WO2020156631A1

WO2020156631A1 - Sensors based on dielectric silicon layers

Info

Publication number: WO2020156631A1
Application number: PCT/EP2019/052017
Authority: WO
Inventors: Klaus Eller; Manfred Amann; Carsten Bornhoevd; Jochen Dauth; Rebecca NEUWIRTH-GAUKEL
Original assignee: Wacker Chemie Ag
Priority date: 2019-01-28
Filing date: 2019-01-28
Publication date: 2020-08-06

Abstract

The sensors comprise at least one dielectric silicon layer. The silicon layer is located between at least two electrodes. The sensors according to the invention are sensitive to certain analytes, light, or electromagnetic fields and thus can detect changes in the area surrounding the sensors. The sensors according to the invention are simple to produce and have a variety of possible applications. The invention additionally relates to a measuring system comprising the sensor, to a method for producing the sensor, and to a measuring method using the sensor.

Description

Sensors based on dielectric silicone layers

The invention relates to sensors comprising at least one dielectric

Silicone layer. The silicone layer is located between at least two electrodes. The sensors according to the invention are sensitive to certain analytes, to light or to electromagnetic fields and can therefore detect changes in their environment. The sensors according to the invention are simple to manufacture and have a wide range of possible uses. Furthermore, the present invention comprises a measuring system comprising the sensor, a method for producing the sensor, and a measuring method using the sensor.

State of the art

WO 2014/105959 A1 describes compositions of silicone polymers that can be used as EAP material (electroactive polymer). Multi-layer systems with electrodes are also described which are used as transducers.

WO 2017/210795 A1 describes an application of a dielectric polymer in capacitive and resistive 2D measuring systems. Here electrodes made of ionically conductive material are used. There are pressure and volume as well

Changes in moisture or resistance are detected and evaluated. The structure here is comparable to a key matrix and is also

evaluated accordingly.

WO 2015/130550 A1 describes a sensor patch with a capacitive one

Measurement structure. Here, a porous material located between two electrodes is influenced by an analyte and this is evaluated electrically (capacitively). The analyte penetrates into the porous dielectric in gaseous form. Use on the skin, for wound monitoring, on tubes and in animals is also described. Wireless evaluation using RFID technology is also listed. WO 2017/052773 describes the use of an EAP polymer and EAP layer system as a shear, tension, torsion and pressure sensor for various

Applications such as measuring shoe insoles. A composite construction of a thin elastomer, such as silicone, with an electrode, such as graphite mixed with silicone, and a dielectric insulator, such as silicone, is explicitly mentioned herein. A degeneration determination by measuring the electrode resistances is also mentioned.

US 2017/0086709 A1 describes the use of a silicone polymer with a flexible electrode as a strain gauge. The use as a plaster on the skin is mentioned, whereby a flexible substrate is used which can consist of an elastomeric silicone film and comes into direct contact with the skin. The application form is also given in combination with a portable evaluation and data storage and communication device.

WO 2016/133841 A1 describes a sensor which can be attached to the skin and which partially penetrates the skin or the bloodstream.

Abdus Sobhan et al. describe in Bioscience, Biotechnology, and Biochemistry, 2018, VOL. 82, NO. 7, 1134-1142 a method for producing a biosensor by immobilizing antibodies for the detection of peanut allergens (Arachis hypogaea, Ara h1-Ara h 13) on single-walled carbon nano-tubes (SWCNT). By using carbon nano tubes, these are

However, sensors are very limited in their application.

The disadvantages of the devices described in the prior art are low elastic deformability, lack of permanent temperature stability. In addition, the conventional systems are usually difficult to manufacture and have no variable usability with regard to different receptors and evaluation methods.

It was therefore the object of the present invention to provide an improved sensor which exceeds the properties of conventional sensors or devices and which is easy to manufacture and vary. Surprisingly, this task is accomplished by sensors with dielectric

Silicone layers released. The advantages of the silicon-based sensors of the present invention over conventional systems are that

Silicone layers are scalable in any area, skin-friendly, permeable to gas and water vapor (gas-selective), sterilizable, biocompatible, bio-resistant, atraumatic, break-resistant, permanently flexible, stretchy and temperature-resistant. Furthermore, binding molecules of any kind can be used for the detection of

simply attach different analytes. Also a combination with actuators (valves, pumps, etc.) and a combined evaluation (impedance

spectroscopic, resistive, mechanical behavior) in one sensor is possible.

Illustrations

Figure 1 shows the sketch of a one-sided biosensor.

Figure 2 shows the sketch of a two-sided biosensor.

Figure 3 shows the model for capacitive evaluation.

Figure 4 shows the model for resistive evaluation.

Figure 5 shows the sketch of a gas sensor.

Figure 6 shows the sketch of a one-sided chemical liquid sensor.

FIG. 7 shows the test setup of a chemical liquid sensor.

Description of the invention

The invention relates to a sensor comprising:

(a) at least two electrodes

(b) at least one dielectric disposed between the electrodes

Silicon layer, which is designed so that

an analyte can be adsorbed or absorbed on one or both sides of the silicone layer,

at least one further layer, which can adsorb or absorb an analyte, is arranged on one or both sides of the silicone layer,

embedded in the silicone layer or between several silicone layers is a material which can absorb an analyte or can react chemically with an analyte, - A material is embedded in the silicone layer or between several silicone layers, which can change its chemical and / or physical properties when exposed to light, or

- A material is embedded in the silicone layer (3) or between several silicone layers (3), which can change its chemical and / or physical properties when exposed to an electromagnetic field.

The common manufacturing and process steps for generating

Electrode masses and application methods are known in the prior art.

Different materials and systems can be used as electrode material.

Solid electrode base materials are produced by mixing conductive substances such as furnace carbon black, acetylene black, graphite, graphene, carbon fibers, carbon nanofibers, metal particles and metal fibers into silicone elastomers. A silicone elastomer base material comparable to the dielectric layers described below is preferably used.

Alternatively, a solid electrode can consist of electrically conductive components (metals, etc.) applied directly to the dielectric layers. The application can be in the form of a spray, screen printed, in one

Gas phase process or in the form of a plasma coating. Simple processes such as brushing or rolling metal-filled inks or carbon black emulsions can also be used.

Electrode liquids are electrically conductive liquids, for example salts dissolved in solvents or ionic liquids. Common liquids are for example at the company: loLiTec lonic Liquids Technologies GmbH,

Germany available. Common contact electrodes for the electrode fluids are known in the prior art and are commercially available.

