Layered structure of a sensor for capacitive measurement of bioelectrical signals

Abstract

The present invention relates to a layered signal measurement circuit (30; 40) for a differential voltage measurement system (1) for measuring the bioelectrical signal (S(k)) of a patient (P). The signal measurement circuit (30; 40) includes: a sensor electrode layer (3; 4) connected to a measurement amplifier circuit via sensor lines (S3; S4); and an active shielding layer (3S1; 4S1) extending on the side of the sensor electrode layer away from the patient, wherein the sensor electrode layer and the active shielding layer are electrically conductive. The signal measurement circuit (30; 40) further includes a first insulating layer (I31; I41) extending between the sensor electrode layer and the active shielding layer.

Description

Technical Field

[0001] This invention relates to a layered signal measurement circuit for a differential voltage measurement system used to measure the bioelectrical signals of a patient. In this configuration, conductive electrode layers or shielding layers are insulated from each other, particularly by insulating layers. Background Technology

[0002] Voltage measurement systems used to measure bioelectrical signals, especially differential voltage measurement systems, are used in medicine to measure electrocardiograms (ECG), electroencephalograms (EEG), or electromyograms (EMG).

[0003] Measuring cardiac activity, especially for cardiac imaging, using the aforementioned voltage measurement system is necessary to match the imaging process to the intense movements of the heart during a heartbeat. This requires the use of conventional sensors that must be securely fastened to the patient's body. A feasible method for heartbeat measurement is capacitive ECG, where the ECG signal is captured purely capacitively without direct contact between the patient and the sensor, particularly through the patient's clothing. For good signal quality of the heartbeat signal, the measured signal amplitude preferably must be large. This can be achieved through a large capacitance between the patient and the sensor. The capacitance can be directly influenced by the size of the coupling surface between the sensor and the patient. The larger the coupling surface, the larger the capacitance achieved.

[0004] Currently, the challenge lies in configuring capacitive sensors to be suitable for use in clinical settings. Capacitive sensors must meet stringent requirements regarding cleanability, sterilizability, and mechanical robustness. Furthermore, capacitive sensors must not impair the detection of medical imaging data; that is, the sensors must be X-ray transparent or MR invisible, and the triggering of the imaging data detection must utilize the detected bioelectrical signal. Moreover, in terms of sufficient measurement signal quality, capacitive sensors used for desired positioning at the patient site must be flexibly or moldably constructed and possess good triboelectric properties.

[0005] Capacitive ECG devices are known, which are layered and integrated into conductive fabrics, where conductivity is achieved, for example, through a vapor deposition process with conductive particles. However, using fabric as an integral part of the sensor element makes the cleaning process difficult. Furthermore, fabrics are opaque to X-rays, making them unsuitable for triggering arbitrary medical image data detection. Additionally, ECG devices with metallic surfaces are known; however, they also lack X-ray transmittance. Summary of the Invention

[0006] In this regard, the object of the present invention is to provide a mechanism that meets the requirements of the clinical environment in terms of watertightness, cleanability, robustness and / or imaging characteristics.

[0007] This objective is achieved by the signal measurement circuit and differential voltage measurement system according to the invention. Other particularly advantageous designs and improvements of the invention are derived from the following description, wherein features of different embodiments or variations can also be combined to form new embodiments or variations.

[0008] In a first aspect, the present invention relates to a layered signal measurement circuit for a differential voltage measurement system used to measure the bioelectrical signals of a patient. The signal measurement circuit includes a sensor electrode layer connected to a measurement amplifier circuit via sensor lines. That is, according to the invention, the sensor electrodes are configured as surface electrodes or layered electrodes. The signal measurement circuit also includes an active shielding layer extending on the side of the sensor electrode layer away from the patient. According to the invention, the active shielding is also configured as surface electrodes or layered electrodes. The sensor electrode layer and the active shielding layer are both electrically conductive. A first insulating layer, also configured as surface electrodes, extends between the sensor electrode layer and the active shielding layer, preventing short circuits between the sensor electrode layer and the active shielding layer.

[0009] The sensor circuitry is used to transmit the measurement signal detected by means of sensor electrodes to the measurement amplifier circuitry. Preferably, the measurement amplifier circuitry includes an operational amplifier, which can be configured as a so-called follower. That is, the negative input terminal (also called the inverting input terminal) of the operational amplifier is coupled to the output terminal of the operational amplifier, thereby generating a high virtual input impedance at the positive input terminal.

