Method and device for determining intracellular and/or extracellular, in particular macromolecular fractions of fluids, preferably of body fluids of living organisms

Abstract

Method and device according to the method for determining intracellular and/or extracellular, in particular macromolecular fractions of fluids, preferably of body fluids of living organisms, with the steps: coupling-in a measurement signal through an electrically non-conductive wall into the fluid to be measured; coupling-out an electrical measurement value that is thereby generated in the fluid to be measured; detecting the coupled-out electrical measurement value at a plurality of different frequencies of the electrical measurement signal; determining the intracellular and/or extracellular, in particular macromolecular fractions of the fluid to be measured by means of evaluation of the detected electrical measurement value at a plurality of frequencies of the measurement signal.

Claims

1. Method for determining intracellular and extracellular fractions of a fluid using a device, wherein the device comprises a coupling-in device, a measurement signal generator, whose measurement signal can be capacitively and inductively coupled in, simultaneously or alternatively, via the coupling-in device through an electrically non-conductive wall into the fluid to be measured, a coupling-out device by which an electrical measurement value that is generated by the coupled-in measurement signal in the fluid to be measured can be capacitively and inductively coupled out, simultaneously or alternatively, through the electrically non-conductive wall, a detecting device configured for detecting the coupled-out electrical measurement value, and an evaluation device configured for determining the intracellular and extracellular fractions of the fluid by calculation from a plurality of measurement values generated at different frequencies of the measurement signal that are detected by the detecting device, wherein the inductive coupling-in is employed for frequencies of the measurement signal lower than the frequencies of the measurement signal for which the capacitive coupling-in is employed, and wherein the inductive coupling-in is supplied by at least one coupling coil, the method comprising the steps of: coupling-in with the coupling-in device a measurement signal generated by the measurement signal generator through the electrically non-conductive wall into the fluid to be measured; coupling-out with the coupling-out device an electrical measurement value that is thereby generated in the fluid to be measured; detecting with the detecting device the coupled-out electrical measurement value at a plurality of frequencies of the measurement signal; and determining with the evaluation device at least one of the intracellular and extracellular fractions of the fluid to be measured by evaluation of the detected electrical measurement value at a plurality of frequencies of the measurement signal.

2. Method according to claim 1, characterized in that the measurement signal is at least one of an electrical, magnetic, and electromagnetic alternating field, wherein the frequency of the alternating field is variable and modulatable.

3. Method according to claim 1, characterized in that each of the coupling-in of the measurement signal into the fluid to be measured, and the coupling-out of the electrical measurement value thereby generated in the fluid to be measured, takes place capacitively, inductively, or capacitively and inductively.

4. Method according to claim 1, characterized in that at least a portion of the measurement signal is capacitively coupled into the fluid to be measured, and at least a portion of the electrical measurement value thereby generated in the fluid to be measured is capacitively coupled out.

5. Method according to claim 1, characterized in that the coupling-in device has at least one pair of coupling-in electrodes, and the coupling-out device has at least one pair of coupling-out electrodes, and characterized in that the capacitive coupling-in of the measurement signal into the fluid to be measured takes place by using the pair of coupling-in electrodes, and the coupling-out of the electrical measurement value thereby generated in the fluid to be measured takes place by using the pair of coupling-out electrodes as four-point measurement.

6. Method according to claim 1, characterized in that each of the coupling-in of the measurement signal into the fluid to be measured, and the coupling-out of the electrical measurement value thereby generated in the fluid to be measured, takes place capacitively, inductively, or capacitively and inductively.

7. Method according to claim 1, characterized in that the determining at least one of the intracellular and extracellular fractions of the fluid comprises detecting electrical impedance of the fluid to be measured at a plurality of different frequencies of the measurement signal according to amplitude and phase.

8. Method according to claim 1, characterized in that the determining at least one of the intracellular and extracellular fractions of the fluid comprises evaluation on the basis of the Cole model.

9. Device configured for determining intracellular and extracellular fractions of a fluid comprising a coupling-in device, a measurement signal generator, whose measurement signal can be capacitively and inductively coupled in, simultaneously or alternatively, via the coupling-in device through an electrically non-conductive wall into the fluid to be measured, a coupling-out device by which an electrical measurement value that is generated by the coupled-in measurement signal in the fluid to be measured can be capacitively and inductively coupled out, simultaneously or alternatively, through the electrically non-conductive wall, a detecting device configured for detecting the coupled-out electrical measurement value, an evaluation device configured for determining the intracellular and extracellular fractions of the fluid by calculation from a plurality of measurement values generated at different frequencies of the measurement signal that are detected by the detecting device, wherein the inductive coupling-in is employed for frequencies of the measurement signal lower than the frequencies of the measurement signal for which the capacitive coupling-in is employed, and wherein the inductive coupling-in is supplied by at least one coupling coil.