The dielectric silicone layer preferably consists of at least one silicone elastomer. The silicone layer preferably has a relative permittivity BR (18 ° C., 50 Hz) of

1, 8 or more, particularly preferably from 2.0 or more, particularly preferably 2.8 or more, preferably from 3.0 or more to 30.0 or less. The measurements of the dielectric material parameters can be carried out by a system from NOVOCONTROL Technologies GmbH & Co. KG, Germany.

A Novocontrol spectrometer from the Alpha-A series with a ZGS Active Sample Cell can be used for temperature-dependent measurements in the range of -200 to 400 ° C. The preferred range is in the range of +20 - 150 ° C. Test specimens are measured in the 200 - 1000 pm layer thickness range with an applied voltage up to 1V RMS in the range 0.1 Hz to 10 MHz, preferably in the range greater than 10 kHz.

In principle, all silicone elastomer compositions known in the prior art can be used as base materials for the silicone layer. For example, addition-crosslinking, peroxidically crosslinking,

condensation crosslinking or radiation crosslinking

Silicone elastomer compositions can be used. Are preferred

peroxidic or addition-crosslinking compositions. Addition crosslinking compositions are particularly preferred.

Suitable silicone elastomer compositions are available, for example, under the names ELASTOSIL ^® and SILPURAN ^® from WACKER Chemie AG, Germany. Other suitable silicone compositions are in

WO 2014/090506 A1, WO 2015/11391 1 A1 and WO 2018/177523 A1.

The silicone elastomer compositions can be formulated in one or two components. The silicone elastomer compositions are crosslinked by supplying heat, UV light and / or moisture. Are suitable

for example the following silicone elastomer compositions: HTV

(addition crosslinking), HTV (radiation crosslinking), LSR, RTV 2 (addition crosslinking), RTV 2 (condensation crosslinking), RTV 1, TPSE (thermoplastic

Silicone elastomer), thiol-ene and cyanoacetamide crosslinking systems.

In the simplest case, the addition-crosslinking silicone compositions contain (A) at least one linear compound which has residues with aliphatic carbon-carbon multiple bonds,

(B) at least one linear organopolysiloxane compound with Si-bonded hydrogen atoms,

or instead of (A) and (B)

(C) at least one linear organopolysiloxane compound which has SiC-bonded radicals with aliphatic carbon-carbon multiple bonds and Si-bonded hydrogen atoms, and

(D) at least one hydrosilylation catalyst.

In a special embodiment, the

Silicone compositions around silicone elastomer compositions with fluorinated side groups, as described for example in WO 2018/177523 A1. In this embodiment, component (A), (B) and / or (C) preferably contain at least 2.5 mol%, particularly preferably at least 5 mol%, fluorinated

Side groups, such as 3,3,3-trifluoropropylmethylsiloxy and / or

Bis (3,3,3-trifluoropropyl) siloxy groups

The addition-crosslinking silicone compositions can be

One-component silicone compositions as well as two-component silicone compositions.

In the case of two-component silicone compositions, the two components of the addition-crosslinking silicone compositions can contain all constituents in any combination, generally with the proviso that one component does not simultaneously contain siloxanes with an aliphatic multiple bond, siloxanes with Si-bonded hydrogen and a catalyst, that is to say essentially not simultaneously the constituents (A), (B) and (D) or (C) and (D) contains.

The used in the addition-crosslinking silicone compositions

As is known, compounds (A) and (B) or (C) are chosen so that crosslinking is possible. For example, compound (A) has at least two aliphatic unsaturated radicals and (B) has at least three Si-bonded radicals

Hydrogen atoms, or compound (A) has at least three aliphatic unsaturated residues and siloxane (B) at least two Si-bonded

Hydrogen atoms, or instead of compound (A) and (B), siloxane (C) is used, which contains aliphatic unsaturated radicals and Si-bonded

Has hydrogen atoms in the above ratios. Mixtures of (A) and (B) and (C) with the abovementioned ratios of aliphatic unsaturated radicals and Si-bonded hydrogen atoms are also possible.

The silicone composition usually contains 30-95% by weight, preferably 30-80% by weight and particularly preferably 40-70% by weight (A), based on the

Total mass of the silicone composition.

The silicone composition usually contains 0.1 to 60% by weight, preferably 0.5 to 50% by weight and particularly preferably 1 to 30% by weight (B), based on the

Total mass of the silicone composition.

If the silicone composition contains component (C), 30-95% by weight, preferably 30-80% by weight, particularly preferably 40-70% by weight (C) are usually present in the formulation, based on the total mass of the

Silicone composition.

The amount of component (D) can be between 0.1 and 1000 parts per million (ppm), 0.5 and 100 ppm or 1 and 25 ppm of the platinum group metal, depending on the total weight of the components.

The addition-crosslinking silicone compositions can optionally contain all the other additives which have also hitherto been used for the division of the

addition-crosslinkable compositions were used. These additives can contain up to 70% by weight, preferably 0.0001 to 40% by weight. These additives can e.g. inactive fillers, resinous

Polyorganosiloxanes, which are different from the siloxanes (A), (B) and (C), reinforcing and non-reinforcing fillers, fungicides, fragrances, theological additives, corrosion inhibitors, oxidation inhibitors, light stabilizers, flame retardants and agents for influencing the electrical Properties, dispersing agents, solvents, adhesion promoters, pigments, Dyes, plasticizers, organic polymers, heat stabilizers, etc. These include additives such as quartz powder, diatomaceous earth, clays, chalk, lithopones, carbon blacks, graphite, metal oxides, metal carbonates, sulfates, metal salts of carboxylic acids, metal dusts, fibers, such as glass fibers, plastic fibers, plastic powder,

Metal dusts, dyes, pigments, etc.

Examples of reinforcing fillers which can be used as a component in addition-crosslinking silicone compositions are pyrogenic or precipitated silicas with BET surface areas of at least 50 m ² / g. The above

Silicic acid fillers can have a hydrophilic character or can be rendered hydrophobic by known processes. The content of the crosslinkable

Silicone compositions of active reinforcing filler are in the range of 0 to 70% by weight, preferably 0 to 50% by weight, based on the total mass of the silicone composition.

The amounts of all components present in the composition are chosen so that they do not exceed a total of 100 percent by weight, based on the total mass of the silicone composition.

The silicone layer can be produced by casting, calendering, extrusion of the silicone compositions or an emulsion or solution of these silicone compositions, laser transfer printing, 3D printing, screen printing and other processes known in the art. Suitable methods are, for example, in

WO 2014/090506 A1, WO 2015/113911 A1 and WO 2016/071241 A1 are described.