[0010] An active shielding layer covers the sensor electrode layer, and both layers preferably have the same base area. In an embodiment of the invention, the active shielding layer is connected to the output of the operational amplifier. The potential of the active shielding layer is controlled or settable. The potential of the active shielding layer is maintained near the potential of the sensor electrode layer, thereby preventing current flow from the sensor electrode layer to the active shielding layer. The active shielding layer deflects interference acting upon it from the environment, preventing such interference from reaching the sensor electrode layer and from coupling to the input there. The active shielding layer is characterized by a high virtual input impedance.

[0011] The sensor electrode layer, the first insulating layer, and the active shielding layer, along with other layers optionally introduced below, together constitute a sensor element positioned at or on the patient in a manner oriented relative to the patient by means of the sensor electrode layer for the purpose of measuring signal detection.

[0012] By employing a layered structure where each layer has a significantly smaller size in one spatial dimension than in the other two, the signal measurement circuitry is made exceptionally thin, allowing it to be well molded onto the patient's body.

[0013] In a second aspect, the present invention relates to a differential voltage measurement system for measuring bioelectrical signals of a patient. The voltage measurement system has at least two signal measurement circuits, each corresponding to an effective signal path, the effective signal path comprising a sensor electrode. The differential voltage measurement system can also include more than two signal measurement circuits. At least one signal measurement circuit, preferably all of the included signal measurement circuits, is configured as described above and below. In particular, the electrodes and / or shielding of at least one signal measurement circuit are constructed in layers or in a layered manner.

[0014] As mentioned earlier, the differential voltage measurement system according to the invention detects bioelectrical signals, such as those of human or animal patients. For this purpose, the differential voltage measurement system has a number of measurement lines or effective signal paths. These measurement lines or effective signal paths, for example as a single cable, connect electrodes positioned at the patient to detect signals to other components of the voltage measurement system, i.e., in particular electronic devices, for evaluating or displaying the detected bioelectrical signals, especially heartbeat signals.

[0015] Those skilled in the art are familiar with the basic operation of differential voltage measurement systems, and therefore will not elaborate further. The differential voltage measurement system can, in particular, be configured as an electrocardiogram (ECG), electroencephalogram (EEG), or electromyogram (EMG).

[0016] In one embodiment, the differential voltage measurement system may include a reference measurement circuit, which includes a reference electrode. The reference electrode, or its associated reference measurement circuit, is used to achieve potential balance between the patient and the ECG measurement device. In another embodiment, the reference measurement circuit also includes sensor lines and an operational amplifier. In yet another embodiment, the reference electrode is configured as a separate, independent sensor element.

[0017] In an implementation, the differential voltage measurement system may include a grounding circuit, preferably configured as a separate sensor element, the grounding circuit including a grounding electrode, through which the patient can at least capacitively couple with the grounding potential during signal detection.

[0018] Therefore, the differential voltage measurement system according to the invention has at least one signal measurement circuit according to the invention. Accordingly, the differential voltage measurement system according to the invention shares the advantages of the signal measurement circuit according to the invention.

[0019] In an advantageous embodiment, the signal measurement circuit further includes a conductive second shielding layer disposed on the side of the active shielding layer opposite to the sensor electrodes and separated from the active shielding layer by a second insulating layer. This passive second shielding layer covering the active shielding layer is used to intercept particularly strong electric fields that could overload the operational amplifier controlling the aforementioned active protection barrier. In these embodiments, even very strong interfering electric fields can be effectively shielded. The second shielding layer is insulated from the active shielding layer via the second insulating layer. This additional layer advantageously ensures an improvement in the detected measurement signal without significantly affecting the overall height of the signal measurement circuit.

[0020] In other embodiments, the signal measurement circuit includes a conductive ground electrode layer to place the signal measurement circuit on a ground potential. That is, the ground electrode is also configured as a surface electrode or layered electrode and is disposed on the side of the second shielding layer opposite to the sensor electrode layer. A third insulating layer separates the ground electrode layer from the passive second shielding layer. A fourth insulating layer isolates the ground electrode layer from the environment.

[0021] As mentioned earlier, the sensor electrode layer, active shielding layer, additional shielding layer, and / or ground electrode layer are advantageously constructed to be of the same size and shape. A larger sensor electrode area results in greater capacitance, thus improving the quality of the detected measurement signal. Therefore, the base area of ​​the aforementioned layers is advantageously in the range of 9 cm² to 64 cm². The basic shape of the layers can vary between circular, quadrilateral, and especially square shapes. Preferably, the aforementioned layers have a size of 24 cm², i.e., a rectangle with side dimensions of 6 cm x 4 cm, or a size of 25 cm², i.e., a square with side dimensions of 5 cm.

