DE10033171A1 - Device for non-invasive, stress-free blood pressure measurement has non-invasive sensor device, signal processing unit for deriving blood flow/blood flow rate at least five times per second - Google Patents

Abstract

The arrangement has a non-invasive sensor device and a signal processing unit for deriving blood flow or blood flow rate at least five times per second from bio-sensor signals and for estimating the degree of propagation of a pulse wave and blood pressure using individual cardio-vascular parameters, e.g. determined using reference measurements, conventional Riva-Rocci measurements. The arrangement (1) has a non-invasive sensor device (4) and a signal processing unit for deriving blood flow or blood flow rate at least five times per second from bio-sensor signals and for estimating the degree of propagation of a pulse wave and blood pressure using individual cardio-vascular parameters determined using reference measurements, by conventional Riva-Rocci measurements, by measuring the individual vessel diameter, by statistical-empirical estimation with and/or without patient data and by pattern evaluation of sensor signals.

Description

environment

Blood pressure measurements are among the most commonly used medical in the world Examination procedures. Regular blood pressure monitoring helps cardiovascular Recognize and deal with risks in good time.

In addition, continuous blood pressure monitoring in medicine has a fixed one Importance and is indispensable in the monitoring and assessment of the physiological Condition of patients in intensive care, anesthesia and in pre- and Aftercare range.

State of the art

The majority of the non-invasive measuring devices are based on sphygmomanometric methods according to Riva-Rocci, in which blood pressure values can only be recorded selectively, a pulse- Pulse or even a continuous measurement cannot be achieved with this. The occlusions of the Extremities can lead to tissue trauma when measuring too often Furthermore, it is a burden for the patient. Furthermore, the determined Blood pressure values deviate considerably from the real values.

An ideal device should therefore be as reliable as possible, non-invasive, without external pressure and can continuously determine blood pressure. The most diverse new approaches here to meet these requirements only partially.

A method for arterial applanation tonometry is proposed in [Sec92]. there is used for non-invasive, continuous blood pressure measurement only partially on an artery pressure exercised. For artifact suppression, the fundamental and harmonics of the registered Filtered out pressure signal and thus reconstructed the pressure signal again.

A continuous process, which works without permanent external pressure, is described in [McQ93] specified. Two non-invasive ultrasound Doppler sensors are used attached to a larger artery. The blood flow signals measured with it are then used to create a simplified mathematical-empirical model of the Characterize the artery parametrically. The so-called resonance frequency determined of the model is correlated with blood pressure. A conventional cuff measurement the system calibrates to blood pressure at irregular intervals.

In [Elt98] a device is presented which initializes with as little as possible Calibration measurements can determine blood pressure accurately and robustly. Also suitable due to the measurement technology without sleeves, the procedure (Quasi) continuous blood pressure monitoring. In the present invention, this is Device for calibration measurements and sensors as well as signal processing significantly improved.

literature

[Elt98] Elter, P., Lutter, N., Müller-Glaser, KD, Stork, W., device for non-invasive blood pressure measurement, patent specification DE 198 29 544
[McQ93] McQuilkin, G., Noninvasive, non-occlusive method and apparatus which provides a continuous indication of arterial pressure and a beat-by-beat characterization of the arterial system, US Patent 5,241,964, 1993
[Sec92] Seca GmbH, method and device for non-invasive continuous blood pressure measurement in humans, published application DE 41 05 447 A1, 1992

Description of the invention Theoretical basics

For non-invasive, continuous determination of blood pressure without external pressure, the Blood pressure derived from other, non-invasively measurable cardiovascular sizes. To In the following, relationships between the three in the propagation of pulse waves in Vascular system of humans accompanying pulse forms (pressure, current and volume pulse) be derived.

An exact mathematical description of the hemodynamic processes in the vascular system itself is extremely difficult. It is more clever to consider simplified models and then map them to the vascular system. A tried and tested model of a vascular segment is that of an elastic hose.