The silicone layer is preferably one or more

Silicone films or a silicone multilayer composite with two or more layers of hardened silicone. The production of thin silicone films is described, for example, in WO 2014/090506 A1. The production of thin multilayer silicone composites is described, for example, in WO 2015/1 1391 1 A1.

The silicone films or the individual layers of the silicone multilayer composites preferably have a film thickness of 0.1 to 400 pm, particularly preferably from 1 to 200 pm, very particularly preferably from 1 to 150 pm and in particular preferably from 2 to 100 gm and in each case a thickness accuracy of ± 5% measured over an area of 200 cm ² ; preferably a thickness accuracy of ± 3% in each case.

The absolute layer thickness can be determined using a SEM analysis

(Scanning electron microscope analysis) and is carried out with a cryosection. The surface quality and roughness are determined, for example, with the aid of a confocal microscope.

In order to bring about an electrically measurable change in the silicone layer, various biological, physical changes or chemical processes on the surface of the silicone layer are conceivable. They range from training or

Dissolution of an ionic or covalent bond, a complex or

Adduct formation or dissolution, through charge transfer interactions, via the physisorption of materials, the detachment of double layers (sacrificial layer), the release of molecules or aggregates bound on the surface up to swelling phenomena on or in the silicone film. Also temperature or

light-induced processes on the surface can be measured capacitively.

Such interactions often require functionalization of the silicone layer. This functionalization can be achieved in volume through functional silanes, siloxanes, silicone resins, crosslinking agents, fillers and blends with organic

Polymers are made. Another possibility is to treat the surface of the silicone layer with functional silanes, siloxanes, silicone resins, particles and organic polymers, e.g. Polyurethanes, poly (meth) acrylates and polyethers.

The following functions are suitable, for example, for immobilizing a binding molecule in and on the silicone layer: non-reactive groups, for example alkyl, aryl, alkenyl, (EO) _n - (PO) m, etc., which enable better phyisorption of the binding molecule. Reactive groups, e.g. hydroxy, amino, mercapto, formyl,

Carboxy, anhydride, epoxy, azido, isocyanato, vinyl ether, vinyl ester, acetal, diazo compounds enable the binding molecule to be covalently bonded to the silicone surface. “Binding molecules” are molecules that are bound sufficiently firmly (covalently or non-covalently) to the surface of the silicone layer and bind specifically to the desired analyte. The binding molecules can come from different classes of substances. They are preferably macromolecules such as proteins or nucleic acids.

The binding molecules are preferably immobilized by chemisorption or physisorption on the surface of the silicone layer or on the surface of an adhesion layer which is arranged on the silicone layer

Binding molecules can be proteins or parts (fragments) of proteins. Analytes can be small molecules, antibodies, proteins and protein fragments (peptides), hormones, protein complexes, small molecules, viruses / phages, whole cells, cell organelles, artificial vesicular systems or DNA / RNA.

Antibodies from a wide variety of vertebrates can be used as binding molecules. Alternatively, the antibodies come from (artificial) gene banks. The antibodies can also be produced recombinantly with microorganisms. Parts (fragments) of the antibodies are sufficient for the activity as a binding molecule (Fab, Fv, VHH, FC). Analytes can contain the corresponding antigens, other antibodies,

Proteins and protein fragments (peptides), hormones, protein complexes,

low-molecular substances, viruses / phages, whole cells, cell organelles, artificial vesicular systems or DNA / RNA.

Proteins not naturally occurring with antibody-like properties can also be used as binding molecules (anticalins, affibodies, etc.). The binding molecules can be directly on the silicone layer or in the form of

Surface structures of any kind of cells, microbes or viruses can be indirectly bound to the film.

The silicone layer equipped with binding molecules can be used to determine many analytes. The primary measuring principle is the change in the capacity of the silicone layer or the layer system through the specific binding of an analyte. Conversely, the capacity can also be changed in that a bound component is changed or split off by an analyte.

An example of such a system is the cleavage of a substrate bonded to the silicone layer, e.g. a peptide by a suitable enzyme activity, one

Polysaccharides by hydrolase or lysis of a cell by a virus.

The removal of a bound molecule or part of a molecule generates the electrical signal.

For low molecular weight substances in particular, enzymes which bind the analyte e.g. bind as a substrate can be used. For this purpose, both the naturally occurring active enzymes can be used as well as inactivated enzymes that can no longer convert the substrate (analyte). The enzyme activity of the binding protein can also be determined with active enzymes. The change in the conformation of the enzyme is used for the measurement.

DNA and RNA molecules can also serve as the binding molecule. Hybridization with analyte DNA and RNS is used in particular.

Receptors can also serve as the binding molecule. Receptors are specific proteins on the cell surface or within the cell that, due to the binding of a specific ligand, lead to an activation or inactivation of a cell process. In the case of receptors on the cell surface, external signals are passed on to the cell. Intracellular receptors are mostly regulators of metabolism. Corresponding analytes are the associated ligands

(Hormones, cell-cell interaction molecules such as CD’s (Cluster of Differentiation) or low-molecular substances of metabolism.

Another class of membrane-bound proteins are transport proteins. These can specifically recognize molecules outside and / or inside the cell and selectively transport them across a membrane system. The

Membrane systems also belong to certain cell organelles. The specific binding property can be used as a binding protein in a biosensor. Certain proteins and glycoproteins are able to bind complex polysaccharides (carbohydrate structures). Most of these are used to identify certain polysaccharides (eg lipopolysaccharides, endotoxins) on cell surfaces (lectins).

Simple sugars (glucose, fructose, lactose etc.) can be used as binding molecules and represent the reverse measurement method of the enzymes or

Transport proteins represent. Appropriate proteins (transporter, enzyme) bind as analyte. Polysaccharides such as Cyclodextrins can be used for the specific binding of low molecular weight compounds as well as for the binding of certain (e.g. tyrosine-rich) proteins. Other polysaccharides can be used to detect lectins.

Specific binding proteins have been identified in various organisms that are used for the selective binding of certain ligands / analytes. Their function is mostly detoxification (e.g. GSH - gluthation-S-transferase), growth advantages over other organisms (e.g. avidin / streptavidin, siderophores) or serve the specific interaction of a protein class (e.g. SNARE-mediated membrane fusion - snap-tag).

In the case of photoactivatable proteins, the absorption (absorption) of a photon only represents a bond in the figurative sense. The absorption of the photon leads to a change in conformation which can be measured in the form of a biosensor (e.g. bacteriorhodopsin).