[0022] A particularly advantageous design for the signal measurement circuit proposes that the first, second, third, and / or fourth insulating layers each extend completely beyond the layer to be insulated. All insulating layers are advantageously constructed to be larger. This means that while the basic shapes of the first, second, third, and / or fourth insulating layers are the same as those of the sensor electrode layer, the two shielding layers, and / or the ground electrode layer, they have a larger base area. In other words, the insulating layers have a larger perimeter than the electrode layer or shielding layer. For example, in the case of a circular basic shape, the insulating layer can extend radially beyond the electrode layer and shielding layer by 2 mm. This ensures that even in the case of flexible deformation or molding of the sensor element formed through the layered structure, short circuits between the conductive electrode layer and the shielding layer are avoided.

[0023] The signal measurement circuit according to the invention is particularly advantageously configured if the layer thicknesses of the sensor electrode layer, active shielding layer, additional passive shielding layer, ground electrode layer, first insulating layer, second insulating layer, third insulating layer, and / or fourth insulating layer are in the range of 50 μm to 500 μm. All layers of the sensor element are particularly preferably configured with a height within the said range, for example, 300 μm. Ideally, this achieves a total height of the sensor element not exceeding 4 mm, typically or even less. The thinner the layer structure overall, the better the sensor element can be molded onto the patient's anatomy, or the more flexible the sensor element is.

[0024] In an advantageous embodiment, the sensor electrode layer, active shielding layer, additional shielding layer, and / or first insulating layer, second insulating layer, third insulating layer, and / or fourth insulating layer are fused or welded to one of the adjacent layers. This results in a force-fitted connection between the layers and increases the stability of the sensor element formed through the layers, additionally ensuring insulation between the conductive electrode layer and the shielding layer. These advantages are particularly effective if all layers are connected to their adjacent layers as described.

[0025] The sensor electrode layer, active shielding layer, additional shielding layer, and / or ground electrode layer are particularly advantageously formed from carbon-particle-rich plastic.

[0026] The base material, namely plastic, here ensures mechanical stability while maintaining at least partially elastic and reversibly deformable properties in the presence of the electrode layer and the shielding layer. Carbon incorporations advantageously contribute to the conductivity of the layer. The carbon particles are preferably nanoparticles. The degree of carbon incorporation filling is related to the desired conductivity and the type of carbon particles. In selecting carbon particles, the increasing influence of mechanical material properties with increasing filling degree is particularly important. In particular, sufficient conductivity can be achieved with very low filling degrees of only a few volume percentages by using carbon nanotubes (CNTs). According to the invention, a maximum of 50% carbon should be incorporated.

[0027] Furthermore, in other embodiments, the basic materials for the electrode layer and the shielding layer can be selected according to the desired frictional performance. Materials with frictional properties matching those of cotton or pulp can be used in particular, as these fabrics are most commonly used as patient clothing and thus serve as the support surface for the materials forming the sensor electrode layer.

[0028] Examples of basic materials include polyurethane (PU) or polyvinyl chloride (PVC).

[0029] The sensor electrode layer is particularly advantageously configured such that it has a surface resistance in the range of 10 kOhm to 100 kOhm. Here, the stated and all subsequent resistance specifications correspond to the requirements of DIN EN 61340-2-3 (VDE 0300-2-3), Electrode 2-3: Surface Resistance of Electrodes and Electrodes with Specified Surface Resistance, for Electrodes with Specified Surface Resistance (IEC 61340-2-3:2016). The desired surface resistance is preferably achieved by the corresponding degree of carbon particle filling in the base material.

[0030] In another preferred embodiment of the signal measurement circuit, the first, second, third, and / or fourth insulating layers are also made of plastic. Preferably, all insulating layers are made of plastic. It is particularly advantageous for an article comprising multiple identical components if all layers of the sensor element are made of the same base material. In particular, it is also possible to achieve the same frictional performance for all included layers. Therefore, particularly uniform deformability is also obtained. Of course, different base materials can be selected for individual insulating layers or all insulating layers, such as the fourth insulating layer, especially on the outside, wherein at least similar frictional characteristics should be guaranteed.

[0031] Even at 100% humidity and temperatures up to 40°C, the base material of a single insulating layer or all insulating layers should maintain good or sufficient insulating properties. These environmental conditions arise when a patient comes into contact with the sensor element for an extended period. That is, in a preferred embodiment, the base material of the insulating layer is hydrophobic, meaning it is water-repellent and absorbs only minimal liquid, thus hardly altering its insulating properties.

[0032] Therefore, the insulating layer is particularly preferably formed of aliphatic or aromatic polyurethane (PU) or polyvinyl chloride (PVC).

