WO2024110210A1 - Method for determining the phases of an arterial signal and measuring device designed therefor - Google Patents

Abstract

The invention relates to a measuring device for determining the phases of an arterial signal, in particular a blood pressure meter. The measuring device comprises an inflatable cuff that can be placed around an extremity, a sensor for detecting the fluid pressure present in the cuff, and a differential pressure sensor for detecting pressure fluctuations between the cuff and the extremity.

Description

Method for determining the phases of an arterial signal and measuring device designed for this purpose

Description

Field of the invention

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.

Background of the invention

In practice, two methods are commonly used to measure the physiologically important blood pressure values SYS and DIAS noninvasively: auscultatory (according to Korotkoff, gold standard) and oscillometric (according to Riva-Rocci). Both methods use a pressure cuff. The oscillometric measurement method is particularly common in blood pressure monitors that perform automated measurements (see G.

Drzewiecki: Noninvasive Arterial Blood Pressure and Mechanics, The Biomedical Engineering Handbook, 2nd Edition, 2000;

Chapter 71.2) .

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:

As the blood enters the artery, a sharp knocking sound begins, the intensity of which increases during the first phase.

In the second phase, the knocking sound is accompanied by an additional murmuring sound.

In the third phase, pure knocking noises can be heard again, which are now particularly strong and sharp.

In the fourth phase, the tones sound muffled and quieter, the typical knocking quality disappears, until in the fifth phase the noises disappear (see also S. Eckert: 100 years of blood pressure measurement according to Riva-Rocci and Korotkoff, Journal für Hypertonie 2006; 10 (3) , 7-13, www.kup.at/hypertonie).

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. When measuring blood pressure alone, these two values are particularly relevant.

Furthermore, the trained physician can also use auscultation to determine the presence of certain diseases based on the transitions between the phases.

However, the above measurement method, which is considered the gold standard, is complex and requires trained personnel.

The historical development and the measurement problems are comprehensively explained in the following publication:

J. Allen: Characterization of the Korotkoff sounds using joint time-frequency analysis, Physiological Measurement 2004; 25, 107-117 . There are also a number of blood pressure monitors that automatically measure systolic and diastolic blood pressure. In simple designs, the monitor only includes a sensor that measures the cuff pressure. The cuff is inflated until no more pressure fluctuations are measured. A The valve is opened and the pressure is measured using the pressure sensor while the air is being released. The pressure in the cuff fluctuates with the pulse. —

Commercially available blood pressure monitors therefore usually have a calculation algorithm which determines diastolic and systolic pressure from the cuff pressure alone. However, this measurement is inaccurate. In particular, in people whose typical pressure curve deviates from the norm, for example due to an illness, this type of blood pressure measurement can show highly inaccurate values because the calculation algorithm is geared towards a standard patient (see J. Liu: Patient-Specific Oscillometric Blood Pressure Measurement, IEEE Transactions on Biomedical Engineering 2015; 63 (6), 1220- 1228).

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. As soon as the last sound is heard, 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. However, the oscillometric measuring devices determine the maximum vascular oscillation and calculate the systolic and diastolic blood pressure using a mathematical formula.

In particular, blood pressure is calculated from the vibrations of the blood vessel wall (oscillometric measurement). Usually, 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.

In the case of a completely irregular heartbeat (absolute arrhythmia), the oscillometric measurement is subject to errors.

Many devices also pump up to the maximum and then measure during the release phase. The pressure built up is sometimes perceived as unpleasant.

It has not yet been possible to sufficiently mechanize the determination of the phases of an arterial Korotkof f signal.

Object of the invention

In contrast, the invention is based on the object of reducing the above-mentioned disadvantages of the prior art. In particular, 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. In particular, the phase transition is determined in connection with a blood pressure measurement. Summary of the invention

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.

Preferred embodiments and developments of the invention can be taken from the subject matter of the dependent claims, the description and the drawings.

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. In the sense of the invention, 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.

In particular, 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. In particular, 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.

In another embodiment, a measurement can be taken during inflation.

Also, as provided in one embodiment, 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.

Furthermore, the measuring device according to the invention 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. In a pulse phase, the arterial pulse causes a pressure fluctuation which can be detected by the additional sensor.

In the sense of the invention, 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.

According to one embodiment of the invention, the sensor 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.

Furthermore, the sensor 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. In particular, when measuring blood pressure, the transition to the first phase, third phase and/or fifth phase can be recorded more precisely (systolic pressure, mean arterial pressure, diastolic pressure). If 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.

In one embodiment, 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.

It is essential for this embodiment of the invention that the sensor detects vibrations in a low frequency range, in particular from 0.1 Hz. The lower In particular, the limit frequency should not exceed 10% of the typical pulse frequency.