A further layer is preferably arranged on one or both sides of the silicone layer and comprises one or more binding molecules, selected from the group comprising antibodies, antibody fragments, enzymes, nucleic acids, receptors, transport proteins, polysaccharide-binding proteins, polysaccharides, oligosaccharides and specific binding proteins.

Another way to immobilize the analyte on the silicone layer is through molecular imprinting, that is, “embossing” molecules in one

Polymer layer. This method offers a possibility to bind molecules in a polymer layer very specifically and is often used in biotechnology, biomedicine or also used for chemical processes. In the manufacturing process, the molecule to be analyzed initially serves as a template (template molecule). In the next step, monomers are docked onto the functional groups of the template. The polymerization then takes place around the template molecule. Once this has been completed, the template must be removed as completely as possible from the polymer matrix, for example by extraction. In this way, characteristic cavities are obtained in which the analyte can bind. By binding the analyte, the dielectric properties of the polymer film change and capacitive measurement is made possible.

Polymerization options include bulk polymerization, in which all components (monomer, crosslinker, initiator and template) are mixed together. Furthermore, there is the possibility that the cavities are only embossed on the surface of a polymer layer, this is mostly used for large analytes such as cells, viruses or bacteria, or the covalent polymerization, in which the cavity is built up step by step, first of all the

functional monomers bound to the specific binding sites of the template. The crosslinker is then used for networking.

Cavities in which an analyte can be specifically bound are preferably present in the surface on one or both sides of the silicone layer.

Typical areas of application for Molecularly Imprinted Polymers are in biotechnology filters, wastewater treatment and recovery of valuable components, such as active ingredients or precious metals, in the synthesis of enantiomer separation or catalysis, chromatography, biosensors and in medicine they can be used as imitators for antibodies, enzymes or Serve receptors. They are also suitable for drug delivery.

Several types of sensors can be implemented with the system according to the invention. The sensors according to the invention are preferably biological sensors (biosensor), chemical sensors (for gaseous or liquid analytes), electromagnetic or optical sensors. Biological and chemical sensors are particularly preferred. The sensor can be configured on one or two sides. With a one-sided

An electrode is arranged directly on one side of the possibly modified silicone layer and a chamber for the on the other side

Electrode fluid. In a two-sided configuration, such chambers are arranged on both sides.

In one embodiment, the sensor therefore comprises at least one chamber which is arranged between the silicone layer and an electrode and which is suitable for receiving a conductive electrode liquid. The

Electrode liquid is in direct contact with the possibly modified silicone layer as well as with the electrode (contact electrode). The sensor particularly preferably comprises two such chambers, in each case on both sides of the

modified silicone layer. In this construction, the electrodes serve as

Contact electrodes for the electrode fluid.

In addition to the contact electrodes, this embodiment preferably comprises two additional electrodes which are connected directly to the possibly modified silicone layer. This makes it possible to measure the longitudinal as well as the transverse impedance of the silicone layer or the layer system.

At least one further layer which can specifically bind an analyte is preferably arranged on the sides of the silicone layer which come into contact with the conductive electrode liquid. Preferably the sensor is bilateral, i.e. on both sides of the silicone layer there is another layer that can specifically bind an analyte.

The further layer preferably comprises one or more binding molecules selected from the group comprising antibodies, antibody fragments, enzymes, nucleic acids, receptors, transport proteins, polysaccharide binding proteins, polysaccharides, oligosaccharides and specific binding proteins. Antibodies and antibody fragments are very particularly preferred. At least one adhesion layer and at least one passivation layer is preferably arranged on one or on both sides of the silicone layer. The adhesive layer enables the immobilization of the binding molecules and preferably comprises the functions described above for immobilizing the

Binding molecules. The passivation layer prevents analytes from being bound non-specifically to the binding molecules. The passivation layer comprises

preferably at least one surfactant, particularly preferred are nonionic surfactants. Suitable nonionic surfactants are, for example

Polyalkylene glycol ethers, alkyl glucosides, alkyl polyglucosides, octylphenol ethoxylates or nonylphenol ethoxylates. Commercially available surfactants are, for example, Triton ^® (example: Triton ^® X-100 (Octoxinol 9), Triton ^® WR 1339 (Tyloxapol)), Tween ^® (example: Tween ^® 20 (Polysorbate 20)), fatty acid esters such as Span ^® (sorbitan fatty acid ester) ), Fatty alcohol polyglycol ether / fatty alcohol ethoxylates (FAEO) such as Brij ^® (example: Brij ^® 30, Brij ^® 58) or sucrose palmitate stearates.

Alternatively, the passivation layer comprises the analyte

binding-neutral proteins (example: BSA, milk proteins).

In a further embodiment, at least one of the electrodes is arranged directly on the silicone layer. Two electrodes are preferably directly on the

Silicon layer arranged.

In a further embodiment, two electrodes are arranged lengthwise on the silicone layer and two electrodes are arranged transversely on the silicone layer. This makes it possible to carry out a capacitive as well as a resistive measurement at the same time.

In a further embodiment, the sensor has two gas-permeable

Electrodes and two gas-permeable silicone foils, between which at least one layer with material that can absorb a gaseous analyte or with which a chemical reaction is arranged (gas sensor).

In a further embodiment the is on one or on both sides

Silicone layer applied at least one further layer with material that can chemically react with the analyte. In a further embodiment, at least one material is embedded in the silicone layer which can undergo a chemical reaction with an analyte.

In a further embodiment, the

Silicon layer an analyte can be directly adsorbed or absorbed. This is particularly preferably physisorption or chemisorption.

In a further embodiment the is on one or on both sides

Silicon film arranged at least one electromagnetically active layer and / or at least one conductive layer. If this electromagnetic and / or conductive layer interacts with an electromagnetic field, the capacitance of the silicone layer systems changes, for example. This embodiment represents a magnetoelectric sensor.

In a further embodiment, there is at least one photoactive layer within a translucent silicone layer or between two translucent silicone layers. If light strikes this photoactive layer, the capacity of the silicone layer systems changes. This embodiment represents an optical sensor

The invention further relates to a measuring device comprising:

- At least one sensor according to the invention, and

- At least one control and evaluation unit, which is connected to the electrodes of the sensor.

Several embodiments of the invention are described below by way of example using FIGS. 1 to 7. However, these are not to be understood as conclusive for the embodiments of the present invention.

Figure 1 (one-sided) and Figure 2 (two-sided) show the basic structure of a biosensor. A dielectric silicone layer (3) is equipped with an electrode (2) and a bioactive layer system (4,5,6). This surface has the property of specific interactions with an analyte to be able to enter. The bioactive layer system can comprise an adhesion layer (4), a passivation layer (5) and binding molecules (6).