[0033] By forming all the layers of the sensor element with a thin, flexible, or layered material, high patient comfort can be achieved because the sensor element can take on any shape corresponding to the individualized patient's anatomy, is compliant, lightweight, and virtually wear-free.

[0034] In an embodiment of the signal measurement circuit, the first insulating layer and / or the second insulating layer, i.e., the insulating layer between the sensor electrode layer and the passive first shielding layer, and the insulating layer between these two shielding layers, are configured such that they have a volume resistivity in the range of 50 MOhm to 50 GOhm. Preferably, both insulating layers have this volume resistivity. The first insulating layer between the sensor electrode layer and the active shielding layer functions as a voltage divider for the measured bioelectrical signal.

[0035] In other embodiments of the signal measurement circuit, the third insulating layer is configured such that it has a volume resistivity in the range of 1 GOhm to 100 GOhm.

[0036] The third insulating layer between the passive shielding layer and the grounding electrode layer positively suppresses the coupling input of interfering external electric fields.

[0037] From the perspective of electrical safety, in an embodiment of the present invention, it can be proposed that the third insulating layer of the insulating grounding electrode layer can withstand a breakdown strength of 1 minute under conditions such as 4kV AC (alternating current).

[0038] In a particularly preferred embodiment of the signal measurement circuit, the sensor electrode layer has a section that is surrounded on both sides by a first insulating layer and an active shielding layer, thus constituting a segment of the sensor circuit. In other words, the sensor electrode layer itself constitutes a segment of the sensor circuit. This segment is configured as a flat conductor element. For shielding and insulation, the segment of the sensor electrode layer is covered on both sides with a first insulating layer and an active shielding layer. The segment of the sensor circuit thus formed has mechanical properties very similar to those of a sensor element formed by different layers, and is therefore flexible, deformable, and especially flat. The segment of the sensor circuit thus formed is advantageously directly connected to the sensor element, i.e., the segment of the sensor electrode layer and the remaining sensor electrode layers belonging to the sensor element are formed in one piece. That is, the sensor circuit segment is configured with the same layers as the sensor element, wherein the materials used for each layer correspond to the configuration of the sensor element. In embodiments, the segment of the sensor circuit thus formed can transition to another circuit shape classically configured.

[0039] In an improved embodiment of the invention, the sensor circuit segment has a length ranging from 20 cm to 200 cm and a width ranging from 2 cm to 6 cm. Thus, particularly, the sensor circuit segment adjacent to the sensor element to be positioned at the patient is configured to be unaffected by interference signals and / or robust relative to them.

[0040] The signal measurement circuit according to the present invention has the following advantages:

[0041] - Capacitive ECG sensors can be manufactured cost-effectively through a structure made of plastic film, especially when the same basic material is used for all layers.

[0042] - By appropriately selecting the base materials of different layers, the signal measurement circuit is stable relative to changes in air humidity and the effects of liquids.

[0043] - The signal measurement circuit can be cleaned and disinfected particularly thoroughly and easily through the outer side of the sensor element, which is formed by a thin film.

[0044] - In any case, the materials used are X-ray resistant and do not impede X-ray imaging performed in parallel with ECG detection.

[0045] - By appropriately selecting the base material of the layer, the signal measurement circuit is particularly flexible and can be molded onto the patient's anatomy, and is also insensitive to motion. Attached Figure Description

[0046] The invention will be described in detail again below with reference to the accompanying drawings and embodiments. Here, in the different drawings, the same parts are given the same reference numerals. The drawings are generally not to scale. The drawings show:

[0047] Figure 1 A view of the differential voltage measurement system installed at the patient location in the first embodiment is shown.

[0048] Figure 2 Views are shown of differential voltage measurement systems including signal measurement circuits according to the invention, respectively, in one embodiment of the invention.

[0049] Figure 3 A view of a signal measurement circuit in another embodiment is shown. Detailed Implementation

[0050] In the accompanying drawings, ECG measurement system 1 is exemplarily used as a starting point for differential voltage measurement system 1 to measure the bioelectrical signal S(k), which is the ECG signal S(k) in this case. However, the invention is not limited thereto.

[0051] Figure 1 A view is shown of a differential voltage measurement system 1 in the form of an ECG measurement system 1 disposed at the patient P in the first embodiment. The voltage measurement system 1 includes an ECG device 17 with its electrical terminals and electrodes connected to the electrical terminals via cables K, each including a sensor line, to measure the ECG signal S(k) at the patient P.

