METHOD AND SYSTEM FOR DETERMINING THE SPEED OF SOUND IN A FLUID IN THE REGION OF AN IMPLANTED VASCULAR SUPPORT SYSTEM

Abstract

The invention relates to a method for determining the speed of sound in a fluid (1) in the region of an implanted, vascular support system (2), comprising the following steps: a) sending an ultrasonic signal (3) by means of an ultrasonic sensor (4) b) reflecting the ultrasonic signal (3) on at least one sound reflector (5), which is visible in the field of vision (6) of the ultrasonic sensor (4) and arranged at a defined distance at least to the ultrasonic sensor (4) or to a further sound reflector (5), c) receiving the reflected ultrasonic signal (8), d) determining the speed of sound in the fluid using the reflected ultrasonic signal (8).

Claims

1.-13. (canceled)

14. A method for determining the speed of sound in blood in the region of a cardiac support system, the method comprising: a) sending an ultrasonic signal by means of an ultrasonic sensor, b) reflecting the ultrasonic signal on at least one sound reflector, the at least one sound reflector being located in a field of vision of the ultrasonic sensor and at a defined distance from the ultrasonic sensor or a further sound reflector, c) receiving the reflected ultrasonic signal, and d) determining a speed of sound in the blood based on the reflected ultrasonic signal.

15. The method of claim 14, wherein the ultrasonic signal is reflected on at least two sound reflectors, the at least two sound reflectors each located at different distances from the ultrasonic sensor.

16. The method of claim 14, wherein the at least one sound reflector has an acoustic impedance greater than the largest acoustic impedance of the blood or less than the lowest acoustic impedance of the blood.

17. The method of claim 14, wherein the at least one sound reflector is configured to be embedded into an embedding material.

18. The method of claim 14, wherein the speed of sound is determined based on a pulse time of flight-based analysis algorithm.

19. The method of claim 14, wherein the speed of sound is determined based on an FMCW-based analysis algorithm.

20. The method of claim 19, wherein a beat frequency is determined.

21. The method of claim 14, wherein the at least one sound reflector is configured to project sound at least partially into a flow path of the blood formed by an inlet cannula.

22. The method of claim 21, wherein the at least one sound reflector comprises the inlet cannula.

23. A system for determining the speed of sound in blood in the region of a cardiac support system, the system comprising: an ultrasonic sensor, which is arranged in or on the support system, at least one sound reflector, the at least one sound reflector located in the field of vision of the ultrasonic sensor and at a defined distance to at least the ultrasonic sensor or a further sound reflector.

24. The system of claim 23, wherein the at least two sound reflectors are each located at different distances from the ultrasonic sensor.

25. The system of claim 23, wherein the at least one sound reflector is embedded into an embedding material.

26. The system of claim 23 comprising an analysis unit configured to store a pulse time of flight-based analysis algorithm.

27. The system of claim 23 comprising an analysis unit configured to store an FMCW-based analysis algorithm.

28. The method of claim 23, wherein the at least one sound reflector is configured to project sound at least partially into a flow path of the blood formed by an inlet cannula.

29. The method of claim 28, wherein the at least one sound reflector comprises the inlet cannula.

30. A cardiac support system comprising: an ultrasonic sensor, which is arranged in or on the support system, at least one sound reflector, the at least one sound reflector located in the field of vision of the ultrasonic sensor and at a defined distance to at least the ultrasonic sensor or a further sound reflector.

31. The method of claim 30, wherein the at least one sound reflector is configured to project sound at least partially into a flow path of the blood formed by an inlet cannula.