10. Device according to claim 9, characterized in that the coupling-in device and/or the coupling-out device have flat electrodes for capacitive coupling-in of the measurement signal and for capacitive coupling-out of the electrical measurement value that is generated in the fluid to be measured and/or at least one coil for the inductive coupling-in of the measurement signal and/or at least one sensor for the measurement of a magnetic field influenced by an electrical value generated in the fluid to be measured.

11. Device according to claim 9, characterized in that the coupling-in device has at least one pair of coupling-in electrodes, and the coupling-out device has at least one pair of coupling-out electrodes, wherein the coupling-out electrodes are substantially disposed between the coupling-in electrodes.

12. Device according to claim 11, characterized in that the coupling-in and coupling-out electrodes are disposed on the outside of a fluid line which conveys the fluid with the intracellular and extracellular fractions that are to be determined.

13. Device according to claim 9, characterized in that the detecting device has a device for detecting an impedance according to amplitude and phase.

14. Device according to claim 9, characterized in that the coupled-out electrical measurement values can be evaluated in the evaluation device on the basis of the Cole model.

15. Dialysis machine comprising a dialyser and at least one device according to claim 9, wherein the dialysis machine is configured to be controlled or regulated depending on the determined intracellular and extracellular fractions of the fluid.

16. Dialysis machine according to claim 15, characterized in that the device is disposed downstream of the dialyzer and is configured for determination of a water fraction, which controls or regulates a transmembrane pressure in the dialysis device.

17. Dialysis machine according to claim 15, characterized in that the device is configured for detection of air bubbles and/or detection of hemolysis, and if air bubbles and/or hemolysis are detected during a dialysis circulation, a warning signal is triggered and/or the dialysis circulation is interrupted by the device.

18. Dialysis machine according to claim 15, characterized in that the intracellular and extracellular fractions of the fluid are determined.

19. Dialysis machine comprising a dialyzer and at least one device according to claim 9.

20. Device according to claim 9, wherein the at least one coupling coil is an exterior coil and for generating a magnetic field.

21. Device according to claim 9, wherein the inductive coupling-in is supplied via a plurality of coupling coils.

22. Device according to claim 9, wherein the capacitive measurement forms a high-pass filter, wherein the electrical measurement value generated in the fluid to be measured by a low-frequency end of the measurement signal is coupled-out inductively by coupling coils, and wherein the electrical measurement value generated in the fluid to be measured by a high-frequency end of the measurement signal is coupled-out capacitively.

23. Device according to claim 9, wherein the coupled-in measurement signal is at least one of an alternating electrical, alternating electromagnetic, or alternating magnetic field frequency-modulated over a frequency of 5 kHz to 1 MHz.

Description

DESCRIPTION OF THE FIGURES

(1) Two example embodiments of the invention are explained in detail below with the aid of the drawings. The drawings show:

(2) FIG. 1 current paths through a blood sample at a low measurement signal frequency

(3) FIG. 2 current paths through a blood sample at a high measurement signal frequency

(4) FIG. 3 Cole equivalent circuit diagram for the blood sample

(5) FIG. 4 idealized locus curve of blood in resistance-reactance diagram

(6) FIG. 5 equivalent circuit diagram for the measurement of impedance using a impedance analyzer

(7) FIG. 6 measured impedance locus curve (Z.sub.disp) of 0.9% aqueous NaCl solution and of blood at variable ultrafiltration volumes

(8) FIG. 7 calculated impedance locus curve of blood at variable ultrafiltration volumes

(9) FIG. 8 comparison between the hemoglobin concentration and the calculated BIS factor (bioimpedance)

(10) FIG. 9 comparison between the plasma protein concentration and the calculated BIS factor (bioimpedance)

(11) FIG. 10a schematic diagram of the measurement setup for determining the blood impedance

(12) FIG. 10b practical measurement setup for determining the blood impedance (blood tube with BCM electrodes applied)

(13) FIG. 11 schematic diagram of a dialysis machine with devices for determining the blood impedance before and after the dialyzer