In a further development of the invention, 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.

In this embodiment of the invention, 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.

It has been found that 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 inventor discovered that the length of this time offset depends on the phase of the arterial signal.

According to an embodiment of the invention, 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.

In particular, 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. For each pulse, 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).

In one embodiment of the invention, 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.

Due to the appropriate dimensioning of the sensor and cuff, 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.

In a preferred embodiment, 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.

According to the invention, on the one hand the pressure in a cuff extending around a vessel is measured. On the other hand the differential pressure between the cuff and the vessel is measured by means of a further sensor, in particular a differential pressure sensor. As described above, a physical signal such as force, voltage, etc., which depends on the differential pressure, can be used to determine the differential pressure. In particular, the sensor can be designed as a vibration sensor and/or surface microphone and have the measuring ranges described above.

Preferably, a sensor with a lower limit frequency of less than 0.5 Hz, preferably less than 0.3 Hz, in particular from 0.05 to 0.5, particularly preferably from 0.08 to 0.12 Hz is used.

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.

Furthermore, as already described above, 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. To determine the phases, 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 .

It has been found that the oscillometric signal reaches an inflection point in this phase .

At the same time, 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. In 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. In the fifth phase, 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.

Short description of the drawings The subject matter of the invention will be explained in more detail below with reference to the drawings Fig. 1 to Fig. 12.

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.

Detailed description of the drawings

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. In particular, 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.

Signals from the differential pressure sensor 15 can be transmitted to an external unit of the measuring device 10 via the connections 15c. 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 .

By means of the differential pressure sensor 15, a second measuring signal is generated, which is described in detail below.

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 .

When inflating, the pressure sensor 16 serves to limit the maximum pressure. For example, 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.

Furthermore, the 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.

Based on the signal from the cuff pressure sensor 16, 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.

In a first phase, the cuff is inflated and the pressure increases to a programmed target value.

This is followed by a short second phase in which the pump is switched off.

Subsequently, in a third phase, the air is let out of the cuff, with the pressure decreasing linearly, until in a fourth phase the air is let out of the cuff so that the pressure in the cuff corresponds to atmospheric pressure.

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 (piezo 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.

Due to the use of the proportional valve, the reduction is essentially linear.

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.

As can be seen in the detailed illustration in Fig. 2, 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.

In this representation, for example, in the same measuring interval of 0-40 seconds, 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.

In addition, 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.

It can be seen that the signal is different, especially in phase 3 where it drops less sharply.

Furthermore, the signal of the cuff pressure sensor is simultaneously plotted in Fig. 5a.

The signal from the cuff pressure sensor approximately follows the signal curve of the oscillatory signal, but is delayed by a time period At.

In this example it can be seen that At is significantly larger in the third phase than in the first phase.

In this embodiment, 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.

In particular, in 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.

In phase 1, blood begins to flow into the artery and a knocking sound occurs.

In phase 2, the artery is open and there are weak tapping sounds, superimposed by murmuring sounds.

In phase 3, the artery is so dilated that pure knocking sounds can be heard again. These are louder than in phase 1.

In 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.

It can be seen that both signals show characteristic changes during the phases . The same applies to the auscultatory signal, which is applied via a band filter, in this embodiment via a pass filter, from 20 to 100 Hz.

As shown below, the signals, both the oscillometric and the auscultatory signal, can be converted into a spectral representation for further processing.

For the oscillometric signal, as shown in Fig. 7, 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.

In the middle of the measurement, 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.

It can be seen that the fundamental wave and the distortion factor change characteristically with the beginning of the phases of the arterial signal marked 1 to 5.

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 Δt 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.

According to the invention, not all decision criteria necessarily have to be taken into account.

In particular, in a simple embodiment, a measuring device is provided in which the auscultatory signal and/or the time difference signal are dispensed with.

Thus, the beginning of 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. For the determination of phase 3, which is relevant for the mean arterial pressure, the maximum of the oscillometric fundamental wave is sufficient.

The additional consideration of the auscultatory signal of the differential pressure sensor and the time difference enables a more precise determination of the phase transitions.

All these values can be included in the calculation model of the measuring device, for example the blood pressure.

The corresponding algorithm can thus determine the blood pressure more accurately by narrowing it down more precisely, especially in phases 1 and 5.

At the same time, the measurement result is less sensitive to atypical signal curves, for example due to diseases.

In a further development, 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.

In particular, the measurement method can be used for diagnostic purposes for the following diseases:

- Arterial stiffness

- Endothelial function

- Heart valve insufficiency

- General heartbeat behaviour (e.g. ejection) or general hemodynamics.

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.