32. The method of claim 31, wherein the at least one sound reflector comprises the inlet cannula.

Description

[0037] The solution presented herein as well as its technical environment are explained below in more detail based on the figures. It is important to note that the invention is not limited by the shown exemplary embodiments. In particular, unless explicitly stated otherwise, it is also possible to extract partial aspects of the facts explained in the figures, and to combine said partial aspects with other components and/or findings from other figures and/or the present description. The following figures show schematically:

[0038] FIG. 1 a sequence of a method presented here in a standard operating sequence,

[0039] FIG. 2a a detailed view of an implantable vascular support system,

[0040] FIG. 2b a detailed view of a further implantable vascular support system,

[0041] FIG. 3 an emission characteristic of an ultrasonic element,

[0042] FIG. 4 an illustration of a system presented here,

[0043] FIG. 5 an illustration of a pulse time of flight-based approach that can be used here,

[0044] FIG. 6 an illustration of an FMCW-based approach that can be used here,

[0045] FIG. 7 example graphs of real parts of impedances,

[0046] FIG. 8a a detailed view of a system presented here, and

[0047] FIG. 8b a detailed view of a further system presented here.

[0048] FIG. 1 shows a schematic representation of a sequence of a method presented here in a standard operating sequence. The illustrated sequence of the method steps a), b), c) and d) with the blocks 110, 120, 130 and 140 is only exemplary. In block 110, an ultrasonic signal is transmitted with an ultrasonic sensor. In block 120, the ultrasonic signal is reflected on at least one sound reflector, which is arranged in the field of vision of the ultrasonic sensor and at a defined distance from the ultrasonic sensor. In block 130, the reflected ultrasonic signal is received. In block 140, the speed of sound is determined in the fluid using the reflected ultrasonic signal.

[0049] In particular, the method steps a), b), and c) can also be executed at least partially or simultaneously in parallel.

[0050] FIG. 2a schematically shows a detailed view of an implantable vascular support system 2. FIG. 2b shows a schematic representation of a detailed view of a further implantable vascular support system 2. FIGS. 2a and 2b are explained jointly below. The reference symbols are used uniformly.

[0051] The method presented here can in principle be integrated into all designs of cardiac support systems. By way of example, FIG. 2a shows the integration into a left ventricular microaxial pump in the aortic valve position, and FIG. 2b shows the integration into an apically positioned radial support system 2.

[0052] The flow direction of the fluid 1 is represented in FIGS. 2a and 2b by arrows. In each case, an ultrasonic sensor 4 is specified, which is arranged in or on the support system 2. The ultrasonic sensors 4 are designed as an ultrasonic transducer in FIGS. 2a and 2b by way of example. In addition, two circumferential sound reflectors 5 are specified along an inner circumference of a flow channel of the support system 2, which are arranged in the field of vision 6 of the ultrasonic sensor 4 and each at a defined distance 7 to the ultrasonic sensor 4. In particular in the embodiment according to FIG. 2a, the flow channel can be formed in the interior of a(n) (inlet) cannula (not shown here) of the support system 2.

[0053] The detailed view according to FIG. 2a shows a tip of a support system 2 with a microaxial pump (not shown here); said tip accommodating the ultrasonic sensor 4. A flow conductive body 10 is in this case by way of example placed directly in front of the ultrasonic sensor 4. Said flow conductive body 10 is not spaced at a distance from the ultrasonic sensor 4 and is permeable for ultrasonic signals. The fluid 1 in this case flows in the direction of the pump. The tip of the support system 2 shown in the detailed view according to FIG. 2a can in a preferred arrangement protrude into a ventricle (not shown here) of a heart with the end shown herein on the left, wherein the pump can be arranged at least partially in the aorta (not shown here). In this arrangement, the support system thus penetrates an aortic valve (not shown here).

[0054] The detailed view according to FIG. 2b relates to a support system 2, which is also referred to as an apical radial pump. The support system 2 comprises a flow machine 11 (a pump in this case), which expels the fluid 1 as shown in radial direction.

[0055] In both exemplary pump variants, the ultrasonic sensor 4, in particular an ultrasonic element of the ultrasonic sensor 4, is usually placed such that the angle to the flow is α=0° (zero degrees); a best possible Doppler shift can therefore be realized.