(14) FIG. 12a variation in the amplitude of the blood impedance at 1 Mhz (injection of air bubbles indicated by *)

(15) FIG. 12b variation in the phase of the blood impedance at 1 Mhz (injection of air bubbles indicated by *)

(16) FIG. 13 variation in the m.sub.Hb/m.sub.pro quotient on occurrence of hemolysis (abscissa values above 110)

(17) The measurement method proposed here is based in a first approximation on the assumption that blood is a suspension of blood cells 5 (predominantly red blood cells) in plasma water 3, containing primarily dissolved ions and protein molecules 4 (of which albumins are present in the greatest quantity). If one applies a measuring current with a low frequency between two electrodes 1, 2, the current flow occurs almost exclusively through the plasma water 3 (FIG. 1), while at a high frequency the measuring current flows through the plasma water 3 and the blood cells 5 (FIG. 2), because the cell membranes of the blood cells 5, which isolate direct current and have an effect similar to capacitors, represent negligible resistance for such high frequencies. Thus the known Cole model, consisting of an ohmic resistor connected in parallel with a series connection of an ohmic resistor and a capacitor (FIG. 3), can be used as an equivalent circuit diagram for such a blood sample. In a resistance-reactance diagram, there thus results, for a plurality of different measurement frequencies, an idealized locus curve as shown in FIG. 4, wherein the volume of plasma water V.sub.plasma can be calculated from the resistance R.sub.E and the volume of the erythrocytes V.sub.RBC from the volume R.sub.I.

(18) $\begin{array}{cc}\begin{array}{c}{V}_{\mathrm{plasma}}={\left(\frac{l_{\mathrm{Schlauch}}.Math.\sqrt{V_{\mathrm{total}}}}{{\rho }_{\mathrm{plasma}}.Math.R_E}\right)}^{\frac{2}{3}}\\ ={\left(\frac{l_{\mathrm{Schlauch}}.Math.\sqrt{V_{\mathrm{total}}}}{{\rho }_{\mathrm{plasma}}}\right)}^{\frac{2}{3}}.Math.{\left(\frac{1}{R_E}\right)}^{\frac{2}{3}}\\ =k_{\mathrm{plasma}}.Math.{\left(R_E\right)}^{-2/3}\end{array}\begin{array}{c}{V}_{\mathrm{RBC}}={\left(\frac{l_{\mathrm{Schlauch}}.Math.\sqrt{V_{\mathrm{total}}}}{{\rho }_{\mathrm{RBC}}.Math.R_I}\right)}^{\frac{2}{3}}\\ ={\left(\frac{l_{\mathrm{Schlauch}}.Math.\sqrt{V_{\mathrm{total}}}}{{\rho }_{\mathrm{RBC}}}\right)}^{\frac{2}{3}}.Math.{\left(\frac{1}{R_I}\right)}^{\frac{2}{3}}\\ =k_{\mathrm{RBC}}.Math.{\left(R_I\right)}^{-2/3}\end{array}& \mathrm{Formula}1,2\end{array}$

(19) Here the two values I.sub.Schlauch and V.sub.total are known from the measurement setup: I.sub.Schlauch is the length and V.sub.total the volume of the measurement area [“Schlauch”=“tube” in German]. The conductivities of plasma water and erythrocytes are ρ.sub.plasma and p.sub.RBC respectively. Since length, volume and conductivity can if necessary be assumed to be constant, they can be combined into the constants k.sub.plasma and K.sub.RBC, which can be determined experimentally. Using the known volumes V.sub.total and V.sub.RBC, the hematocrit value Htc can be calculated from the determining equation:
Hct=(V.sub.RBC/V.sub.total)×100%  Formula 3

(20) If necessary, the hemoglobin concentration in the blood can be calculated from the Hct. With the known volumes V.sub.total, V.sub.RBC and V.sub.plasma, the volume of solids can be calculated. Assuming that the solids consist substantially of proteins, the volume of proteins in the blood V.sub.protein is consequently determined:
V.sub.protein=V.sub.total−V.sub.RBC−V.sub.plasma  Formula 4

(21) If one assumes that the solids consist substantially of proteins, the density of protein (D.sub.protein=1.4 kg/l) can be used to calculate the plasma protein concentration C.sub.protein:
c.sub.protein=(V.sub.protein×D.sub.protein)/V.sub.plasma  Formula 5

(22) Thus it is possible from a bioimpedance measurement of the blood to determine the protein concentration, the hematocrit value, and if necessary the hemoglobin concentration.