In addition to the ankle/brachial index, carotid wall analysis, and endothelial function measurement, the measurement of arterial stiffness is playing an increasingly important role in cardiovascular prevention. Arterial elasticity or extensibility (the opposite of stiffness) is defined as the relative change in volume in relation to changes in pressure.

The non-invasive examination of endothelial function serves as an early marker of the development and progression of early arteriosclerosis. Early diagnosis and risk stratification are goals in the prevention of coronary heart disease. 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.

In all these applications, 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.

In this context, for example, the relationship between reduced noise level in phase 2 and vascular calcification as well as the relationship between vascular elasticity and time difference signal around the phase transition 3/4 is relevant.

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.

In the equivalent circuit, 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.

TI Reference symbol list

10 Measuring device

11 Cuff 12 Cable to cuff pressure sensor

13 Pump

15 Sensor/Differential pressure sensor

15a Piezo layer

15b Conductive layer 15c Connection

16 Pressure sensor / cuff pressure sensor

17 Microcontrollers

18 Valve

20 Arm 21 Artery

22 bones

Claims

Expectations : 1. Measuring device for determining the phases of an arterial signal, in particular a blood pressure monitor, comprising an inflatable cuff that can be placed around an extremity, a sensor for detecting the pressure in the cuff and a differential pressure sensor for detecting pressure fluctuations between the cuff and the extremity. 2. Measuring device according to the preceding claim, characterized in that the differential pressure sensor is designed as a surface sensor, in particular as a piezo sensor, in particular a thin-film sensor. 3. Measuring device according to one of the preceding claims, characterized in that the differential pressure sensor has a measuring range from 0.05 Hz and/or up to 500 Hz, in particular from 0.2 and/or up to 200 Hz, in particular from 0.5 and/or up to 100 Hz. 4. Measuring device according to one of the preceding claims, characterized in that the measuring device comprises at least one bandpass filter which divides the signal of the differential pressure sensor into a low frequency and a high frequency component, in particular with a cutoff frequency between 10 and 30 Hz. 5. Measuring device according to one of the preceding claims, characterized in that the measuring device comprises a device for calculating the phases of the arterial signal, which calculates the beginning or the end of at least one phase taking into account a time offset between the signal of the sensor for detecting the pressure in the cuff and the differential pressure sensor. Measuring device according to one of the preceding claims, characterized in that the differential pressure sensor extends over less than 50%, preferably less than 20%, particularly preferably less than 10%, in particular over from 3 to 20%, preferably 5 to 15% of the length of the cuff. Measuring device according to one of the preceding claims, characterized in that the measuring device comprises a proportional valve for releasing the pressure from the cuff. Method for determining the phases of an artery signal, in particular with a measuring device according to one of the preceding claims, wherein the pressure in a cuff extending around a vessel is measured, wherein a differential pressure between the cuff and the vessel is measured via a sensor. Method according to one of the preceding claims, characterized in that a sensor with a lower limit frequency of less than 0.5 Hz, preferably less than 0.3 Hz, in particular from 0.05 to 0.5 Hz, particularly preferably from 0.08 to 0.12 Hz is used. . Method according to one of the preceding claims, characterized in that the signal of the sensor is divided into a low frequency and a high frequency component, in particular into an oscillometric and an auscultatory component, in particular with a cutoff frequency between 10 and 30 Hz. . Method according to one of the preceding claims, characterized in that the phases of the arterial signal are Consideration of an offset between the cuff pressure and the differential pressure between the cuff and the vessel, and/or characterized in that the phases of the arterial signal are determined taking into account a time offset between the cuff pressure and the differential pressure between the cuff and the vessel, and/or characterized in that the method is carried out while the pressure is being released from a cuff, and/or characterized in that the pressure is reduced essentially linearly. . Method according to one of the preceding claims, characterized in that an oscillometric signal, an auscultatory signal and a differential signal between the cuff pressure and the pressure between the cuff and the vessel are generated to determine the phases of an arterial signal. . Method according to the preceding claim, characterized in that the phases of the arterial signal are determined via turning points and/or local minima or maxima of the signals, in particular characterized in that the start of a - first phase by an inflection point of the oscillometric signal and/or by a local maximum of the distortion factor of the oscillometric signal, and/or - second phase by a local maximum of the auscultatory signal and/or a turning point of the Difference signals, and/or - third phase by a local maximum of the oscillometric signal and/or by a local minimum of the distortion factor of the oscillometric signal, and/or -fourth phase by a second phase by a local maximum of the auscultatory signal and/or a local maximum of the difference signal, and/or - fifth phase is 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. . Method according to one of the preceding claims, characterized in that the method is used to limit the time of the systolic and/or diastolic pressure during a blood pressure measurement. . Medical measuring device, in particular blood pressure measuring device, designed to carry out a method according to one of the preceding claims.