The most general form of the equation of motion for flowing, incompressible liquids is the Navier-Stoke equation system:

with ρ = density, = velocity of the liquid in the three spatial directions, ∇ = Nabla operator, p = pressure, _A = external force density, η = viscosity of the liquid, Δ = Laplace operator.

The terms on the left represent the inertial force densities, the local and the convective acceleration. On the right side, grad p is the pressure gradient, and η.Δ corresponds to the friction force density.

To solve this partial, non-linear system of equations Eq. (1) can with regard to assumptions simplifying the arterial system. The convective Acceleration can be limited to sufficiently low speeds be ignored. The non-linear components can also be neglected because the phase velocity is much greater than the axial velocity component and this in turn is much larger than the radial speed component. As well the second derivatives after z can be neglected; external forces, such as Gravity, for example, and the radial pressure drop are also negligible.

Under these assumptions, the last equation from Eq. (1) can be represented in cylinder coordinates according to FIG. 1 as follows:

Due to the principle of superposition, this linearized differential equation can be solved with a harmonic separation approach from the superimposition of back and forth waves:

Here are: ω = 2.π.f = angular frequency, P _k , V _k = complex amplitudes of the k. harmonic component of the pressure or the velocity (index H: incoming wave, index R = returning wave), and γ _k is the propagation measure or the propagation constant of the k. harmonic component.

For the sake of simplicity, the differential equation is only to be solved below for a harmonic (k = 1) and for a passing wave:

p (z, t) = Pe ^{j ω t- γ z}

ν _Z (r, z, t) = V (r) .e ^{j ω t- γ z} (4)

With γ = α + jß = α + jω / C _Ph . α is the damping coating, β is the phase coating, which in turn can be expressed as the quotient of angular frequency and phase velocity C _ph . The solution of the system of equations must meet the boundary conditions on the inner wall of the vessel R ₀ . The hose in the z direction is assumed to be very rigid or longitudinally fixed. Furthermore, the axial acceleration component must assume the value zero at r = 0 due to the rotational symmetry.

The time-variant inner radius consists of a static and a dynamic part:

R ₀ (z, t) = R _{0, stat} + ΔR ₀ .e ^{j ω t- γ z} (6)

The Bessel differential equation Eq. (2) solve:

where J _{n represent} Bessel functions of the first kind of the nth order:

To determine the total flow i (z, t) in the hose, the following is integrated across the cross-section:

Eq. ( 9 ) can be calculated using the calculation rule 2 / xJ ₁ (x) = J ₀ (x) + J ₂ (x) and the substitution

simplify too

For the back and forth waves existing in the arterial system, this results in pressure and current pulses

The time courses of the pressure and the current pulse are sketched in FIG . It is clearly evident that according to Eq. (12) the pressure pulse from the sum, the current pulse from the difference of a back and forth wave.

If two locations z ₁ = 0 and z ₂ = 1 are now considered in accordance with FIG. 8 on a branch of the vessel, the following results with f (z, t) = F (z) .e ^{j ω t} :

With the calculation rules ½ (e ^x + e ^-x ) = cosh (x) and ½ (e ^x - e ^-x ) = sinh (x) the pressure P ₁ can be determined to:

Define the reflection factor _{r r} with

in this way, starting from Eq. (13) also determine the following calculation formula for blood pressure P ₁ :

Conclusions

• A) The blood pressure p _{1, total} , composed of the harmonic components p ₁ , can be calculated according to Eq. (14) can be determined from the harmonic components I ₁ and I _{2 of} the two current pulses i _{1, ges} and i _{2, ges} , the cardiovascular parameter Z _I and the degree of expansion γ.
• B) Eq. (16) offers to determine the blood pressure P _{1, total} , from the blood flow I _1, the cardiovascular parameter Z _I, the reflection factor r _r and the degree of spread γ.