(23) FIG. 5 shows the circuit diagram of an impedance analyzer for the determination of a bioimpedance Z.sub.1: via the electrodes with the impedance Z.sub.2, the current source and voltage electrodes of the impedance analyzer are connected to the bioimpedance. The impedance analyzer itself has the internal resistance (or internal impedance) Z.sub.3, over which the voltage drop is measured.

(24) Assuming that the internal resistance of the impedance analyzer is significantly greater than the bioimpedance Z.sub.1 to be measured, the major part of the measuring current i.sub.1 flows over the bioimpedance, and the voltage drop u.sub.3 measured by the BCM corresponds to the voltage drop over the bioimpedance u.sub.1. The voltage Z.sub.disp indicated by the BCM is calculated according to:
Z.sub.disp=u.sub.3/i.sub.1  Formula 6

(25) If the bioimpedance is sufficiently low compared to the internal resistance, Z.sub.disp corresponds to the bioimpedance Z.sub.1. If this is not the case, because the bioimpedance assumes values that are too high, the measured impedance Z.sub.disp no longer corresponds to the bioimpedance Z.sub.1.

(26) This is the case when the impedance of blood in the blood tube is measured with capacitive coupling. FIG. 6 shows such impedance locus curves when measuring either a 0.9% NaCl solution or blood concentrated by ultrafiltration in the blood tube:

(27) With the aid of a voltage divider and a current divider, the resistance ratio, which corresponds to the indicated impedance Z.sub.disp, can be calculated:

(28) $\begin{array}{cc}\begin{array}{c}\frac{u_3}{u_1}=\frac{Z_3}{Z_2+Z_3}.Math.u_3\\ =u_1.Math.\frac{Z_3}{Z_2+Z_3}\end{array}\begin{array}{c}\frac{i_2}{i_1}=\frac{Z_2+Z_3}{Z_1+Z_2+Z_3}.Math.i_1\\ =i_2.Math.\frac{Z_1+Z_2+Z_3}{Z_2+Z_3}\\ =\frac{u_2}{Z_1}.Math.\frac{Z_1+Z_2+Z_3}{Z_2+Z_3}\end{array}\begin{array}{c}{Z}_{\mathrm{disp}}=\frac{u_3}{i_1}\\ =\frac{Z_1.Math.Z_3}{Z_1+Z_2+Z_3}\end{array}& \mathrm{Formula}7\text{-}9\end{array}$

(29) When the bioimpedance Z.sub.1 is known, for example in the case of 0.9% aqueous NaCl solution in the tube, the internal resistance of the BCM Z.sub.3 can be calculated from the indicated impedance Z.sub.disp with the aid of Formula 4.

(30) When the impedances Z.sub.3 (determined with the aid of a 0.9% aqueous NaCl solution) and Z.sub.2 (from the electrode geometry) are known, the bioimpedance Z.sub.1 can be calculated by means of Formula 4. FIG. 7 shows this for the concentrated blood:

(31) In the next step, the volumes of erythrocytes V.sub.RBC and plasma water V.sub.plasma can be calculated from the resistances.
V.sub.plasma=k.sub.RBC.Math.(R.sub.1).sup.−2/3
V.sub.plasma=k.sub.plasma.Math.(R.sub.E).sup.−2/3  Formula 10, 11

(32) When the total volume V.sub.total is known, it can be used with V.sub.RBC to calculate the hematocrit value Htc:

(33) $\begin{array}{cc}\mathrm{Hkt}=\frac{V_{\mathrm{RBC}}}{V_{\mathrm{total}}}=k_{\mathrm{Hkt}}.Math.{\left(R_I\right)}^{-2/3}& \mathrm{Formula}12\end{array}$

(34) That is to say, Hct is related to the Cole resistance R.sub.I. In the laboratory experiment, the hemoglobin concentration in the blood, rather than the hematocrit value, was determined; the two values are, however, closely correlated. FIG. 8 shows the relationship between the hemoglobin concentration and the BIS factor (without k.sub.Hct) given in Formula 12:

(35) The theoretically expected linear relationship between the BIS factor and the hemoglobin concentration can be clearly discerned.

