WO2024110210A1 - Procédé de détermination des phases d'un signal artériel et dispositif de mesure conçu à cet effet - Google Patents

Procédé de détermination des phases d'un signal artériel et dispositif de mesure conçu à cet effet Download PDF

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Publication number
WO2024110210A1
WO2024110210A1 PCT/EP2023/081386 EP2023081386W WO2024110210A1 WO 2024110210 A1 WO2024110210 A1 WO 2024110210A1 EP 2023081386 W EP2023081386 W EP 2023081386W WO 2024110210 A1 WO2024110210 A1 WO 2024110210A1
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WIPO (PCT)
Prior art keywords
signal
cuff
pressure
sensor
measuring device
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PCT/EP2023/081386
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German (de)
English (en)
Inventor
Bonev SLAVTCHO
Original Assignee
Bopuls Ug
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Publication of WO2024110210A1 publication Critical patent/WO2024110210A1/fr

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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/02—Detecting, measuring or recording pulse, heart rate, blood pressure or blood flow; Combined pulse/heart-rate/blood pressure determination; Evaluating a cardiovascular condition not otherwise provided for, e.g. using combinations of techniques provided for in this group with electrocardiography or electroauscultation; Heart catheters for measuring blood pressure
    • A61B5/021—Measuring pressure in heart or blood vessels
    • A61B5/022—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/02—Detecting, measuring or recording pulse, heart rate, blood pressure or blood flow; Combined pulse/heart-rate/blood pressure determination; Evaluating a cardiovascular condition not otherwise provided for, e.g. using combinations of techniques provided for in this group with electrocardiography or electroauscultation; Heart catheters for measuring blood pressure
    • A61B5/021—Measuring pressure in heart or blood vessels
    • A61B5/022—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers
    • A61B5/02208—Measuring pressure in heart or blood vessels by applying pressure to close blood vessels, e.g. against the skin; Ophthalmodynamometers using the Korotkoff method
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B5/00—Measuring for diagnostic purposes; Identification of persons
    • A61B5/72—Signal processing specially adapted for physiological signals or for diagnostic purposes
    • A61B5/7225—Details of analog processing, e.g. isolation amplifier, gain or sensitivity adjustment, filtering, baseline or drift compensation

Definitions

  • the invention relates to a method for determining the phases of an arterial dynamic process and a measuring device designed for this purpose.
  • the measuring device is designed in particular as a blood pressure measuring device.
  • the gold standard for measuring blood pressure is the Korotkoff method (auscultatory method, see CE Grim: Auscultatory BP: still the gold standard, Journal of the American Society of Hypertension 2016; 10 (3) , 191-193).
  • the extremity for example the upper arm, is subjected to pressure using a cuff and systolic and diastolic pressure are determined by auscultation when the pressure is released.
  • the method used for measuring blood pressure The so-called Korotkof f sounds that occur can be divided into five phases:
  • the knocking sound is accompanied by an additional murmuring sound.
  • the blood pressure at the beginning of the first phase is called systolic blood pressure and the blood pressure at the transition from the fourth to the fifth phase is called diastolic blood pressure.
  • the trained physician can also use auscultation to determine the presence of certain diseases based on the transitions between the phases.
  • the systolic blood pressure which is measured when the left ventricle is emptying, is determined by the first occurrence of vascular sounds while the cuff pressure is being released.
  • the artery in the upper arm is still compressed by the inflated cuff.
  • the pressure from the left half of the heart is higher than the cuff pressure, so that the blood begins to flow through the artery again.
  • the vibration of the artery now increases.
  • the diastolic blood pressure has been reached. This is the blood pressure after the ventricle has relaxed and fills with blood again.
  • the artery in the upper arm is now completely permeable again and the cuff pressure has been released.
  • the oscillometric measuring devices determine the maximum vascular oscillation and calculate the systolic and diastolic blood pressure using a mathematical formula.
  • blood pressure is calculated from the vibrations of the blood vessel wall (oscillometric measurement).
  • the systolic and diastolic blood pressure values are derived from the maximum vibration of the vessel walls (also called “mean arterial pressure”).
  • the exact algorithms for deriving the blood pressure values are usually not disclosed by the manufacturers.
  • the invention is based on the object of reducing the above-mentioned disadvantages of the prior art.
  • it is an object of the invention to determine at least one phase transition of an arterial signal as precisely as possible in a simple manner.
  • the phase transition is determined in connection with a blood pressure measurement.