[0056] FIG. 3 shows a schematic representation of an emission characteristic 12 of an ultrasonic element (not shown here). The emission characteristic 12 of an ultrasonic sensor or an ultrasonic element of the ultrasonic sensor is generally lobe-shaped with a main beam direction straight ahead. This is shown in FIG. 3 as an example for a circular disk ultrasonic transducer with a diameter of 3 mm at f.sub.0=4 MHz. In other words, FIG. 3 illustrates the field of vision 6 of the ultrasonic sensor (not shown here). A field of vision width 13 can be measured along the ordinate (y-axis) and a field of vision length 14 can be measured along the abscissa (x-axis).

[0057] FIG. 4 shows a schematic illustration of a system presented herein. The system comprises an ultrasonic sensor 4 and two sound reflectors 5, which are arranged at a different (defined) distance 7 to the ultrasonic sensor 4. The reflectors 5 project into the fluid 1 by way of example.

[0058] Each boundary layer between two acoustic impedances has a reflection factor at which a part of the sound energy is reflected according to the parameter F.

[00002]$\Gamma =\frac{Z_{w2}-Z_{w1}}{Z_{w2}+Z_{w1}}.Math.\Gamma .Math.\le 1$

[0059] In this case, Z.sub.w1, is the wave impedance before the step point and Z.sub.w2 is the wave impedance after the step point.

[0060] The slightly different acoustic impedance of red blood cells and blood serum, for example, provides the reflected signal, which is usually used to calculate the Doppler frequency shift, from which the flow speed of the blood can be determined.

[0061] A(n) (additional) reflector proposed here should preferably have the highest possible reflection factor, which can be achieved in particular by an impedance mismatch with the blood, i.e., the acoustic impedance of the reflector should differ as clearly as possible from the blood, for example by the reflector being made of an air-filled cavity or a metal.

[0062] The method with only one reflector 5 can be faulty as soon as more than one unknown medium is present between the ultrasonic sensor 4 and the reflector 5. For example, the acoustic impedance (formula symbol: Z.sub.w1) and thus the speed of sound (formula symbol: C.sub.1) of the adjustment layers 15 could change over the years due to water diffusion, or deposits 16 of cell layers (with their own acoustic impedance Z.sub.w2 and speed of sound C.sub.2) could occur on the ultrasonic sensor 4, thus creating an additional material layer of unknown thickness and/or unknown speed of sound, as shown in greater detail in FIG. 4. In this context, the different speeds of sound of the different media are shown in FIG. 4 by way of example, namely the speed of sound C.sub.1 of the adjustment layers 15, the speed of sound C.sub.2 of the deposits 16 and the speed of sound C.sub.3 of the fluid 1 (here: blood).

[0063] FIG. 5 shows a schematic illustration of a pulse time of flight-based approach usable herein. In order to explain the illustration according to FIG. 5 and/or the pulse time of flight-based approach, reference is also made to the illustration of the system according to FIG. 4.

[0064] In addition to the ultrasonic power reflected continuously by each scatter particle of fluid 1 (here: blood; in particular at the respective boundary from blood serum to blood cells), there are clear echoes at the reflectors 5, which can be identified in the received amplitude-time data. In addition, the impulse time of flight from the ultrasonic sensor 4 to the reflector 5 and back to the ultrasonic sensor 4 can be calculated. Since the mechanical design of the (heart) support system 2 and thus the (defined) distance 7 between the ultrasonic sensor 4 and reflector 5 is known, the desired speed of sound c is determined with the formula

[00003]$c=\frac{2s}{t}$

where s is the known (defined) distance 7 between the ultrasonic sensor 4 and reflector 5 and t is the measured signal time of flight.

[0065] When using two reflectors 5 with different distances 7, as shown in FIG. 4, the time of flight t.sub.R1 of the impulse scattered on the first reflector 5 is therefore

[00004]$t_{R1}=2\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}\right]$

[0066] And the time of flight t.sub.R2 of the pulse scattered on the second reflector 5 is

[00005]$t_{R2}=2\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}+\frac{s_4}{c_3}\right]$

[0067] where s.sub.1 is the thickness of the adjustment layers 15, s.sub.2 is the thickness of the deposits 16, s.sub.3 is the distance between deposits 16 and the first (left) reflector 5 and s.sub.4 is the distance between the first (left) reflector 5 and the second (right) reflector 5, and where c.sub.1 is the speed of sound in the adjustment layers 15, c.sub.2 is the speed of sound in the deposits 16, and c.sub.3 is the speed of sound in the fluid 1 (here: blood).