(36) For the concentration of plasma proteins the following applies:

(37) $\begin{array}{cc}{c}_{\mathrm{protein}}=D_{\mathrm{protein}}.Math.\frac{V_{\mathrm{protein}}}{V_{\mathrm{plasma}}}=D_{\mathrm{protein}}.Math.\hspace{1em}\frac{V_{\mathrm{total}}-V_{\mathrm{RBC}}-V_{\mathrm{plasma}}}{V_{\mathrm{total}}-V_{\mathrm{RBC}}}\propto k_{\mathrm{pro}}.Math.\frac{1-{\left(R_E\right)}^{-2/3}-{\left(R_I\right)}^{-2/3}}{1-{\left(R_I\right)}^{-2/3}}& \mathrm{Formula}13\end{array}$

(38) If one plots this factor against the plasma protein concentrations measured in the laboratory (FIG. 9), one again obtains a linear relationship, as theoretically expected.

(39) These two linear relationships can be used to determine, with the aid of the bioimpedance, the concentrations of plasma proteins and hemoglobin, and/or the hematocrit value, in “real time” during the dialysis.

(40) By way of an example, a particular embodiment in connection with a dialysis machine is described below with the aid of FIGS. 10a and 10b along with FIG. 11:

(41) FIG. 11 is a schematic diagram of the blood flow in a typical dialysis arrangement with an arterial blood withdrawal 21, blood pump (peristaltic roller pump) 22, heparin feed 23, arterial bubble catcher 24, dialyzer 25, venous bubble catcher 26, injection port 27 and venous blood return 28. Devices are incorporated before and after the dialyzer for determining intracellular and/or extracellular fractions of fluids in the extracorporeal blood circulation. These comprise: coupling-in electrodes 1, 2, for coupling in the measurement signal (measuring current) from the measurement signal generator 13 into the blood to be measured, and coupling-out electrodes 11, 12, for coupling out the voltage drop (measurement signal) produced in the blood by the measuring current, with the voltage drop being measured by a detecting device 14. Not shown is the evaluation device, by means of which the intracellular and/or extracellular, in particular macromolecular fractions of fluids are calculated from the voltage drops and/or phase shifts at different measurement frequencies.

(42) In order to determine intracellular and/or extracellular fractions of fluids in plasma water during the dialysis, and use these fractions expediently for monitoring, controlling and/or regulating the dialysis, the bioimpedance of the blood must be measured continuously or periodically at short time intervals. The coupled-in measurement signal is an electrical, electromagnetic and/or magnetic alternating field, which is frequency-modulated (wobbled) over a wide bandwidth (10 Hz to 10 MHz, preferably 1 kHz to 1 MHz, particularly preferably 5 kHz to 1 MHz). The detection of the electrical measurement value is carried out by means of a device for processing measurement values with high amplitude resolution and temporal resolution, using correlation with the coupled-in measurement signal, for example by means of a impedance analyzer, in order to be able to carry out the determination of the impedance according to amplitude and phase with a high resolution.

(43) The measurement should preferably be carried out without galvanic contact between electrode and blood. Thus a pair of electrodes 1, 2 for capacitive injection of a measuring current, and a pair of electrodes 11, 12 for capacitive measurement of the voltage drop, are applied to the blood tube.

(44) With the arrangement shown here it is additionally possible to detect air bubbles and blood clots in the blood tube. For this purpose, measurement for example at a frequency of 1 MHz and with a sampling rate of 30 samples per second can be carried out continuously. In this a small electrode spacing, of for example 20 mm, is advantageous.

(45) The amplitude |Z.sub.blood| and the phase angle φ.sub.blood of the blood impedance are obtained in this manner:
Z.sub.blood=|Z.sub.blood|*e.sup.iφ.sup.blood

(46) Then the moving average for the amplitude of |Z.sub.blood| and the phase angle φ.sub.blood over the last 64 values is calculated:

(47) $\begin{array}{cc}\stackrel{\_}{.Math.Z_{\mathrm{blood}}.Math.}\left(t\right)=\frac{1}{64}\underset{k=1}{\overset{64}{.Math.}}.Math.Z_{\mathrm{blood}}.Math.\left(t-k\right)\stackrel{\_}{{\phi }_{\mathrm{blood}}}\left(t\right)=\frac{1}{64}\underset{k=1}{\overset{64}{.Math.}}{\phi }_{\mathrm{blood}}\left(t-k\right)& \mathrm{Formula}15,16\end{array}$

(48) The differences between the moving average and the present measured value Δ|Z.sub.blood| and Δφ.sub.blood show abrupt changes in blood impedance on the passage of air bubbles or clots.
Δ|Z.sub.blood|=|Z.sub.blood|−|Z.sub.blood|
Δφ.sub.blood=φ.sub.blood−φ.sub.blood

(49) FIG. 12a (top) shows the difference in the amplitude, and FIG. 12b (bottom) the difference of the phase angle, in the measured blood impedance when air bubbles are injected (at the points indicated by asterisks (*)) into the blood tube through a septum. In the measurement shown here 0.9% saline solution was used instead of blood; the delivery rate of the blood pump was 600 ml/min.