  • the object of the invention is already achieved by a measuring device and by a method for determining the phases of an arterial signal according to one of the independent claims.
  • the invention relates to a measuring device for determining the phases of an arterial signal.
  • the measuring device is thus designed in such a way that it determines at least one phase transition of the arterial signal.
  • it is not necessary for the phase transition to be displayed to the user in any way. Rather, the phase transition can be determined in order to carry out a calculation, in particular a calculation of the associated blood pressure value.
  • the measuring device is therefore specially designed as a blood pressure monitor.
  • the measuring device is designed for the automated determination of at least one blood pressure value, in particular SYS and/or DIAS, by a combination of an oscillometric and an auscultatory measuring method.
  • the measuring device comprises an inflatable cuff that can be placed around an extremity.
  • the measuring device can be designed as an upper arm or forearm measuring device.
  • the cuff can be inflated by means of a pump until the arterial blood flow stops and then By releasing the pressure from the cuff the arterial signal is measured, in particular to determine the blood pressure.
  • a measurement can be taken during inflation.
  • the pressure can be kept constant over a certain period of time, which extends over several pulse beats.
  • the measuring device has a sensor for detecting the fluid pressure present in the cuff.
  • This sensor can be used in particular to detect blood pressure.
  • This pressure sensor detects in particular the absolute pressure.
  • the sensor can in particular be arranged in the measuring device itself, which is connected to the cuff via a hose.
  • the measuring device comprises a differential pressure sensor for detecting pressure fluctuations between the cuff and the extremity.
  • This additional sensor can be located on the underside of the cuff.
  • the sensor is positioned over the artery as intended.
  • a differential pressure sensor in the sense of the invention is understood to be a sensor which detects dynamic pressure fluctuations between the extremity and the cuff.
  • the sensor is therefore not designed as an absolute pressure sensor, but rather responds to changes in the pressure between the extremity and the cuff.
  • the arterial pulse causes a pressure fluctuation which can be detected by the additional sensor.
  • a differential pressure sensor is understood to be any sensor that measures a physical quantity which is related to a change in pressure.
  • the sensor can therefore also measure a force which is essentially proportional to the pressure or a strain which is caused by a pressure fluctuation.
  • the pressure fluctuations cause pulsating tensions in the material and/or surface deformations or vibrations between the skin and the cuff. These tensions/vibrations can be a measure of the differential pressure in the sense of the invention.
  • the additional sensor can in particular also be defined as a sensor for detecting the transmural pressure.
  • the sensor is preferably designed as a surface sensor, in particular as a thin-film sensor. It can in particular be a piezoelectric sensor. Thin-film sensors that function according to the piezoelectric principle comprise a piezoelectric layer that is surrounded by two conductive layers. The conductive layers can in particular be a plastic with a metal vapor-deposited thereon.
  • the senor can be characterized as a vibration sensor in a lower frequency range. Furthermore, according to one embodiment, the sensor can be characterized as a surface microphone in an upper frequency range.
  • the senor can be characterized as a strain sensor, which measures the pressure fluctuations by detecting the strain of the material induced by pressure fluctuations. It has been found that such a sensor can be used to generate an additional measurement signal which makes it surprisingly easy to determine the phase transitions of the arterial signal more precisely.
  • the transition to the first phase, third phase and/or fifth phase can be recorded more precisely (systolic pressure, mean arterial pressure, diastolic pressure).
  • the measuring device is used as a blood pressure measuring device, the phase transition can be easily defined more precisely and systolic and/or diastolic blood pressure can be determined more precisely.
  • the measurement signal from the sensor for recording the cuff pressure is used to determine the blood pressure.
  • the sensor can record the arterial signal in a frequency range that includes both the auscultatory and the oscillometric components. This allows the arterial signal, i.e. the arterial dynamic process, to be represented more accurately.
  • Oscillometric and auscultatory signals are derived from the measured arterial signal and evaluated pulse by pulse using both methods.
  • the pulse-wise curves are suitable for the automated determination of all five phases of the arterial dynamic process.
  • the differential pressure sensor preferably has a measuring range from 0.05 Hz and/or up to 500 Hz, in particular from 0.2 Hz and/or up to 200 Hz. In particular, the differential pressure sensor has a measuring range from 0.5 Hz and/or up to 100 Hz.
  • the senor detects vibrations in a low frequency range, in particular from 0.1 Hz.
  • the measuring device comprises a bandpass filter which divides the signal of the differential pressure sensor into a low frequency and a high frequency component.