The average speed pulse can also be used as an alternative to the current pulse:
= i / πR 2 | 0. Furthermore, the pressure p ₂ at the point z ₂ can alternatively also be determined.

measurement methods

In order to determine the blood pressure p _{1, ges} according to conclusion I., the current pulses i _{1, ges} and i _{2, ges are} measured at least two positions z ₁ and z _{2 of} a vascular branch at least five times per second, from which the is calculated harmonic components I ₁ and I ₂ , then determines it and possibly with a third current pulse i _{3, total} at a third measuring point z _3, the degree of propagation γ and can then with the z. As previously determined by calibration size of the harmonic components _I Z p ₁ of the blood pressure according to Eq. (14) and thus the total blood pressure p _{1, total} according to Eq. (3) determine.

Alternatively, conclusion II. To determine p _{1, ges} suggests measuring the current pulse I ₁ at a position z _{1 of} a vascular branch at least five times per second, then using it and with a further current pulse i _{3, ges} at a further measuring point z ₃ or by means of a size associated with the heart action, e.g. B. an electrocardiogram (EKG) to estimate the degree of expansion γ, in order to then determine the blood pressure p _{1, ges} and the variables Z _I and r _r .

According to this general measurement description, it should be closer to that used to calculate blood pressure necessary determination of the individual parameters.

The harmonic components I ₁ , I ₂ and I _{3 of} the current pulses i _{1, ges} , i _{2, ges} and i _{3, ges} are easiest to calculate using a Fourier transformation.

In addition to the angular frequency ω, Z _I is dependent on the radius R ₀ , the blood density ρ and the viscosity η, where ρ and η always, R ₀ can be assumed to be constant in a first approximation. Z _I can also be statistically and empirically intraindividually assumed to be constant or approximated by individual patient data such as age and gender or calibration measurements. If an even more precise determination of 21 is desired, then the volume pulsation R ₀ must be taken into account, which, however, is approximately in phase with the pressure pulse. The dynamic component ΔR ₀ according to Eq. (6) is usually only a few percent of the static part R _{0, stat} . For example, on the large common carotid artery with R _{0, stat} = approx. 5 mm with a pressure change of approx. 70 mmHg ΔR ₀ = approx. 0.15 mm (= 3%). You could influence this. B. determine by further empirical approximation or calibration measurements. Of course, the radius R _{0 can} also be measured directly, for example using means of locally high-resolution optical tomography or using ultrasound echo methods for larger vessels.

The reflection coefficient r _r essentially depends on the peripheral terminating resistance of the artery and, like Z _I , can be approximated statistically and empirically intraindividually, through individual patient data such as age and gender or calibration measurements.

When determining the propagation measure γ, the damping measure α and the phase measure β = ω / c _ph must be determined. In a first approximation, lossless transmission (α = 0) can be assumed. Then Eq. (14) and (16) to:

The phase velocity c _Ph in the phase measure β is generally different for each harmonic pulse wave due to the dispersive behavior of the human vascular system. Low-frequency parts of the pulse waves propagate more slowly than higher-frequency ones. If freedom from reflection is assumed, the frequency-dependent phase velocities can be determined by the phase difference of the individual harmonic components of the current pulses I ₁ and I ₂ . However, since there are reflections in the arterial system, this is only an approximate solution (the so-called "apparent" phase velocity). However, the "true" phase velocities can be approximated empirically or by calibration measurements. When using Eq. (18) or measurement method II. Lends itself to determining the spatially averaged phase velocity by means of an EKG-R wave and the current pulse I ₁ , ie the velocity of propagation of the pulse waves between the heart and the peripheral measurement point 21 . If the actual "true" phase velocities are to be determined, this requires a further flow measurement at a third location z _{3 of} the vascular system, because in the case of reflection according to Eq. (12) three equations have to be established to determine the three unknowns P _H , R _R and γ.