[0052] To measure the ECG signal S(k), according to the present invention, at least one first sensor electrode and a second sensor electrode are formed in layers, respectively, and the first and second sensor electrodes are disposed at, on, or below the patient P. Currently, the sensor electrodes are disposed on different sensor elements 13, 14. The electrodes are connected via signal measurement cables K, using their sensor lines S3, S4 (see...). Figure 2 The first and second electrodes are connected to the ECG device 17 via terminals 25a and 25b, which are typically plug-in connectors. Here, the first and second electrodes, together with the signal measurement cable K, form part of a signal detection unit, by means of which the ECG signal S(k) can be detected.

[0053] The third electrode 5 serves as a reference electrode to achieve potential balance between the patient P and the ECG device 17. The third electrode 5 is here located near or on the right leg of the patient P via another separate sensor element 15 (“Right-Leg-Drive” or “RLD”).

[0054] Furthermore, multiple additional contacts for other output lines (potential measurement) can be placed at the patient P via other terminals (not shown) on the ECG device 17 and used to generate appropriate signals. Additionally, the sensor element 1a can have other sensor electrodes (not shown here).

[0055] Voltage potentials UECG34, UECG45, and UECG35 are formed between the electrodes for measuring the ECG signal S(k).

[0056] The directly measured ECG signal S(k) is displayed on the user interface of ECG device 17.

[0057] During ECG measurement, patient P is at least capacitively coupled to ground potential E via a grounding circuit including ground electrode 6, which is also configured as a separate sensor element 16.

[0058] The signal measurement cable K or corresponding sensor lines S3, S4 leading from the first sensor electrode layer 3 and the second sensor electrode layer 4 to the ECG device 17 are part of the effective signal paths 6a, 6b. The signal measurement cable K leading from the reference electrode 5 to the ECG device 17 in this case corresponds to a part of the third effective signal path 7N. The third effective signal path 7N transmits interference signals that have been coupled into the patient P and the electrodes.

[0059] Cable K has a shield S, which is schematically shown here as a dashed column surrounding all valid signal paths 6a, 6b, 7N. However, unlike what is shown here, shield S does not collectively surround all cables K, but rather shields each cable K individually. Terminals 25a, 25b, 25c preferably have poles integrated for shield S. These poles then converge onto a common shield terminal 25d. Shield S is currently configured for each cable K as a carbon-rich plastic layer or film surrounding the sensor circuitry, extending along the entire sensor circuitry S3, S4 down to below the layered sensor electrodes.

[0060] In addition, such as in Figure 1 As shown, ECG device 17 can have an external interface to provide terminals for, for example, printers, storage devices, and / or even networks. ECG device 17 also has signal measurement circuitry 30 according to an embodiment of the invention associated with the corresponding terminals 25a, 25b (see, for example, [link to relevant documentation]). Figure 2 ).

[0061] Figure 2 A view of a differential voltage measurement system 1 according to another embodiment of the present invention is shown, the differential voltage measurement system 1 including two signal measurement circuits 30 according to embodiments of the present invention. The two signal measurement circuits 30 have the same construction, therefore corresponding components of the signal measurement circuits 30 are shown only once for overview purposes.

[0062] The arrangement of the sensor electrodes is illustrated here in the form of a substantially capacitive ECG measurement circuit. Because the sensor electrode layer is electrically conductive, ohmic coupling can also be implemented in parallel with capacitive coupling. The patient P and the sensor electrodes are spatially close to each other; specifically, the patient is positioned above the sensor electrodes, which are components of sensor element 13.

[0063] The sensor electrode is configured as a thin-film conductive surface electrode layer 3 and, in this embodiment, has a basic circular shape with a base area of ​​25 cm². The sensor electrode layer 3 is oriented toward the patient P.

[0064] The sensor element 13, which contacts the patient P, also includes an active first shielding layer 3S1 and a passive second shielding layer 3S2, both electrically conductive, disposed on the side of the sensor electrode layer 3 facing away from the patient P. The sensor element 13 also includes a grounding electrode layer 3E, by means of which the signal measurement circuit 30 according to the invention can be connected to a ground potential. The listed conductive electrode layers 3, 3E and shielding layers 3S1, 3S2 are separated from each other or from the environment by electrically insulating insulating layers I31, I32, I33 and I34. Specifically, the fourth insulating layer I34 forms the outer side of the sensor element 13, which is disposed on the side of the sensor electrode layer 3 facing away from the patient P. Insulating layers I31, I32, I33 and I34 also have a basic circular shape, which has a diameter 5 mm larger than that of electrode layers 3, 3E or shielding layers 3S1, 3S2, so as to reliably insulate the remaining layers and prevent short circuits even if the sensor element is deformed, especially bent.