  • the cut-off frequency can be between 10 and 30 Hz, in particular between 15 Hz and 25 Hz.
  • the signal of the differential pressure sensor is divided into two signal components, namely a low-frequency component and a high-frequency component.
  • the low frequency component can serve as a vibration signal, particularly as an oscillometric signal.
  • the high frequency component can serve as an auscultatory signal. In the higher frequency range, the sensor therefore works as a surface microphone.
  • the signal from the differential pressure sensor in particular the oscillometric signal component, has a time offset compared to the pressure signal from the cuff. This time offset can be determined in particular at an inflection point of the respective pulse signal.
  • the measuring device therefore comprises a device for calculating the phases of the arterial signal, which determine the beginning and/or the end at least one phase taking into account a time offset between the signal from the fluid pressure sensor and the differential pressure sensor.
  • local maxima, minima and/or inflection points of the signal curve can be used to determine phase transitions of the arterial signal.
  • the evaluation can be carried out in particular on a pulse-by-pulse basis.
  • a feature of the signal curve can be extracted (e.g. the fundamental wave amplitude, distortion factor, etc.) and plotted over time.
  • the differential pressure sensor preferably extends over less than 50%, particularly preferably less than 20%, very preferably less than 10% of the length of the cuff. In particular, it extends over 3 to 20%, preferably 5 to 15% of the length of the cuff (in the longitudinal direction).
  • the differential pressure sensor has an active measuring area of 50-1000 mm 2 , preferably 200-700 mm 2 , particularly preferably 300-500 mm 2 .
  • the length of the cuff which is particularly designed as an upper arm or forearm cuff, is understood to be the area along the extremity in which the cuff can be inflated.
  • the length of the sensor is understood to be the length of the active measuring area.
  • the cuff can in particular have a length of 10 to 20 cm.
  • the cuff is typically designed for an upper arm circumference of 20 to 35 cm.
  • the sensor which is designed as a thin-film sensor, can in particular have a Length of 5 to 20 mm, preferably 8 to 15 mm and/or a width of 10 to 80 mm, preferably 25 to 40 mm.
  • the differential pressure sensor delivers a narrower signal during a pulse than the cuff pressure sensor.
  • the effect of the arterial pulse is locally limited. Therefore, the pulse contour can be measured more precisely with the additional sensor than by measuring the cuff pressure.
  • the measuring device comprises a proportional valve for releasing the pressure from the cuff.
  • the measuring device is thus designed in such a way that the pressure drops essentially linearly during the measuring phase. This also improves the measuring accuracy and facilitates further processing of the measuring signal.
  • the proportional valve can be designed to be electronically controlled via a pressure sensor or, depending on its design, to lead to an almost linear progression when releasing pressure.
  • the invention further relates to a method for determining the phases of an arterial signal.
  • the method is carried out in particular with the measuring device described above.
  • the pressure in a cuff extending around a vessel is measured.
  • the differential pressure between the cuff and the vessel is measured by means of a further sensor, in particular a differential pressure sensor.
  • a physical signal such as force, voltage, etc., which depends on the differential pressure, can be used to determine the differential pressure.
  • the sensor can be designed as a vibration sensor and/or surface microphone and have the measuring ranges described above.
  • the sensor signal is preferably divided into a low frequency and a high frequency component, in particular into an oscillometric and an auscultatory component.
  • the cutoff frequency between low frequency and high frequency components is preferably between 10 and 30 Hz, in particular between 15 and 25 Hz.
  • Phases of the arterial signal can be determined, in particular by taking into account an offset between the cuff pressure and the differential pressure between cuff and vessel.
  • the time offset between the cuff pressure and the differential pressure can be used to determine the phases of the arterial signal.
  • the method is preferably carried out during the depressurization of the cuff, wherein the pressure is preferably reduced substantially linearly.
  • an oscillometric signal, an auscultatory signal and a differential pressure signal can be generated.
  • Phases of the arterial signal can be determined in particular via turning points and/or local minima or maxima of the signals.
  • Turning points in signal curves are also referred to as trend changes. This is the point in the measurement signal at which the signal changes its curvature behavior, i.e. either changes direction from a right turn to a left turn or vice versa.
  • Turning points and/or local minima or maxima of the signal are used in particular in a pulse-width evaluation.
  • the beginning of the first phase of the arterial signal is defined by the fact that the blood begins to flow into the artery and a sharp tapping sound begins .