The measurement rate for blood flow determination of at least 5 Hz results from the following consideration: Human pulse rates are in the range between approx. 30 and 180 beats per minute, which corresponds to a minimum of 0.5 Hz or a maximum of 3 Hz. It can be shown that human pulses with the first harmonic components can be adequately described depending on the desired resynthesis accuracy. For example, the first 10 harmonics in resynthesis achieve a signal approximation of up to 99.5%. The relevant bandwidth of a pulse signal can be set in the range from approx. 2.5 Hz to 30 Hz ( Fig. 3). To fulfill the sampling theorem, at least twice the bandwidth must be sampled, i.e. at least 5 Hz. The same applies to the registration of the cardiac action or excitation, e.g. B. by means of an EKG.

Description of the system

The technical realization of the invention, Fig. 4 through Fig. 7, are removed. It consists of a main unit ( 1 ) with sensors ( 11 ), a removable sensor housing ( 4 ) with further sensors ( 12 ), which is connected to the main unit ( 1 ) via a self-winding cable ( 13 ). The sensor housing ( 4 ) can be completely integrated into the main housing ( 1 ). When removing the sensor housing ( 4 ) according to FIG. 6, the sensor distance ( 11 ) and ( 12 ) is automatically determined by the device, for example via the cable unwinding. It is also possible to connect the sensor housing ( 4 ) to the main housing ( 1 ) wirelessly, e.g. B. by radio, which is of particular interest in measurement variant II, if sensor ( 12 ) as an ECG lead z. B. is carried out by means of a chest strap.

The measured and calculated biosignals (blood flow, pulse, blood pressure, spread, EKG, etc.) can be displayed graphically and / or alphanumerically on the LCD display ( 2 ). It is also possible to make settings using the keypad ( 3 ).

According to variant I, the sensors ( 11 ), ( 12 ) according to FIG. 8 measure at least two blood flow signals i _{1, ges} and i _{2, ges} at two different locations z ₁ and z _{2 of} a vascular branch of the body (preferably on an arterial branch), see above that, apart from other parameters, it can also be used to determine the degree of dispersion γ. The device can be attached to the human extremities with bands ( 5 ), ( 6 ), e.g. B. the arm ( 7 ), placed on superficial vessels. For this purpose, the brachial artery ( 8 ) in the crook of the arm, the radial or ulnar artery ( 9 ) on the wrist, or the digital artery ( 10 ) on the finger are available as larger vessels on the arm. Since the phase velocities of the pulse waves are quite high (approx. 5 to 15 m / s), it is advisable not to place the sensors ( 11 ) and ( 12 ) too close to one another for higher measuring accuracy. At a distance of 1 = 5 cm and an average phase velocity of C _Ph = 10 m / s, there is a temporal phase difference of 5 ms, so that the dynamic phase differences to be recorded have to be resolved in the 1/10 millisecond range. If the measuring distance I, for example between the brachial artery ( 8 ) and the radial artery ( 9 ), increases to approx. 25 cm, the requirements for temporal resolution are of course significantly lower (approx. Millisecond range), since the phase differences to be recorded have increased ,

With measurement variant II., The requirements for temporal resolution are further drastically reduced. The sensor ( 12 ) for detecting the heart action is implemented, for example, as a simple ECG lead, which is designed as a wireless breast sensor including a chest strap and continuously sends the EKG signal by radio to the main device ( 1 ). Another possibility is, for example, an EKG lead left arm-right arm according to Einthoven I, with one electrode in the main device ( 1 ) and the other in the removable sensor ( 12 ). The sensor ( 11 ) in the main device ( 1 ) still measures the blood flow in a blood vessel, e.g. B. ( 8 ), ( 9 ), ( 10 ). To determine the average phase velocity, the time of propagation of the pulse waves from the heart (EKG-R wave) to the peripheral measuring point of the blood flow measurement is now used. Another measure for registering the heart action is, for example, the heart tones.