[0068] Since the adjustment layers 15 with the speed of sound c.sub.1 and the deposits 16 with the speed of sound c.sub.2 act equally on both impulses, the difference of the signal times of flight t.sub.R2−t.sub.R1 only contains components in the sought (fluid) range or in the (fluid) range relevant here with the (sought) speed of sound c.sub.3:

[00006]$t_{R2}-t_{R1}=2\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}+\frac{s_4}{c_3}\right]-2\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}\right]=2\frac{s_4}{c_3}$

[0069] Since the distance s.sub.4 of the two reflectors 5 to one another is known, the speed of sound c.sub.3 can be determined independent of the influence of additional layers between the ultrasonic sensor 4 and the reflector 5.

[0070] One possibility for determining the times of flight t.sub.R1 and t.sub.R2 or t.sub.R1−t.sub.R2 is the calculation of the cross-correlation 17 of the transmission pulse 3 (pulse of the transmitted ultrasonic signal 3) to the receiving pulses 8 (pulses of the received and reflected ultrasonic signals 8) reflected on the ultrasonic reflectors 5 and delayed by the times of flight t.sub.R1 or t.sub.R2. The time-discrete cross correlation 17 can be calculated as follows for an energy signal:

[00007]$\text{?}=\text{?}=\underset{\text{?}}{\overset{\text{?}}{.Math.}}\text{?}$$\text{?}\text{indicates text missing or illegible when filed}$

where R.sub.xy [n] is the discrete cross-correlation at time n, and the operator “star” as an acronym for the cross-correlation, x* [m] is the conjugated complex transmission signal over all time shifts m, and y[m+n] is the receiving signal at time n over all time shifts m.

[0071] The illustration according to FIG. 5 shows an example of the result of this calculation. FIG. 5 shows the pulse of the emitted ultrasonic signal 3, the pulses of the received reflected ultrasonic signals 8 and the (time-discrete) cross correlation 17 over time 18. The time interval t.sub.R1−t.sub.R2 can be determined from the distance between, e.g., the two tips (peaks) in the cross-correlation signal 17—after reverse-recalculating the discrete time steps.

[0072] FIG. 6 shows a schematic illustration of an FMCW-based approach usable here. In order to explain the illustration according to FIG. 6 or the FMCW-based approach, reference is also made to the illustration of the system according to FIG. 4.

[0073] The (ultra)sound reflectors 5 represent the dominant targets in the emission range of the ultrasonic sensor 4, in particular due to their high reflection factor. Their beat frequencies can therefore be clearly detected in the calculated spectrum. Since the mechanical design of the (heart) support system and thus the distance between the ultrasonic sensor 4 and the reflector 5 (formula symbol x) is known, the desired speed of sound c is determined by the formula

[00008]$c=2\star s_x\star \frac{\left(\frac{\mathrm{bw}}{T}\right)}{f_{\mathrm{beat},x}}$

where s.sub.x is the known distance between the ultrasonic sensor and reflector x, bw/T is the slope of the frequency ramp, and f.sub.beat,x is the resulting beat frequency in the base band. In particular, since the reflectors 5 are installed in a fixed location, the resulting beat frequency is only influenced by their distance to the ultrasonic sensor 4 and the corresponding time of flight of the frequency ramp in the fluid (here: blood), and in particular contains no speed-dependent portion.