  • the distortion factor of the oscillometric signal has a local maximum.
  • the distortion factor is the measure of distortion of a sinusoidal signal.
  • the distortion factor indicates the extent to which the harmonics overlap the sinusoidal fundamental oscillation and can be determined mathematically.
  • the second phase when the tapping sound of the first phase is accompanied by an additional murmuring sound, causes a local maximum of the auscultatory signal and/or an inflection point of the difference signal.
  • the third phase in which pure knocking sounds can be heard, is decisive for the mean arterial pressure.
  • the third phase there is a local maximum of the oscillometric signal and/or a local minimum of the distortion factor of the oscillometric signal.
  • the fourth phase in which the tones become muffled and the typical knocking disappears, can be determined by a local maximum of the auscultatory signal and/or a local maximum of the difference signal.
  • the transition from the fourth phase to the fifth phase is decisive for the diastolic pressure.
  • the noises disappear completely. This can be determined by an inflection point of the oscillometric signal, a local maximum of the distortion factor of the oscillometric signal, an inflection point of the auscultatory signal and/or an inflection point of the difference signal.
  • This phase determination can be used to narrow down the time of measurement of the systolic and/or diastolic pressure and/or the mean arterial pressure, particularly in the case of blood pressure measurement.
  • the invention further relates to a medical measuring device which is designed to carry out the method described above.
  • the measuring device can in particular be designed as a blood pressure measuring device.
  • the measuring device has a processor and a non-volatile memory.
  • a program is stored in the non-volatile memory, which is designed to carry out the method steps according to the method described above.
  • Fig. 1 is a schematic representation of a measuring device according to the invention.
  • Fig. 2 is a block diagram of the structure of the measuring device.
  • Fig. 3 shows the pressure curve in the cuff during a measurement.
  • Fig. 4 shows the signal curve of the cuff pressure sensor and the differential pressure sensor.
  • Fig. 4a is an enlarged representation of the signals.
  • Fig. 4b is a spectral representation of the signals.
  • Fig. 5a and 5b are graphs of the low frequency and high frequency components of the differential pressure sensor.
  • Fig. 6 compares the phases of the arterial signal with different signal curves.
  • Fig. 7 shows the waveform of the oscillometric signal during a measurement.
  • Fig. 8 shows the signal curve of the auscultatory signal during a measurement.
  • Fig. 9 shows the time difference curve of the signal of the differential pressure sensor and the cuff pressure sensor during a measurement.
  • Fig. 10 is a decision table that explains the signal evaluation.
  • Fig. 11a and Fig. 11b are schematic representations of the differential pressure sensor.
  • Fig. 12 is an equivalent circuit diagram of the differential pressure sensor.
  • Fig. 1 is a schematic representation of an embodiment of a measuring device 10 according to the invention during its use. The illustration on the right is a section.
  • the measuring device 10 comprises an inflatable cuff 11 which is placed around an extremity, in this illustration around the user's arm 20.
  • the measuring device 10 can be designed as an upper arm or lower arm measuring device for measuring blood pressure.
  • the cuff pressure is measured via the cuff pressure sensor (see Fig. 2).
  • the cuff pressure sensor (16 in Fig. 2) can be arranged in an external unit that is connected to the cuff 11 via the line (e.g. hose) 12.
  • the cuff pressure sensor measures the pneumatic pressure inside the cuff 11.
  • the arm 20 includes the bone 22 and the artery 21.
  • the cuff 11 is designed such that a differential pressure sensor 15 is arranged adjacent to the artery 21 on the underside of the cuff 11.
  • the differential pressure sensor 15 is designed as a thin-film sensor with a piezo layer.
  • the cuff 11 extends over a length of 1 m over the arm 20 .
  • the sensor 15 extends over a length l s .
  • l s is smaller than 1 M - in particular, l s is less than 20% of 1 M .
  • Fig. 2 is a schematic representation of the components of the cuff inflation device 11 .
  • the measuring device comprises a pneumatic pump 13 , by means of which the cuff 11 is inflated via the line 12 .
  • the pressure sensor 16 serves to limit the maximum pressure.
  • the measuring device is set up in such a way that a stored maximum pressure is initially generated. If there is still a pulse at this pressure, i.e. the artery is not compressed, the pressure is increased again.
  • the pump 13 is controlled via the microcontroller 17 .
  • the pressure sensor 16 also serves as a cuff pressure sensor 16 , i.e. it measures the pressure in the cuff during the measuring phase.