(50) The injection of the air bubbles into the blood tube thereby becomes particularly apparent through a change in the phase angle.

(51) The arrangement shown here also enables the reliable detection of any hemolysis which might occur in the blood in the blood tube. In this case, when red blood cells are destroyed there is a reduction in the intracellular components of the blood, in particular in the hematocrit (Hct), which can be determined reliably by the method.

(52) At the same time there is a decrease in the intracellular hemoglobin fraction, which can preferably be determined by measurement at high frequency. This hemoglobin from the destroyed blood cells dissolves in the plasma and raises the plasma's protein content, which can likewise be determined by the method. When quotients are formed and monitored, hemolysis becomes evident through a sudden fall in the otherwise substantially constant quotients. The standard filtration of medium molecular proteins by the dialysis filter has little influence on the protein content of the plasma, since the large albumin and globulin molecules do not pass through the filter. Thus the quotient from the hemoglobin mass and the plasma protein mass remains constant in normal cases, i.e. when the dialysis treatment of the blood is correct (without damage to the erythrocytes):
const.=m.sub.Hb/m.sub.pro=(V.sub.blood*c.sub.Hb)/(V.sub.plasma*c.sub.Pro)=c.sub.Hb/((1−Hct)*c.sub.Pro

(53) When hemolysis occurs, however, there is a fall in m.sub.Hb and a simultaneous rise in m.sub.pro, resulting in a dramatic change in their quotient, as shown in FIG. 13 (abscissa values above 110).

(54) As well as measurement by means of capacitive coupling-in of the measurement current as described above, analogous contact-free measurement using inductive coupling-in of the measurement current via an exterior coil (coupling coil) is also conceivable for the person skilled in the art. Using the coil, magnetic fields of different frequencies are thereby generated outside the blood tube. As a measurement signal, the magnetic field that arises can for example be measured from outside using a GMR sensor (giant magnetoresistance sensor).

(55) Due to the magnetic field injected, eddy currents are generated in the measurement area, which counteract the injected magnetic field. At low frequencies of the magnetic field, only small eddy currents form in both the extracellular and intracellular space, because the currents cannot pass through the cell membranes. The attenuation of the injected magnetic field is therefore only slight, and the GMR sensor would measure only a slight diminution of the magnetic field. At higher frequencies of the magnetic field, the eddy currents can pass through the cell membranes, and the injected magnetic field is attenuated to a greater extent. The measurement of the magnetic field influenced by the induced eddy currents is also possible in a known manner using other magnetic field sensors (e.g. Hall sensors or receiver coils).

(56) Capacitive measurement forms a high-pass filter, and is therefore particularly suitable for higher frequencies. Thus in an alternative embodiment the low-frequency end of the measurement signal spectrum is coupled-in inductively by means of coupling coils (not shown), and the electrical measurement value thereby generated in the fluid to be measured is likewise coupled-out inductively by means of coupling coils (not shown). In contrast, the high-frequency end of the measurement signal spectrum is coupled-in capacitively, as shown in FIG. 11, by means of flat electrodes 1, 2, and the electrical measurement value thereby generated in the fluid to be measured is likewise coupled-out capacitively by means of flat electrodes 11, 12, so that in this case capacitive and inductive coupling on the excitation and measurement sides are used alongside each other, simultaneously or intermittently.

LIST OF REFERENCE SIGNS

(57) 1. coupling-in electrode 2. coupling-in electrode 3. fluid, blood 4. protein 5. blood cells 11. coupling-out electrode 12. coupling-out electrode 13. measurement signal generator 14. detecting device 15. blood tube 16. wall of the blood tube 21. arterial blood withdrawal 22. blood pump 23. heparin feed 24. arterial bubble catcher 25. dialyzer 26. venous bubble catcher 27. injection port 28. venous blood return