The entire surface of the bioactive side is wetted by an electrode liquid (7). The analyte (8) is or is supplied in this liquid (7). Other biological components or impurities (10) can also be found in the liquid (7). Through the passivation layer (5) only specific bonds between the binding molecule (6) and the analytes (8) are effective. The specifically bound analyte is outlined as reference number (9).

The binding events cause a measurable change in the layer system. A distinction can be made between capacitive and resistive change. Analyte binding events influence the measurable capacitance between the electrode (2) and the electrode liquid (7), into which a contact electrode (1 1) is introduced.

Furthermore, a resistive measurement of the analyte bonds can be carried out in combination with the capacitive measurement. For this purpose, an independent control and

Evaluation unit (15) connected to two electrodes (13, 12) which are in direct contact with the bioactive layer (4, 5, 6) can be used to resistively detect analyte bonds. A change in conductivity in the electrolyte (7) can also be determined thereby or only. Geometric

Dimensions, function and designs of the two electrodes (12, 13) are known in the prior art. The use of CNT (carbon nanotubes) or other electrically conductive additives is described in several applications. An interference of CNT to obtain a conductive

Silicone polymers are suitable, for example. Temperature effects are a disadvantage of resistive measurement. These are less with capacitive measurements

problematic.

By skillful control of the sensor with the control and

Evaluation unit (14) can also have a mechanical change effect

(Deflection / piezo effect) can be achieved. The electrostatic deformations are used with a high electrical field or high voltage. A static deflection / deformation or a dynamic one can be generated with any vibration frequencies and shapes.

Mass deposits, acting forces, deformations etc. can be detected and the mechanical changes can also be used to make doses, releases or closures, for example. With oscillating excitation, a pump or mixing effect can also be generated by the sensor itself.

A chemical gas sensor is shown in FIG. A gas-sensitive material (53) is introduced between at least two dielectric silicone layers (3). The electrodes (51, 52) are gas-permeable and, for example, made of diffusion-open material such as GNT or silicon black filled with conductive carbon black (for example ELASTOSIL ^® 3162 available from WACKER Chemie AG, Germany). The control and evaluation unit (14) enables an impedance evaluation of the changes in the material (53) with regard to permittivity and loss factor. A change in the material (53) with regard to the electrical conductivity can be detected by that of the control and evaluation unit (14) and / or by the control and evaluation unit (15). Both can be measured separately or simultaneously. The measurement by the control and evaluation unit (15) is a direct measurement of the current flow through the material (53) between the electrodes (12) and (13). A diffusing gas (54) is absorbed in the material (53) or interacts with the material (53) in any way. This interaction is measured with the control and evaluation units (14) and (15).

A chemical liquid sensor is shown in FIG. A chemically reactive additive (61) is mixed into the silicone layer (3) or into a surface coating (4). A chemical reaction with the electrode liquid (7), such as, for example, removing the additive (62), changes the effective potential areas of the capacitive system. This results in a change in the impedance which is measured with the control and evaluation unit (14). In addition, independently of this, the impedance between the electrodes (12) and (13) can be measured with the evaluation (15). The change in impedance between the electrodes (12) and (13) can also be used to assess the degree of reaction. The sensor according to the invention can also be used as an optical sensor. Optical sensors are based on the effect of the change in resistance

Exposure to light. A photoactive layer is introduced through two electrodes in a transparent silicone body. An attached

The voltage source supplies an electrical current depending on the light irradiation. This is proportional to the illuminance. The structure is comparable to FIG. 5 (chemical gas sensor), an LDR (Light Dependent Resistor) layer (53) being enclosed by optically transparent layers (51, 52). Penetrating optical radiation changes the resistance between the electrodes (12, 13).

The invention further relates to a method comprising the following steps:

a) Provision of a measuring device comprising:

- At least one sensor according to the invention

- At least one control and evaluation unit, which with the electrodes of the

Sensor is connected;

b) contacting an analyte with the sensor or exposure to light on the sensor;

c) Measurement of the change in at least one property of the silicone layer as an electrical signal via the control and evaluation unit.

The property of the silicone layer is preferably the volume, the mass, the permittivity, the magnetic permeability, the polarizability (electrical permittivity) and / or the electrical resistance and the electrical charge.

The longitudinal impedance or transverse impedance of the silicone layer or the layer system is particularly preferably measured.

The electrical evaluation can be carried out using various methods and must always be coordinated in combination with the design of the sensor.

The simplest evaluation is in Figure 1 with an electrical control and

Evaluation unit (14) outlined.

Suitable procedures are: - Time-based capacitance measurement, which measure the charging and discharging times, and determines the frequency response on an oscillator circuit.

- Charge-based capacity measurement, which works by means of charge transfer from reference capacities.

- capacitance measuring bridge / measuring bridge circuits, which are similar to one

Wheatstone measuring bridge makes a comparison with known components.

In FIG. 1, the measurement by means of a charge-based and discharge-based method is highly error-prone due to an electrochemical potential that arises between the electrode liquid (7) and the bioactive layer (6, 5, 4). For this reason, a two-sided liquid sensor structure is to be preferred in the application (cf. FIG. 2, sectional sketch 20). Furthermore, due to the double amount of analyte bonds, there is a two-sided liquid electrode with a higher ratio

Useful signal to basic signal (relative signal change) advantageous.

An electrical model of an evaluation is given in FIG. 3. Contact resistance, parasitic impedances and the like are caused by the

Components R_E considered. The starting capacity of the body is outlined with the value C_S. The changes due to connections to the active layer or other interaction of the analyte with the silicone layer are simulated with the capacitance C_Rec. Here, the amount of the capacity and the loss factor (tan_delta) can vary in strength and frequency

Have changes according to the connections or interaction.

A model of the resistive evaluation is given in FIG. 4. Parasitic

Resistors are simulated again with the components R_E. A

Deposition / binding / release or other interaction of the analyte with the silicone layer causes a measurable change in the R_Rec. This can vary depending on the strength of the physical, biochemical effect

frequency-dependent change in measured impedance in magnitude and phase.

In one embodiment of the invention, an electrical voltage is applied via the control and evaluation unit in order to elastically deform the silicone layer. The maximum applied electrical voltage is preferably 5000 V or less, particularly preferably 2000 V or less, particularly preferably 1000 V or less. Depending on the dielectric layers

Layer thicknesses and applied voltage, max. Field strengths in the range of 200 V / mm or less, preferably 100 V / pm or less, particularly preferably 80 V / pm or less, occur.