[0074] When using two reflectors 5 with different distances 7, as shown in FIG. 4, the beat frequency f.sub.beat,R1 of the frequency ramp reflected at the first reflector is therefore

[00009]$f_{\mathrm{beat},R1}=2\left(\frac{\mathrm{bw}}{T}\right)\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}\right]$

and the beat frequency f.sub.beat,R2 of the frequency reflected on the second reflector is

[00010]$f_{\mathrm{beat},R2}=2\left(\frac{\mathrm{bw}}{T}\right)\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}+\frac{s_4}{c_3}\right]$

where s.sub.1 is the thickness of the adjustment layers 15, s.sub.2 is the thickness of the deposits 16, s.sub.3 is the distance between deposits 16 and the first (left) reflector 5 and s.sub.4 is the distance between the first (left) reflector 5 and the second (right) reflector 5, and where c.sub.1 is the speed of sound in the adjustment layers 15, c.sub.2 is the speed of sound in the deposits 16, and c.sub.3 is the speed of sound in the fluid 1 (here: blood).

[0075] Since the adjustment layers 15 with the speed of sound c.sub.1 and the deposits 16 with the speed of sound c.sub.2 act equally on both frequency ramps, the difference of the beat frequencies f.sub.beat,R2 f.sub.beat,R1 only contains components in the searched (fluid) range or in the (fluid) range relevant here with the (searched) speed of sound c.sub.3:

[00011]$f_{\mathrm{beat},R2}-f_{\mathrm{beat},R1}=2\left(\frac{\mathrm{bw}}{T}\right)\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}+\frac{s_4}{c_3}\right]-2\left(\frac{\mathrm{bw}}{T}\right)\left[\frac{s_1}{c_1}+\frac{s_2}{c_2}+\frac{s_3}{c_3}\right]=2\left(\frac{\mathrm{bw}}{T}\right)\frac{s_4}{c_3}$

[0076] Since the distance s.sub.4 of the two reflectors 5 to one another is known, the speed of sound c.sub.3 can be determined irrespective of the influence of additional layers between the ultrasonic sensor 4 and the reflector 5.

[0077] To determine the beat frequencies, the ultrasonic frequency f.sub.0 is influenced by frequency modulation as an example. Without limitation, sine-wave-shaped, saw-tooth-shaped, triangular or rectangular modulation types can be used. It is particularly preferred that the ultrasonic sensor or the ultrasonic element of the sensor provide a broadband resonance and that the ramp time of flight (formula symbol: T) is much greater than the time of flight of the frequency ramps to the ultrasonic sensor 4 (ultrasound transducer) to the (ultra)-sound reflectors 5 and back again. The echoes of the successively emitted, modulated ultrasonic frequency reflected at the reflectors 5 are overlaid with the instantaneous transmission frequency ramp. The base band signal generated in this way contains the beat frequencies to be determined. These are converted by the transformation into the frequency range, e.g., by discrete Fourier transformation (DFT) or fast Fourier transformation (FFT).

[0078] The illustration according to FIG. 6 shows a possible realization of the previously described FMCW-based approach by means of a sawtooth modulation. The upper diagram of FIG. 6 shows the graph of the frequency 19 versus time 18. It can be seen that both the ultrasonic signal 3 (transmission signal) emitted by the ultrasonic sensor and the reflected ultrasonic signals 8 (receiving signals) received by the ultrasonic sensor (three here as an example) are shaped in the manner of a sawtooth. In this case, three receiving signals 8 shifted relative to the transmission signal 3 and to one another are applied as examples, which would for example be the case if three ultrasonic reflectors arranged at different distances to the ultrasonic sensor were used.

[0079] The FMCW approach regularly works with a periodic frequency modulation, in this case periodic sawtooth modulation, which should be as time-linear as possible to ensure the best-possible accuracy of the measurement. The modulation is usually performed cyclically. Such a cycle from the lowest to the highest frequency is also referred to as a signal burst. The duration of a corresponding cycle is shown in the upper diagram of FIG. 6 as a so-called chirp duration 22. In addition, a usable chirp duration 23 is marked.

[0080] The ultrasonic sensor in this case sends an example of a linear frequency-modulated signal with a sawtooth-shaped change of the transmission frequency 3. The same signal is received by the ultrasonic sensor after a reflection on one of the ultrasonic reflectors. The received signal 8 differs in the time, wherein the time difference 21 between the frequency shifts is generally proportional to the distance of the reflective ultrasonic reflector from the ultrasonic sensor. At the same time (assuming a linear frequency change), the difference frequency 20 between the transmission signal 3 and the receiving signal 8 is the same at any point in time and is thus also a measure for the distance to the reflective ultrasound reflector. This frequency difference can be evaluated in particular in the frequency range.