Doppler technology is used to obtain the blood flow signals i _{1, ges} , i _{2, ges} , i _{3, ges} . According to FIGS. 9 and Fig. 10 either ultrasound at a well-defined frequency in the MHz range (the default value about 8 MHz for shallow vessels) or laser radiation with a well defined wavelength preferably irradiated in the near infrared range of about 700-3000 nm in the tissue.

In the case of ultrasound, the sound waves radiated at an angle α with the frequency f ₀ on moving blood particles at the speed v are scattered back with a Doppler frequency shift Δf:

Here c is the speed of sound propagation and is in human tissue approx. 1450 m / s. In this way, the backscattered ultrasound is obtained on Ultrasound detector a measure of blood flow.

In the optical case, the tissue is irradiated with coherent laser light of wavelength λ ₀ . However, the light is also scattered on static tissue parts due to the significantly shorter free path length compared to ultrasound, so that the photons - if at all - hit a moving blood particle at any angle and, depending on that, receive a more or less large Doppler shift. Thus, a maximum Doppler frequency shift Δf _max can be detected at the photodetector only in the case of single scattering on a moving blood particle v:

n is the refractive index of tissue.

The advantages of the laser Doppler method compared to the ultrasonic Doppler method are that the light can be coupled into the tissue without contact gel (although in principle it would also be possible with ultrasound without gel, which, however, is very high coupling losses is connected). Secondly, you get a much higher one local resolution so that blood flow is measured even in the smallest vessels such as capillaries can be. However, this also increases the signal bandwidth of the detected signal. The advantage of the ultrasound procedure lies in the possibility of absolutely reducing the blood flow measure, whereas the laser Doppler technique only allows a relative measurement and thus relies on an initial calibration.

The received sensor signals are gem. 11 is supplied with an analog part (14) to the block diagram of Fig., and then further processed in a digital Signalverabeitungseinheit (15). The system also includes memory modules ( 16 ), an input / output unit ( 17 ) and a display unit ( 18 ).

In the analog section, Fig. 12, the sensors (11), (12), (19) (It may, as shown in the following, other than the two sensors 1 and 2 further sensors are used. These are "virtual" sensor m ( 19 ) summarized.) Excited ( 20 ). In the case of flow measurement, this takes place in the ultrasound case via an oscillator, in the optical case with a laser diode driver circuit which ensures constant optical power and protection of the laser diode from being switched on or off. No excitation is required for the ECG. The received signals are then bandpass filtered ( 22 ) and amplified with automatic gain control (AGC) ( 23 ). It is important that the analog modules have constant group or phase delays so as not to falsify the phase information of the signals, which would be unfavorable with regard to the phase or pulse wave velocity determination. Useful lower limit frequencies of the bandpass are around 100 Hz in order to eliminate the high signal component, 50 Hz network interference and the low-frequency plethysmographic signal components in the frequency range from 1 Hz to approx. 30 Hz. In flow measurement, they result in wall or sensor movements and, in the visual case, in the time-changing blood volumes. These optical plethysmographic components are disturbing for the flow determination, but can also contribute further information to the determination of the blood pressure. The time course of the plethysmographic signal is qualitatively linked to the volume pulse and thus to the pressure pulse.

In the case of ultrasound, the signals are demodulated after amplification ( 24 ) and bandpass filtered again. Before the AD conversion ( 27 ), an amplification ( 26 ) is carried out again before it is gem. Fig be further processed in the digital signal processing unit (15). 13,.

Both when recording the heart action (ECG) and in the visual case, there is one Demodulation is not necessary. With laser flow measurement, this is already done at the photodetector by interference of the frequency-shifted and non-frequency-shifted beams automatically.

The upper limit frequency of the bandpass ( 22 ) or ( 25 ) depends on the sampling rate of the AD converter ( 15 ). It is used to suppress noise and aliasing effects. When ultrasound is an upper limit frequency of about 20 kHz (25), useful for laser beams of about 200 kHz (22), in detecting the cardiac action about 2 kHz (22).