  • valve 18 is arranged in the line 12. To release air from the cuff, the pump 13 is switched off and the valve 18 is opened. The air now escapes via the line 12 while the measurement is being carried out.
  • the valve 18 is also controlled via the microcontroller 17.
  • the microcontroller 17 can control the valve 18 in such a way that the pressure in the cuff decreases almost linearly when deflated.
  • Fig. 3 shows the pressure curve during use of the measuring device.
  • the cuff is inflated and the pressure increases to a programmed target value.
  • the arterial signal is measured during the third phase, i.e. while the air is being released from the cuff.
  • the sensor signals shown in Fig. 4 are generated.
  • the x-axis shows the time in seconds.
  • the y-axis for the differential pressure sensor is a standardized amplitude of the measurement signal, while for the cuff pressure sensor it is the absolute pressure in mmHg.
  • the cuff pressure signal pulsates with the pulse and decreases continuously over the measurement due to the pressure release.
  • the differential pressure signal on the other hand, which in this embodiment is recorded in a measuring range between 0.1 and 100 Hz, pulsates clearly with the pulse and changes, as will be shown in more detail below, among other things due to the different phases of the arterial signal.
  • the maximum pulse amplitude is significantly higher for the differential pressure sensor than for the cuff pressure sensor. In a measuring interval, this can be 0.5 to 5% of the maximum amplitude for the cuff pressure sensor and 40 to 80% of the maximum amplitude for the differential pressure sensor.
  • the maximum pulse amplitude (at time 15 s) for the cuff pressure signal is about 2% of the measuring range and for the piezo sensor about 60% of the measuring range.
  • the piezo sensor can detect the noise (alternating components > 20 Hz) through direct contact with the skin surface and the large sensor surface.
  • the noise is absorbed by the cuff material and the Cuff air bubble so strongly dampened that the cuff pressure sensor practically no longer measures alternating components > 15 Hz (amplitudes >6 dB above noise level, see spectra of the measurement signals according to Fig. 4b).
  • the differential pressure signal is split into a low-frequency oscillatory signal and a high-frequency auscultatory signal. This is done using a bandpass filter.
  • Fig. 5a shows a typical differential pressure signal during a pulse (thin line) as an example for phase 1 and phase 3 of the arterial signal.
  • the signal from the cuff pressure sensor approximately follows the signal curve of the oscillatory signal, but is delayed by a time period At.
  • At is significantly larger in the third phase than in the first phase.
  • At is determined at an inflection point of the signal. According to another embodiment, however, it is also conceivable to determine At at a local minimum or local maximum.
  • Fig. 5b shows the signal curve of the high-frequency auscultatory signal of the differential pressure sensor. It can be seen that the signal curve in phase 1 also differs from phase 3.
  • phase 3 the envelope of the signal curve is flatter and wider (dashed line).
  • Fig. 6 shows the course of the oscillometric signal and the auscultatory signal compared with the phases of the arterial signal.
  • phase 1 blood begins to flow into the artery and a knocking sound occurs.
  • phase 2 the artery is open and there are weak tapping sounds, superimposed by murmuring sounds.
  • phase 3 the artery is so dilated that pure knocking sounds can be heard again. These are louder than in phase 1.
  • phase 4 the sounds become quieter and the typical knocking sounds disappear and are replaced by fading dull sounds until the sounds stop at the transition to phase 5.
  • the oscillometric signal is formed in this part from the low frequency component in the frequency range from 0.5 to 20 Hz.
  • the normalized signal curve is plotted against the cuff pressure signal curve.
  • both signals show characteristic changes during the phases .
  • the signals both the oscillometric and the auscultatory signal, can be converted into a spectral representation for further processing.
  • a fundamental wave is formed over the measurement time, which is plotted over a standardized amplitude.
  • the signal curve for a pulse wave is thus shown (beat to beat, as also in Fig. 8 and 9).
  • the fundamental wave rises and falls again.
  • the vibrations detected by the differential pressure sensor are at their highest.
  • a distortion factor can be derived from the fundamental wave.
  • the distortion factor is relatively high because the fundamental wave deviates quite strongly from an ideal sine wave.
  • Fig. 8 shows the auscultatory signal of the differential pressure sensor, with time on the x-axis and the relative volume in dB (A) on the y-axis.
  • the auscultatory signal is therefore the signal of the differential pressure sensor, which works as a surface microphone in this area.
