flow rate, the fluid volume flow through the assist system can be determined using the (known) geometric boundary conditions in the assist system (known cross section of the range of measurement or the observation region through which the fluid can flow). Said fluid volume flow can help, at least approximately, to detect the blood volume that is actually conveyed through a (heart) assist system. This knowledge of the volume of blood that is actually pumped by a ventricular assist system or a cardiac assist system is medically of great importance, in particular for controlling the (implanted) assist system.

[0083] The solution presented herein makes possible one or more of the following advantages, in particular: [0084] Pulsed Doppler measurement or PWD-based flow rate measurement or volume flow measurement is also made possible even with a large distance between the measurement window and the ultrasonic transducer. [0085] Resolution of the geometrically induced ambiguity of the Doppler shift on the basis of geometric boundary conditions in the assist system.