In the signal processing unit ( FIG. 13), which is controlled by an overall control ( 28 ), artifacts and other disturbances are first eliminated by adaptive or parametric filters ( 29 ) and then the blood flow is determined continuously, ie at least five times per second ( 30 ). This can be done in the time domain using simple zero crossing counting. This procedure is simple and quick, but inaccurate. More precise procedures work in the frequency domain. For this purpose, spectrograms are determined from the time signal for short time intervals, which can be estimated via a Fast Fourier Transformation (FFT) or parametrically. It can be shown that the mean flow velocity can be determined via the first weighted moment in the spectral range:

In the optical case, the procedure is the same. It can also be shown here that the first weighted moment in the spectral range is proportional to the mean flow velocity:

When registering the heart action z. B. with an EKG is none other than filtering further signal processing necessary.

The propagation measure γ ( 33 ) is determined on the basis of at least two flow pulses i _{1, ges} and i _{2, ges} or a variable linked to the heart action (e.g. EKG) and a flow pulse. For this purpose, the time signals i _{1, ges} and i _{2, ges} are transformed into the frequency range _, for example by means of an FFT, and the amount and phase of the harmonic components I ₁ and I ₂ are thus obtained. As explained above, the phase measure β = ω / C _Ph can then be calculated. To determine the mean phase measure using an EKG and a flow pulse, the time difference from the start of the pulse waves at the heart (EKG-R wave) to the peripheral measuring point of the blood flow measurement is now used.

Since only the "apparent" phase velocities can be determined with these methods, the "true" phase velocities and attenuations can be approximated empirically or by calibration measurements to increase the measuring accuracy. The exact calculation of the "true" phase velocities and the damping requires a further flow measurement at a location 23 of the vascular system ( 19 ). This also increases the measuring accuracy and reduces the number of calibration measurements required.

Then the individual harmonic components of blood pressure p ₁ ( 35 ) can be calculated with the parameter Z _l and possibly the reflection factor r _r . Ultimately, the continuous course of the blood pressure p _{1, total} , in particular the systolic, diastolic and mean blood pressure, can thus be determined.

The cardiovascular parameter Z _l and the reflection factor r _r are determined, depending on the desired reliability and measurement accuracy, analytically according to the individual calibration ( 32 ) or statistically and empirically with individual patient data such as age and gender. In order to increase the measuring accuracy and reduce the number of calibration measurements, the essential parameter in Z ₁ the vessel radius R _{0 can} also be measured directly by means of a further non-invasive sensor ( 19 ). Local high-resolution optical tomography is suitable for determining the radii of small vessels, and the known ultrasound echo methods can be used for larger vessels.

Gravity can be used for calibration. By raising or lowering a well-defined height of the extremity to which the sensors are attached, other parameters can be determined on the basis of the resulting additional, known hydrostatic blood pressure. Furthermore, the pulse contours in the blood flow and blood pressure can provide information about the necessary cardiovascular sizes of the vessel radius R ₀ and reflection factor _{r r} .

To ensure that the sensors ( 11 ), ( 12 ), ( 19 ) provide optimal signals and that the device can be easily applied to humans, the flow sensors can be positioned manually or automatically. In the automatic case, the sensor signals can be continuously tracked before, during or between the measurement and calculation phases, so that the initial optimal application or slipping of the device is always based on the same optimal measurement distance (above the desired blood vessel). The associated control algorithm ( 31 ) is implemented in the signal processing unit ( 15 ), the hardware controls ( 21 ) are integrated in the sensors ( 11 ), ( 12 ), ( 19 ) themselves and in the analog part ( 14 ).