[0065] With the current single-layer thickness of 50 μm, sensor element 13 has a total height of 0.8 mm, which is advantageously thin because it causes less interference to the patient and can be easily deformed. Currently, the single layer is made of the same base material, namely polyurethane, in which carbon particles are mixed into the conductive electrode layers 3, 3E or shielding layers 3S1, 3S2 to achieve the desired conductivity.

[0066] In the embodiment described, the sensor electrode layer 3 is configured with a corresponding degree of carbon filling such that it has a surface resistance in the range of 100 kOhm. The first insulating layer and the second insulating layers I31 and I32 are configured such that they have a volume resistivity of 5 GOhm. Conversely, the third insulating layer I33 is configured such that it has a volume resistivity of 20 GOhm.

[0067] Sensor element 13 is coupled to the measurement amplifier circuit via sensor line S3. The eight different layers currently integrated into sensor element 13 transition to sensor line S3 at the line input on the sensor element side and continue to guide within sensor line S3 (where, for overview purposes, in sensor line S3 and further in…). Figure 2 The right part only shows conductive layers 3, 3S1, 3S2, and 3E, but not the insulating layer that also continues there. Sensor electrode layer 3 is disposed in the center of sensor line S3 as a core conductor element, which extends around the core conductor through the remaining layers that are configured in a tubular shape consistent with sensor element 13. That is, sensor line S3 essentially depicts the layer structure of sensor element 13.

[0068] The construction of the signal measurement circuit 30 is described in detail below. Patient P wears clothing that is non-conductive under normal conditions. Therefore, the sensor electrode 3 can be capacitively coupled to patient P.

[0069] The measurement amplifier circuit, including the operational amplifier 27 extending from the sensor line S3, is surrounded by an active protection barrier 25 forming an active shielding layer 3S1 at the sensor element 13 and a passive protection barrier S forming a passive shielding layer 3S2 at the sensor element 13. The operational amplifier 27 is configured as a so-called tracker. That is, the negative input terminal 27a of the operational amplifier 27 is coupled to the output terminal 28 of the operational amplifier 27. In this way, a high virtual input impedance is achieved for the operational amplifier 27 at the positive input terminal 27b. This advantageously results in almost no current flowing between the sensor electrode layer 3 and the active shielding layer 3S1 or the active protection barrier 25 due to the voltage adjustment between the output terminal 28 and the positive input terminal 27b. Furthermore, the positive input terminal 27b of the operational amplifier 27 is maintained at an electrical bias voltage by means of a resistor 26 that switches towards the measurement device ground (also called the "measurement ground"). This allows the positive, high-impedance input terminal 27b to be placed at the desired measurement potential. In this way, the DC component can be suppressed, especially during the main capacitive coupling.

[0070] The measurement signal from the sensor electrode layer 3 is coupled into the high-impedance input terminal 27b.

[0071] The active protection barrier 25 and the passive protection barrier S, or the active shielding layer 3S1 and the passive shielding layer 3S2, completely surround the measurement amplifier circuit or completely isolate the sensor electrode layer 3 from the environment in order to achieve effective shielding.

[0072] The passive protection barrier S is also connected to the equipment grounding terminal 31.

[0073] These two protective barriers 25 and S are surrounded by a ground layer 3E over the entire signal measurement circuit 30, which is coupled to a ground potential E. The ground layer 3E, or the ground barrier ES forming the ground layer 3E at the sensor element 13, advantageously leads to further improvements in the extraction of interference signals via the ground contact, improved shielding against external electromagnetic interference fields, and improved extraction of electrostatic charges.

[0074] Another electrode, which forms the ground electrode 6, is also disposed in a separate sensor element 16 shown here, to couple the patient P at least capacitively and / or ohmically to the ground potential E.

[0075] The other electrode of the reference electrode 5 or the associated measurement circuit 36 ​​is used in another sensor element 15 for potential derivation, for example, as a so-called drive neutral electrode (DNE).

[0076] Reference electrode 5 and ground electrode 6 are only schematically shown. Figure 2 The left portion is depicted. Actual size relationships and electrode shapes may differ from the depicted form in practice.

[0077] The differential voltage measurement system 1 may optionally include switching devices in the form of a switch matrix 33. In the presence of multiple sensor electrodes, the switching devices are used, for example, to select which of the sensor electrodes is used for further signal processing based on the patient's anatomy.

[0078] The differential voltage measurement system 1 can also be a signal processing device in the form of a signal processing box 34. This signal processing device is configured to perform preprocessing on the detected measurement signal to remove interference components. The signal processing device 34 can be configured to perform standard processing using frequency-based filters, such as bandpass or bandstop filters, but can also perform extended interference suppression, for example, as described in German patent application DE 102019203627A.