  • a characteristic signal curve is also obtained over the five phases of the arterial signal.
  • Fig. 9 shows the signal curve of the time difference ⁇ t between the cuff pressure signal and the differential pressure signal.
  • the x-axis shows the time in seconds and the y-axis shows the time difference in milliseconds ms.
  • a signal curve characteristic of the different phases of the arterial signal is also obtained.
  • Fig. 10 shows a decision table showing the start of the different phases.
  • a measuring device in which the auscultatory signal and/or the time difference signal are dispensed with.
  • phase 1 and the beginning of phase 5 which are relevant for systolic and diastolic pressure, can already be determined via an inflection point of the oscillometric fundamental wave.
  • the accuracy can be increased by taking into account the distortion factor, which forms a local maximum at the beginning of phase 1 and phase 5.
  • the distortion factor which forms a local maximum at the beginning of phase 1 and phase 5.
  • phase 3 which is relevant for the mean arterial pressure, the maximum of the oscillometric fundamental wave is sufficient.
  • the corresponding algorithm can thus determine the blood pressure more accurately by narrowing it down more precisely, especially in phases 1 and 5.
  • the measurement result is less sensitive to atypical signal curves, for example due to diseases.
  • the measuring method according to the invention and the measuring device according to the invention can also be used to detect atypical phase progressions for diagnostic purposes.
  • the measurement method can be used for diagnostic purposes for the following diseases:
  • Diagnosis can be carried out in particular by evaluating the signal curves using artificial intelligence. Increased arterial stiffness precedes arteriosclerosis, which then significantly increases the risk of chronic wounds and amputations due to reduced blood flow. Arterial stiffness can be detected particularly using pulse curve profiles.
  • Endothelial dysfunction can be measured as a disturbance of endothelium-dependent vasodilation. This can be determined in particular on the basis of the rate of increase of the pressure curve.
  • Heart valve insufficiency describes a leaky heart valve, i.e. the heart valve does not close completely. Blood can flow back through the open valve, which reduces the heart's pumping capacity. In advanced stages, this can lead to heart failure. Heart valve insufficiency can be detected based on the time it takes for the pressure to rise.
  • the method according to the invention is particularly advantageous in that the phase of the arterial signal can be assigned to a measured pressure curve with high accuracy.
  • the course of the pulse-related signals can thus serve as a basis for advanced diagnostics, especially vascular health diagnostics.
  • An evaluation of the new signals over longer periods of time using artificial intelligence can serve both to expand the knowledge base about the cardiovascular system and to improve early diagnosis. This can support the timely and correct application of appropriate therapies and the evaluation of their success.
  • Fig. 11a and Fig. 11b are schematic representations of the differential pressure sensor 15.
  • the differential pressure sensor 15 comprises a piezo layer 15a, which is formed as a thin film.
  • the piezo layer 15a is surrounded by two conductive layers 15b.
  • the conductive layers 15b can be formed as metal-deposited plastic layers.
  • the sensor 15 thus forms a capacitor in the equivalent circuit shown in Fig. 12, the capacitance of which changes.
  • the circuit includes the capacitance of the sensor , the capacitance of the line, i.e. the cable , in particular of the coaxial cable leading to the terminals 15c, as well as the internal resistance R ⁇ of the amplifier circuit.
  • An output voltage U is generated as a function of the differential pressure acting on the sensor 15.
  • the invention made it possible to easily and automatically record the phases of an arterial signal with high accuracy.

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  • Health & Medical Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • Vascular Medicine (AREA)
  • Cardiology (AREA)
  • Biomedical Technology (AREA)
  • Heart & Thoracic Surgery (AREA)
  • Physiology (AREA)
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Abstract

L'invention concerne un dispositif de mesure pour déterminer les phases d'un signal artériel, en particulier un tensiomètre. Le dispositif de mesure comprend un brassard gonflable qui peut être placé autour d'un membre, un capteur pour détecter la pression de fluide dans le brassard, et un capteur de pression différentielle pour détecter des fluctuations de pression entre le brassard et le membre.
PCT/EP2023/081386 2022-11-22 2023-11-10 Procédé de détermination des phases d'un signal artériel et dispositif de mesure conçu à cet effet WO2024110210A1 (fr)

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DE102022130899.5 2022-11-22
DE102022130899.5A DE102022130899B4 (de) 2022-11-22 2022-11-22 Verfahren zur Bestimming der Phasen eines Arteriensignals sowie hierfür ausgebildetes Messgerät

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