There are several implementations as adaptive flow sensors, both in the acoustic and in the optical case. According to FIG. 14, the first possibility includes, in both cases, mechanical tracking of the actual sensor ( 37 ) consisting of transmitting ( 38 ) and receiving piezo crystal ( 39 ) or laser diode ( 38 ) and photodetector ( 39 ), for example a movable one Stepper motor ( 36 ) that can travel along on a spindle ( 40 ). The sensor unit ( 11 ), ( 12 ), ( 19 ) is then applied to the arm in such a way that the travel path runs across the blood vessel. The motor ( 36 ) is addressed via the hardware control ( 21 ) and regulated via the control algorithm ( 31 ).

A second possibility is an ultrasound or laser array FIG. 15. There are many individual sensors ( 41 ) (consisting of transmit and receive piezo crystal or laser diode and photodetector), for example, arranged in a linearly offset manner in two rows, which can be controlled or selected via a multiplexer circuit ( 21 ). The control takes place via the control algorithm ( 31 ).

Another alternative in the optical case can be realized with light guiding devices such as optical fibers. In the basic diagram according to FIG. 16, a virtual array is built there with the aid of micromechanically adjustable mirrors ( 45 ), in which only one sensor element ( 43 ) is required. A pair of light guide elements of the bundle ( 42 ) is selected by rotating the mirror ( 45 ) by the angle ϕ and / or by driving a distance d along the optical beam axis ( 44 ), so that exactly one light supply of the strand is selected. A further possibility for selecting a pair of light guide elements is expanded ( 47 ) in FIG. 17, beam expander ( 46 ), and then a pair of light guide elements is selected with an electronically controllable LCD shutter ( 48 ) by darkening ( 49 ) and non-darkening ( 50 ).

Claims (22)

1. Device for non-invasive blood pressure measurement on living beings, characterized in that

• a) a non-invasive sensory device
  • - measures two quantities linked to the blood flow or the blood flow velocity at different positions on the blood vessel system of a living being, or
  • - measures a size associated with blood flow or blood flow velocity in the blood vessel system of a living being and a size linked with heart action,
• b) a signal processing unit
  • - estimates the blood flow or the blood flow rate from the at least two received bio-signals at least five times per second, by means of two variables linked to the blood flow or blood flow velocity or one variable each associated with the blood flow or the blood flow velocity and the heart action, estimates the spread of the pulse waves,
  • - and determines the blood pressure continuously and / or quasi-continuously with the blood flow or the blood flow velocity, the spread of the pulse waves and individual cardiovascular parameters,
• c) where the individual cardiovascular parameters can be determined
  • by means of reference measurements via variation of the hydrostatic blood pressure due to a well-defined difference in height between the measuring point of the blood flow pulse and the heart level,
  • - using the conventional Riva-Rocci method,
  • - by measuring the individual vessel diameter,
  • - by means of statistical-empirical estimation with and / or without individual patient data,
  • - by means of a sample evaluation of the detected sensor signals.