[0079] Furthermore, the differential voltage measurement system 1 may include a triggering device 35. The triggering device 35 constitutes a control signal for identifying the patient P's heartbeat or heart rhythm and generating a control signal that includes trigger or start time information for the medical imaging device. Based on the control signal from the triggering device 35, the imaging device calculates the time for image data detection.

[0080] Figure 3 A detailed view of the signal measurement circuit 40 according to the invention in another embodiment is shown. Here, the layered structure of the sensor element 14 according to the invention is shown again in particular.

[0081] The layered signal measurement circuit for the differential voltage measurement system 1 shown herein includes a sensor electrode layer 4 connected via sensor line S4 to a measurement amplifier circuit (not shown), the differential voltage measurement system 1 being used to measure the bioelectrical signal of patient P. The sensor electrode layer is oriented towards patient P in the region of sensor element 14, and the measurement signal S(k) is capacitively coupled into the layer.

[0082] The sensor element 14 also includes an active shielding layer 4S1 extending on the side of the sensor electrode layer 4 facing away from the patient P. The active shielding layer 4S1 is used to shield the sensor electrode layer 4 from electromagnetic interference fields from the environment. Both the sensor electrode layer 4 and the active shielding layer 4S1 are conductive, thus forming a thin-film surface electrode.

[0083] In the embodiment described, the sensor element includes at least one non-conductive insulating layer I41 extending between the sensor electrode layer and the active shielding layer, and separating the two conductive layers 4, 4S1 from each other. Another non-conductive insulating layer (not shown) can be provided on the side of the active shielding layer 4S1 facing away from the patient P, forming the outer side of the sensor element 14 and outwardly defining the sensor element 14.

[0084] Currently, the sensor electrode layer 4 and the active shielding layer 4S1 have a rectangular cross-section, which has a size of 30 cm² with a side dimension of 6 cm by 5 cm. The rectangular shape of the conductive layers 4 and 4S1 also presupposes the basic shape of the sensor element 14.

[0085] At least one insulating layer I41 and (if present) an additional outer insulating layer also extend completely beyond the layers 4, 4S1 to be insulated, i.e., have a larger bottom area, wherein the side dimensions of the insulating layers are 6.4 cm and 5.4 cm, respectively, in order to avoid short circuits between layers 4, 4S1 that are conducted when the signal measurement circuit 40 moves or deforms.

[0086] The sensor element 14 currently has at least three layers 4, 141, and 4S1, each with a layer thickness of 50 μm. Therefore, (in the presence of another outer insulating layer) the overall thickness of the sensor element is achieved to a maximum of 0.15 mm to 0.5 mm (which is advantageously thin), resulting in high flexibility of the signal measurement circuit 40 without reducing patient comfort.

[0087] The sensor electrode layer 4 and the active shielding layer 4S1 are formed of a carbon-particle-rich plastic film, here PVC, to achieve the desired conductive properties. In this case, the fill level is set such that the sensor electrode layer 4 has a surface resistance of 50 kOhm.

[0088] The insulating layer, especially insulating layer I41, is currently also made of a plastic film, but without carbon inclusions. Insulating layer I41 is configured to have a volume resistivity of 5 GΩ.

[0089] To prevent damage to the sensor element 14 during molding or during the transport or storage of the signal measurement circuit 40, the sensor electrode layer 4, the active shielding layer 4S1, and the insulating layer are soldered together. This connection method operates across the entire surface of the layers, making it particularly stable and easily implemented for various plastics.

[0090] Using plastic as the base material for the layered structure of sensor element 14 is particularly advantageous because plastic provides very good mechanical durability under conditions of high flexibility.

[0091] In this embodiment of the signal measurement circuit 40, the sensor electrode layer 4 has a segment SEG, which is surrounded on both sides by an insulating layer I41 and an active shield S41, thus forming segment A of the sensor line S4. That is, segment A is formed in the form of a flat conductor structure, which has a length of 30 cm when the width is 3 cm. With the above-mentioned layer thickness of 50 μm, the height of segment A is a total of 250 μm to 400 μm (or higher), which is always very thin. Because at least the sensor electrode layer 4 extends directly into segment A via the segment SEG, the coupling between the sensor element 14 and the sensor line S4 is preferably stable and not easily interfered with. Because the patient's body often extends into the area of ​​the sensor line S4 during the molding of the sensor element 14, patient comfort and interference-free signal transmission can be ensured by means of segment A. After segment A, the sensor line S4 transitions into a typical core conductor element, as shown in reference Figure 2 As described.