The sensors described here can be used in a variety of applications. The following applications are preferred: sensor patches, measurement of the release of substances from the polymer matrix (active substance patch), mobile biosensor for travel, medical applications (for example sensor patches, diagnostics, measurement of active substance release), pharmaceutical applications (for example drug screening), biotechnology (for example fermentation control, Media analysis), environmental analysis (e.g. drinking water control,

Wastewater control, pollutant monitoring), food industry (e.g. quality control, freshness assessment, maturity measurement), chemical industry

(e.g. process control), industrial production (e.g. detection of leakage), protective equipment (e.g. gas masks, gloves), agrochemicals (e.g. pest infestation), and analytics (e.g. environmental pollution).

Examples

In the examples described below, all parts and percentages are by weight unless otherwise stated. Unless otherwise stated, the examples below are carried out at a pressure of the surrounding atmosphere, that is to say at about 1000 hPa, and at room temperature, that is to say at 25 ° C., or at a temperature which changes when the

Reactants at room temperature without additional heating or cooling. The following examples illustrate the invention without being restrictive.

In the examples, silicone films (ELASTOSIL ^® film 2030 250 / xxx,

SILPURAN ^® Film 2030 250 / xxx, WACKER Chemie AG, Germany) by customary processes for polymerizing silicones (e.g. hydrosilylation, Pt catalyzed crosslinking). The film thickness was adjusted by knife coating (20 gm or 50 gm). In the case of chemically functionalized films in which certain functional groups are arranged on the surface, suitable chemical compounds were incorporated in the process of film production or subsequently applied to the film surface. Carboxyl-functionalized film: incorporation of 10-undecenoic acid trimethylsilyl ester (UA-TMS); COOH exposure.

Epoxy-functionalized film: incorporation of allyl glycidyl ether or limonene 1, 2 epoxy. Amine-functionalized film: knife coating a reactive 10 gm thick layer containing aminopropyl groups.

All components were in p.A. Quality sourced from the chemical trade (Sigma-Aldrich). The antibodies (AK) were obtained as dissolved proteins from Sigma-Aldrich or from The Binding Site. The electrolyte-containing aqueous solution used was usually 10 mmol / L potassium phosphate buffer (KPi) with a pH of 7.5 or phosphate-buffered saline (PBS). The proteins used (e.g. binding proteins such as antibodies,

Analytes) were dissolved in this buffer. To prevent non-specific binding, 0.5% Tween 20 (PBST) was added to PBS.

Antibody solutions (AK - in PBS with details of the dilution used)

AK 1: Sheep anti-human IgG (The Binding Site) 1: 2000

AK 2: Donkey anti-sheep IgG with bound peroxidase (Sigma-Aldrich) 1: 5000

AK 3: sheep anti-rabbit IgG with bound peroxidase (Sigma-Aldrich) 1: 5000

Capacitive measurement:

The change in capacity was registered with an LCR meter. The measuring device was used: LCR-6100 RS Pro, Precision LCR Meter, from IET Labs Inc. NY, USA. Measurements were made at different frequencies (10/20/100 kHz), with a voltage of 1 V applied to the silicone film. The changes in capacitance (Cp in nF / pF) and loss factor (tan_delta) associated with film changes / bonding events were recorded the correlating loss resistance (Rp in kOhm) of an imaginary

Equivalent circuit diagram. Detection of peroxidase (TMB test):

The presence of peroxidase was measured in a colorimetric test using 3, 3 ', 5,5'-tetramethylbenzidine (TMB) as the substrate (TMB dissolved in KPi buffer). The colorimetric quantitative determination of the absorption at 650 nm is transferred with a colored solution into 96 well plates and measured with a Synergy 2 microplate reader from Biotek, Germany. TMB shows a blue color after activation by peroxidase (Emax: 650 nm; usual reaction time 30 min at 25 ° C). After acidifying the solution (stopping the reaction by adding 0.1 mol / L sulfuric acid), the product turns yellow (450 nm). In the examples, the POD is often present as a coupling product bound to an antibody (POD-labeled antibody).

example 1

Non-covalent attachment of an antibody to an unmodified silicone film.

Process steps: binding of the binding protein (antibody), washing off unbound components, saturation of the film to avoid unspecific binding during subsequent use (measurement of an analyte).

A 50 μm unmodified silicone film (ELASTOSIL ^® Film 2030 250/50;

10x10 cm from WACKER Chemie AG) was incubated in 50 mL of a solution of AK1 in potassium phosphate buffer (KPi) for 1 h at 25 ° C. The solution was dumped and the film was washed three times with 50 mL KPi. Antibody AK1 is thus non-specifically bound to the film surface. To saturate potential vacant binding sites, the washed film was incubated for 1 h at 25 ° C. in PBST. The film was then washed in PBS (three times 50 ml) and was available for binding tests.

Example 2

Covalent attachment of an antibody to an epoxy-functionalized

Silicone film.

The stable covalent attachment of the antibodies is achieved by reaction of the nucleophilic amino acid residues on the protein surface (such as -NH2, -SH, phenol, Imidazole) with the epoxy groups accessible on the film under mild

physiological conditions. Due to high ionic strength in the solution of the

Antibody can make sufficient contact with the hydrophobic film surface. An epoxy-equipped film 6 x 6 cm was incubated with 10 mL antibody solution (AK1) in 1 mol / L KPi buffer at pH 7.5 for 16 h at 35 ° C with gentle swirling. The film was then washed three times with 50 mL of a 50 mmol / L KPi buffer pH 7.5 and 0.5 mol / L NaCl. To saturate unreacted epoxy groups, the film was incubated in 50 mL 1 mol / L ethanolamine pH 8 for 12 h at 35 ° C with gentle swirling and finally washed three times with 50 mL KPi buffer. To saturate potential yet unoccupied

Binding sites, the washed film was incubated for 1 h at 25 ° C. in PBST. The film was then washed in PBS (three times 50 mL) and stood for

Binding tests available.

Example 3

Covalent attachment of an antibody to a carboxyl-functionalized

Silicone film.

Carboxyl groups can be cross-linked spontaneously with NH2 groups. The reaction succeeds after activation of the COOH groups with a carbodiimide EDAC (1-ethyl-3- (3-dimethylaminopropyl) carbodiimide, hydrochloride) under mild conditions (pH 5-7) with stabilization with N-OH-succinimide (NHS).