2. Device according to claim 1, characterized in that the flow sensors ( 11 ), ( 12 ) optically determine the quantities associated with the blood flows or the blood flow velocities according to the Doppler effect with coherent light. 3. Device according to claim 1, characterized in that the flow sensors ( 11 ), ( 12 ) determine the quantities associated with the blood flows or the blood flow velocities acoustically with ultrasound according to the Doppler effect. 4. The device according to claim 1, characterized in that in the signal processing unit ( 15 ) the blood flows or the blood flow velocities are determined from the zero crossings of the received sensor signals ( 29 ), ( 30 ). 5. The device according to claim 1, characterized in that in the signal processing unit ( 15 ) the blood flows or the blood flow velocities are determined by moment formation of the spectrograms of the received sensor signals, the spectrograms being able to be estimated via a Fourier transformation, curve fitting or by means of parametric methods ( 29 ), ( 30 ). 6. The device according to claim 1, characterized in that in the signal processing unit ( 15 ) the harmonic components of the blood flows or the blood flow velocities are estimated by a Fourier transformation, curve fitting or by means of parametric methods ( 30 ). 7. The device according to claim 1, characterized in that in the signal processing unit ( 15 ), the phase velocities of the propagation measure ( 33 ) are determined with the phase differences of the harmonic components of the blood flows or the blood flow velocities or via cross-correlation, and that this estimate is optionally statistical-empirical and / or improved by individual calibration measurements ( 32 ). 8. The device according to claim 1, characterized in that in the signal processing unit ( 15 ), the phase velocities of the propagation measure ( 33 ) are determined with the time differences of the propagation times of the pulse waves of electrocardiogram and blood flow or blood flow velocity, and that this estimate is optionally statistical-empirical and / or is improved by individual calibration measurements ( 32 ). 9. The device according to claim 1, characterized in that the attenuation is neglected in the signal processing unit ( 15 ) when calculating the propagation measure ( 33 ), or is included as a statistical-empirical and / or constant determined by individual calibration measurements ( 32 ). 10. The device according to claim 1, characterized in that at a third point a further non-invasive sensor ( 19 ) measures a size linked to the blood flow or the blood flow rate on the blood vessel system of a living being, in the digital signal processing unit ( 15 ) the blood flows or blood flow rates are determined ( 30 ) in order to then precisely determine the phase velocities and the damping of the propagation measure ( 33 ) with the harmonic components. 11. The device according to claim 10, characterized in that the further sensor ( 19 ) optically determines the quantities associated with the blood flows or the blood flow velocities according to the Doppler effect with coherent light. 12. The device according to claim 10, characterized in that the further sensor ( 19 ) determines the quantities associated with the blood flows or the blood flow velocities acoustically using ultrasound according to the Doppler effect. 13. The apparatus according to claim 10, characterized in that in the signal processing unit ( 15 ) the blood flows or the blood flow velocities are determined from the zero crossings of the received sensor signal ( 29 ), ( 30 ). 14. The device according to claim 10, characterized in that in the signal processing unit ( 15 ) the blood flows or the blood flow velocities are determined by torque formation of the spectrograms of the received sensor signal, the spectrograms being able to be estimated via a Fourier transformation, curve fitting or by means of parametric methods ( 29 ), ( 30 ). 15. The device according to claim 10, characterized in that in the signal processing unit ( 15 ) the harmonic components of the blood flows or the blood flow velocities are estimated by a Fourier transformation, curve fitting or by means of parametric methods ( 30 ). 16. The device according to claims 1 or 10, characterized in that a cardiovascular parameter is determined by measuring the vessel radius with the means of optical tomography with a further non-invasive sensor ( 19 ). 17. The device according to claims 1 or 10, characterized in that a cardiovascular parameter is determined by measuring the vessel radius according to the ultrasound echo method with a further non-invasive sensor ( 19 ). 18. Device according to claims 1 or 10, characterized in that the sensor devices ( 4 ) with the main device ( 1 ) communicate wirelessly telemetrically. 19. Device according to claims 1 or 10, characterized in that the sensors ( 11 ), ( 12 ), ( 19 ) for signal optimization are automatically and / or manually positioned on the living being via suitable blood vessels. 20. The apparatus according to claim 23, characterized in that the positioning of the sensors ( 11 ), ( 12 ), ( 19 ) with a mechanical movement ( 36 ) - ( 40 ), which can be controlled via a digital control algorithm ( 31 ) ( 21 ). 21. The apparatus according to claim 23, characterized in that the sensors ( 11 ), ( 12 ), ( 19 ) each form a matrix structure ( 41 ) which can be controlled ( 21 ) via a digital control algorithm ( 31 ). 22. The apparatus according to claim 23, characterized in that the optical sensors ( 11 ), ( 12 ), ( 19 ) via light guide devices ( 42 ) are coupled to the skin, in which a pair of light guide elements, for example via adjustable mirrors ( 45 ), via LCD -Shutter ( 48 ) or Bragg cells is selected.