[0081] The frequency plots of the upper diagram in FIG. 6 are in this example used to generate a frequency spectrum 25 by overlaying/multiplying with the instantaneous transmission signal and by means of a subsequent fast Fourier transformation 24, wherein said frequency spectrum 25 carries the difference frequencies 20 in addition to the background noise 26. In a simplified manner, the receiving signal is multiplied with the instantaneous transmission signal, followed by a Fourier transformation of the base band time signal, from which the difference frequencies 20 result, which are also referred to herein as beat frequencies. The minimum range resolution of FMCW systems is

[00012]$\Delta r=\frac{c}{2\star \mathrm{bw}}=s_4$

is defined. Accordingly, when two ultrasonic reflectors 5 are, e.g., placed at a distance of Δr=s.sub.4=6 mm to each other, and at a(n) (expected) speed of sound in blood c of about 1540 m/s (used to determine the approximately required or particularly advantageous bandwidth), it is possible to work with a bandwidth bw≈128 kHz≤150 kHz.

[0082] But a significantly higher range accuracy can be achieved by the additional use of techniques, such as the so-called zero padding (concatenating or padding of zeros) or high-performance frequency estimation methods. This can contribute to a significantly more precise determination of the speed of sound c in the blood. The achievable accuracy depends in particular on the frequency estimation method and/or the signal-to-noise ratio.

[0083] The particularly advantageous linearity can in particular be achieved over the desired frequency band when using piezo elements (as ultrasonic elements), preferably when the quality of the resonance (wide-band resonance) is reduced by backing (amplification). The illustration according to FIG. 7 shows example plots of real components 27 of the impedances of 8 MHz piezo elements versus the stimulation frequency 28. In the case shown, a frequency ramp with the example bandwidth bw=150 kHz could be placed in the frequency band 29 highlighted in gray.

[0084] FIG. 8a shows a schematic of a detailed view of a system presented here. FIG. 8b shows a schematic of a detailed view of a further system presented here. FIGS. 8a and 8b are explained jointly below. The reference symbols are used uniformly.

[0085] To achieve the best possible reflection, the surface of the reflector should be parallel to the incident ultrasonic wavefront. Since non-planar surfaces such as superimposed reflectors can lead to turbulence in the flow (disadvantageous for Doppler ultrasonic measurement), to the formation of thrombi, and to additional blood damage (hemolysis) due to shear forces, it is expedient to embed the reflectors 5 into an embedding material 9, as illustrated by way of example in FIGS. 8a and 8b. The embedding material 9 is used here as an example for providing a smoother surface or a surface without corners and/or edges in comparison to the reflector surface. It is particularly preferred to embed the at least one reflector 5 into a planar surface, in particular by means of the embedding material 9. The embedding material 9 should as much as possible have the same acoustic impedance as the fluid 1 (here: blood) and be as thin as possible, so that there are no additional reflections or diffractions of the acoustic impulse, unless this additional diffraction is desired. For example, the (or each) reflector 5 with an acoustic impedance C.sub.4 can be embedded into a silicone with an acoustic impedance C.sub.3′, wherein C.sub.3′ is similar to the acoustic impedance C.sub.3 of blood.

[0086] The solution presented here in particular has one or more of the following advantages: [0087] By supplementing at least one ultrasonic reflector in the emission range of the ultrasonic system, the speed of sound can be determined from the resulting pulse time of flight and/or the ramp time of flight from the reflector. [0088] The known speed of sound increases the measurement accuracy of the flow measurement. [0089] The speed of sound depends on the composition of the blood and can in this case be determined and used directly.

[0090] The FMCW approach does not require a very precise time difference to be measured; an equivalent frequency difference can be determined instead, which significantly reduces the technical effort.