A carboxyl group-bearing silicone film (UA-TMS modified silicone film after TMS cleavage; 6x6 cm) was dissolved in 50 mL of a solution of 20 mg EDAC, 30 mg NHS in 0.1 mol / L MES buffer ((2- (N-morpholino ) ethanesulfonic acid), pH 5.4 for 15 min at 25 ° C. with gentle swirling, the solution was poured off, the film was washed three times with 50 mL PBS and the activated film was then washed with 10 mL antibody solution (AK1) in KPi buffer Incubated for 12 h at 25 ° C. and washed three times with 50 mL of a 50 mmol / L KPi buffer pH 7.5

Saturation of potential still vacant binding sites, the washed film was incubated for 1 h at 25 ° C in PBST. The film was then washed in PBS (three times 50 mL) and was available for binding tests. Example 4

Covalent attachment of an antibody to an amine-functionalized silicone film.

Amino groups can form a covalent bond with aldehydes under mild conditions. If a dialdehyde is used, amines can be used

are cross-linked (amino group on binding protein with amino group on silicone film). An amine-containing silicone film was washed in PBS and

then in a 2% aqueous glutardialdehyde solution under gentle

Swirl incubated at 25 ° C for 60 min. After washing three times with 50 mL PBS, the activated film was incubated in 10 mL antibody solution (AK1) in KPi buffer for 12 h at 25 ° C. The film was then washed three more times with 50 ml of a 50 mmol / L KPi buffer pH 7.5. To saturate potential vacant binding sites, the washed film was incubated for 1 h at 25 ° C. in PBST. Finally, the film was washed in PBS (three times 50 mL) and was available for binding tests.

Example 5

Prevention of non-specific bonds to the silicone film:

Many molecules can bind to silicone surfaces in a non-specific way. In order to ensure the specificity of the sensor effect, substances can be applied which can prevent non-specific binding. The effectiveness of such substances was tested by incubating the sensor film with aqueous solutions of various substances (emulsifiers, surfactants) and measuring the binding of analytes to these pretreated films.

Pieces of film (4x4 cm) were incubated in 20 mL PBS solution for 5 h with gentle agitation. The solution contained the compounds to be tested in a concentration of 0.05% (w / v), 0, 1% (w / v) and 0.5% (w / v). The foils were briefly washed with PBS and transferred to a protein solution, and incubated for a further 60 min with gentle agitation. The films were washed three times for 15 minutes with PBS and then the presence of protein on the surface was measured. A peroxidase-labeled antibody (AK3) was used for the detection. The presence of protein on the slide was determined by an enzymatic test of the peroxidase determined (see above TMB test). For this purpose, the film was incubated with a TMB solution and the formation / absence of a color was measured. The test was carried out with untreated films (films without binding protein), with films to which non-specific protein was adsorbed and films to which protein was covalently bound. In all cases, comparable results were obtained: A good prevention of non-specific binding of protein is achieved with a content of 0.1% (w / v) of a solution with different Triton ^® types (example: Triton ^® X-100 (Octoxinol 9), Triton ^® WR 1339 (Tyloxapol)), various Tween ^® types (example: Tween ^® 20 (Polysorbate 20)), fatty acid esters such as Span ^® (sorbitan fatty acid ester),

Fatty alcohol polyglycol ether / fatty alcohol ethoxylates (FAEO) such as Brij ^® (example: Brij ^® 30, Brij ^® 58) or sucrose palmitate stearate.

Example 6

Capacitive measurement to detect the binding of a freely soluble analyte (protein, antibody) to a binding molecule bound to the silicone (antibody with covalent attachment):

An analyte can be measured by bringing the sensor film equipped with a binding molecule into contact in a capacitive measuring cell with a solution which contains the analyte to be measured. In the example, a sensor film equipped with an antibody is brought into contact with a solution which contains another antibody which is recognized very specifically by the binding molecule. An analyte solution containing an antibody is used as a control, but is not recognized by the binding molecule on the sensor film (this ensures good comparability).

The capacitance Cp (in pF) was measured with an LCR meter with an applied voltage of 1V and a frequency of 20 kHz. The measured in parallel

Resistance was always in the high-resistance range (indicator of the integrity of the silicone film). The initial values depend on the respective membrane preparation. The relative changes in capacity values are relevant.

A sensor foil equipped with binding molecules (silicone foil on metal foil;

Connection of AK1 to epoxy film cf. Production above) of 6x6 cm was installed in a capacitive measuring cell and rinsed in KPi buffer on the film side. To the 5 ml of a solution containing an anti-sheep antibody (AK2 solution) as the analyte to be determined were added to the measuring cell with the film equipped with the binding molecule AK1 (sheep antibody). After adding the solution, a significant increase in capacity was measured. After a short time an increased value was reached, which was stable over 60 min. After washing the measuring cell with three times 5 mL KPi buffer, the increased value was retained (see Table 1).

In a parallel experiment, the same experiment was carried out with an analyte solution which contained an anti-rabbit antibody (AK3 solution). No increase in capacity was measured after the solution was added. After this

Washing the measuring cell with three 5 mL KPi buffers retained the value (see Table 1).

At the end of the experiment, the film was removed from the measuring cell and examined for the presence of bound analyte. Since the analytes were labeled with peroxidase, the presence could be determined using a TMP test. The membrane from the experiment, which had shown an increase in capacity, showed a clearly positive TMB test (color development). The film from the test, which had shown no increase in capacity, showed no color development in the TMP test (see Table 1)

Table 1. Change in capacity when analytes are added to a film with a sheep antibody. Detection of the bound analyte (peroxidase-labeled antibodies AK2 and AK3).

Example 7

Capacitive measurement to detect the binding of a freely soluble analyte (protein, antibody) to a binding molecule bound to the silicone (antibody with non-covalent attachment)

The experiment corresponds in all aspects to the procedure set out in Example 6. The only difference was the use of an unmodified silicone film, on which the sensor molecule (antibody AK1) is only bound by non-covalent interaction (see Example 1). The

Initial capacities differ depending on the film preparation. The film used in the example had a thickness of 20 pm. The results are shown in Table 2. Table 2: Change in capacity when analytes are added to a film with a non-covalently bound sheep antibody. Detection of the bound analytes (peroxidase-labeled antibodies AK2 and AK3). A specific signal is only measured with the specific antibody.