[0092] Finally, it should be reiterated that the device described in detail above is merely an embodiment, and these embodiments can be modified in different ways by those skilled in the art without departing from the scope of the invention. Therefore, the differential voltage measurement system can be not only an ECG device, but also other medical devices for detecting bioelectrical signals such as EEG, EMG, etc. Furthermore, the use of the indefinite articles "a" or "one" does not preclude the related features from existing multiple times.

[0093] While not explicitly stated, but meaningful and relevant to the invention, various embodiments, their sub-aspects, or features can be combined or interchanged with each other without departing from the scope of the invention. Where applicable, the advantages of the invention described with reference to one embodiment also apply to other embodiments, even if not explicitly mentioned.

Claims

1. A layered signal measurement circuit (30; 40) for a differential voltage measurement system (1) for measuring the bioelectrical signal (S(k)) of a patient (P), the signal measurement circuit (30; 40) comprising: - Sensor electrode layer (3;) connected to the measurement amplifier circuit via sensor lines (S3; S4) 4), and - An active shielding layer (3S1; 4S1) extending on the side of the sensor electrode layer opposite to the patient. The sensor electrode layer and the active shielding layer are both electrically conductive. The signal measurement circuit (30; 40) further includes: - A first insulating layer (I31; I41) extending between the sensor electrode layer and the active shielding layer. The signal measurement circuit also includes a conductive second shielding layer (3S2), which is disposed on the side of the active shielding layer away from the sensor electrode and is separated from the active shielding layer by a second insulating layer (I32).

2. The signal measurement circuit according to claim 1, the signal measurement circuit further includes a conductive ground electrode layer (3E), the ground electrode layer (3E) is disposed on the side of the second shielding layer away from the sensor electrode and is separated from the second shielding layer by a third insulating layer (I33) and separated from the environment by a fourth insulating layer (I34).

3. The signal measurement circuit according to claim 2, wherein the sensor electrode layer, the active shielding layer, the additional shielding layer and / or the ground electrode layer have a 9cm... 2 Up to 64cm 2 The base area within the range.

4. The signal measurement circuit according to claim 2 or 3, wherein the first insulating layer, the second insulating layer, the third insulating layer and / or the fourth insulating layer each completely extend beyond at least one of the layers to be insulated.

5. The signal measurement circuit according to claim 3, wherein the sensor electrode layer, the active shielding layer, the additional shielding layer, the ground electrode layer, the first insulating layer, the second insulating layer, the third insulating layer and / or the fourth insulating layer have a layer thickness in the range of 50 μm to 500 μm.

6. The signal measurement circuit according to claim 3, wherein the sensor electrode layer, the active shielding layer, the additional shielding layer, the ground electrode layer and / or the first insulating layer, the second insulating layer, the third insulating layer and / or the fourth insulating layer are respectively fused or welded to at least one of the adjacent layers.

7. The signal measurement circuit of claim 3, wherein the sensor electrode layer, the active shielding layer, the additional shielding layer and / or the ground electrode layer are formed of a plastic rich in carbon particles.

8. The signal measurement circuit according to any one of claims 1 to 3, wherein the sensor electrode layer is configured such that the sensor electrode layer has a surface resistance in the range of 10 kOhm to 100 kOhm.

9. The signal measurement circuit according to claim 2 or 3, wherein the first insulating layer, the second insulating layer, the third insulating layer and / or the fourth insulating layer are made of plastic.

10. The signal measurement circuit according to any one of claims 1 to 3, wherein the first insulating layer and the second insulating layer are configured such that the first insulating layer and the second insulating layer have a volume resistivity in the range of 50 MOhm to 50 GOhm.

11. The signal measurement circuit according to claim 2 or 3, wherein the third insulating layer is configured such that the third insulating layer has a volume resistivity in the range of 1 GOhm to 100 GOhm.

12. The signal measurement circuit (40) according to any one of claims 1 to 3, wherein the sensor electrode layer has a segment (SEG) which is surrounded on both sides by a first insulating layer (I41) and an active shielding layer (4S1) and thus constitutes a segment (A) of the sensor line (S4).

13. The signal measurement circuit (40) according to claim 12, wherein the segment of the sensor line has a length in the range of 20cm to 200cm and a width in the range of 2cm to 6cm.

14. A differential voltage measurement system (1) for measuring a bioelectrical measurement signal (S(k)) of a patient (P), the differential voltage measurement system (1) having at least two signal measurement circuits (30; 40) respectively corresponding to the effective signal paths (6a, 6b) of the voltage measurement system, wherein at least one of the signal measurement circuits is configured according to any one of claims 1 to 13.

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