Example 8

Capacitive measurement to detect a chemical reaction in the film mixed reaction components with reaction components in the

Electrolyte solutions

A modified silicone film (Elastosil ^® 2030A + udecenoic acid trimethylsilyl ester) is applied to a PET carrier film with a box doctor in the thickness of 50 miti and cured. A laboratory setup for the verification of the measurement evaluation of the chemical reaction is shown in Figure 7 with sketch 70. Of the

Silicone film is pulled off and placed on a PU carrier ring (71) which is stretched between two glass caps (72). The glass caps (72) are filled with electrolyte liquid (7) (tap water (pH 7)) until the silicone film is completely wetted. A silver wire ring for contacting (1 1) is inserted in the caps. This is connected to an LCR meter (LCR-6100 RS Pro) (14). The experimental data are shown in Table 3.

Table 3: Test data of chemical sensor test.

A stationary value of the capacitance at 2.4 nF was found. This was followed by the addition of sodium hydroxide solution (NaOH 1 mol / L) and a time check of the pH with indicator strips. Capacity initially increased slightly. A check of the pH showed a reduction in time. A visible abreaction and associated milky failure was evident in the electrolyte areas.

Sodium hydroxide solution was added accordingly in order to approximately maintain a pH of 1 1. As a result, the capacity increased more strongly. The pH remained largely the same after 20 min without the addition of sodium hydroxide solution. A reaction and the capacitive evaluability could thus be demonstrated. The results are shown in Table 4.

Table 4: Measured values of chemical sensor test.

Reference numerals of the figures

1: Sectional sketch of one-sided liquid sensor

2: first electrode

3: silicone layer

4: Adhesive layer or surface modification

5: passivation layer

6: binding molecules

7: electrode fluid

8: free analyte in the electrode fluid

9: specifically bound analyte

10: non-specific analyte

11 first or second electrode (contact electrode)

12 third electrode (contact electrode)

13 fourth electrode (contact electrode)

14 first electrical control and evaluation unit

15 second electrical control and evaluation unit 20 sectional sketch of two-sided liquid sensor

50 sectional sketch of gas sensor

51 first electrode

52 second electrode

53 gas sensitive material layer

54 analyte

60 sectional sketch of one-sided chemical liquid sensor 61 chemically reactive material

62 detached or reacted material

70 sectional sketch of laboratory test

71 carrier ring

72 glass caps

Claims

Expectations

1. A sensor comprising:

(a) at least two electrodes; and

(b) at least one dielectric disposed between the electrodes

Silicone layer (3), which is designed such that

an analyte can be adsorbed or absorbed on one or both sides of the silicone layer (3),

at least one further layer, which can adsorb or absorb an analyte, is arranged on one or both sides of the silicone layer (3),

a material is embedded in the silicone layer (3) or between a plurality of silicone layers (3) which can absorb an analyte or can react chemically with an analyte,

- A material is embedded in the silicone layer (3) or between several silicone layers (3), which can change its chemical and / or physical properties when exposed to light, or

2. The sensor according to claim 1, comprising (c) at least one chamber, which is arranged between the silicone layer (3) and one of the electrodes and which is suitable for receiving a conductive electrode liquid (7).

3. The sensor according to claim 2, wherein on the side or sides of the silicone layer (3) which comes into contact with the conductive electrode liquid (7), at least one further layer is arranged which can specifically bind an analyte (8).

4. The sensor according to claim 3, wherein the further layer comprises one or more binding molecules (6) selected from the group comprising antibodies, antibody fragments, enzymes, nucleic acids, receptors, transport proteins, polysaccharide binding proteins, polysaccharides, oligosaccharides and specific binding proteins.

5. The sensor according to claim 4, wherein on one or both sides of the silicone layer (3) at least one adhesive layer (4) and at least one

Passivation layer (5) is arranged.

6. The sensor according to claim 5, wherein the passivation layer (5) comprises a surfactant.

7. The sensor according to any one of claims 1 to 6, wherein at least one of the electrodes is arranged directly on the silicone layer (3).

8. The sensor of claim 1, wherein the sensor is two gas permeable

Has electrodes (51, 52) and two gas-permeable silicone foils (3) arranged between the electrodes (51, 52), between which at least one layer with material (53) which absorbs a gaseous analyte (54) or with it a chemical reaction can enter, is arranged.

9. The sensor according to claim 1, wherein on one or both sides of the silicone layer (3) at least one further layer (4) with material (61) is applied, which can enter into a chemical reaction with the analyte.

10. The sensor according to claim 1, wherein in the silicone layer (3) at least one material (61) is embedded, which can enter into a chemical reaction with the analyte.

11. The sensor of claim 1, wherein on one or both sides of the

Silicon layer (3) an analyte can be adsorbed or absorbed directly.

12. The sensor according to claim 11, wherein the analyte can be bound to the silicone layer (3) by physisorption or chemisorption.

13. The sensor according to claim 11 or 12, wherein cavities are present in the surface on one or both sides of the silicone layer (3), which can specifically bind an analyte.

14. Measuring device comprising:

- At least one sensor according to one of claims 1 to 13, and

- At least one control and evaluation unit (14, 15) which is connected to the electrodes of the sensor.

15. A method for producing a sensor according to one of claims 1 to 13, comprising the following steps:

- providing at least one silicone layer (3),

- Modification of the silicone layer (3) on one or both sides by

(i) functionalization of the silicone layer with functional groups,

(ii) introducing material into the silicone layer (3),

(iii) introducing material between two or more silicone layers (3),

(iv) applying material to the silicone layer (3) or

(v) formation of cavities in the surface of the silicone layer (3),

- Providing at least two electrodes, and

- Arrangement of the modified silicone layer between the electrodes.

16. The method according to claim 15, wherein the modification of the silicone layer (3) consists in applying a further layer on the silicone layer (3) which comprises a material which can absorb or adsorb an analyte or can react with the latter.

17. A measurement method comprising the following steps:

a) Provision of a measuring device comprising:

- At least one sensor according to one of claims 1 to 13

- At least one control and evaluation unit (14, 15) which is connected to the electrodes of the sensor;

b) bringing an analyte into contact with the sensor, exposure to light to the sensor or exposure to an electromagnetic field on the sensor; and c) measuring the change in at least one property of the silicone layer (3) as an electrical signal via the control and evaluation unit (14, 15)

18. The measuring method according to claim 17, wherein the property of the silicone layer (3) is the volume, the mass, the permittivity, the magnetic Permeability, polarizability and / or electrical resistance and electrical charge.

19. The method according to claim 18, wherein in step c) the longitudinal impedance and / or the transverse impedance of the silicone layer (3) is measured.

20. The method according to claim 19, wherein by the control and

Evaluation unit (14, 15) an electrical voltage can be applied to elastically deform the silicone layer (3).