WO2021112802A1

WO2021112802A1 - Impedimetric/capacitive reusable blood glucose measurement with molecular imprinted polymers

Info

Publication number: WO2021112802A1
Application number: PCT/TR2020/051199
Authority: WO
Inventors: Zihni Onur UYGUN
Original assignee: Ege Universitesi
Current assignee: Ege Universitesi
Priority date: 2019-12-04
Filing date: 2020-12-01
Publication date: 2021-06-10
Anticipated expiration: 2022-06-04

Abstract

The present invention relates to a sensor that can be used to measure blood glucose using molecular imprinting technology. The present invention generally relates to fast and precise, single frequency based impedimetric/capacitive electrochemical sensors in order to determine the amount of sugar (glucose, analyte) in liquid samples.

Description

DESCRIPTION

IMPEDIMETRIC/CAPACITIVE REUSABLE BLOOD GLUCOSE MEASUREMENT WITH MOLECULAR IMPRINTED POLYMERS

Technical Field

The present invention relates to a sensor that can be used to measure blood glucose using molecular imprinting technology. The invention generally relates to a fast and precise, single frequency based impedimetric/capacitive electrochemical sensor that determines the amount of sugar (glucose, analyte) in liquid samples. This electrochemical sensor contains molecularly imprinted polymers (MIPs) for use in real-time measurements of sugar molecules in liquid samples. The robust and strong structure of the molecularly imprinted polymers allows the sensor to be reused

Known State of the Art (Prior Art)

The glucose molecule or dextrose is found in the blood in a chair conformation. When we look at its general structure, its state in the chair conformation, namely the a-D-Glucose form, this is the form that circulates freely in the blood and is phosphatized when it enters the cell, and it is the primary energy source of the body. Its sensitive and rapid measurement is extremely important in people suffering from diabetes. Glucose measurement strips and devices known in the art are disposable electrodes where glucose oxidase (GO_x) or glucose dehydrogenase (GDH) bio-recognizing enzyme systems are placed on the surface, which recognize glucose in the blood sample and enable it to be degraded enzymatically. The measurement is carried out by transmitting the electrons given by the bioelectroactive molecules formed by glucose, which undergoes a biochemical change, to the measuring device as an electrochemical signal. The use of enzyme systems in strips has limitations such as sensitivity -specificity problems and the systems allowing single use. For example, since the GO_x system works in an oxygen dependent manner, its sensitivity is affected by the oxygen level in the blood. In contrast, although GDH has high sensitivity, it is not as specific as GO_x, as it can interfere with other substances. In addition, the interaction of enzyme systems with humidity and temperature in the environment can affect the accuracy of the measurements. The most important disadvantages of enzyme systems are that they are open to some interactions such as hematocrit and oxygen levels in the sample, and they are not suitable for reuse in a cost-effective and effective way. Normally, the electrodes that measure with this method can only be reused by removing the glucose and therefore the blood sample to be measured from the strip. On the other hand, in these measuring systems carrying enzymes, the necessary conditions for the enzyme activity will be affected during the cleaning process, so reuse is not possible.

Today's blood glucose analyzers, which are the biosensor class, are analysis systems developed by immobilization of a biological sensor, that is, an enzyme, receptor, antibody, DNA or protein molecule, on a physicochemical transducer that has affinity only for the analyte molecule. The signals resulting from the interaction between the biological sensor and the analyte molecule are transmitted to the analysis system by the "transducer" and the measurement is performed by analyzing the concentration-dependent response of this signal.

Biosensors can be catalytic or affinity based, depending on the biosensor on them. In catalytic-based biosensor systems, an enzyme enzymatically degrades the analyte molecule and measurement is performed on the resulting secondary molecules or with tertiary molecules that can make these secondary molecules measurable. In affinity-based biosensors, analysis is performed by measuring the degree of antigen-antibody, DNA-DNA, protein- ligand binding.

Affinity-based biosensors are more advantageous than catalytic ones in that they do not require any additional molecules other than the receptor-ligand pairs given. Both types of biosensors can be designed based on electrochemical, optical and piezoelectricity according to the "transducer" type. If the signals emerging as a result of these physicochemical changes are electrochemical, "transducer" type electrical signals can be sensors. For example, if electroactive products are formed as an enzyme converts a substrate or the conductivity changes due to the charge distribution of the electrode surface, it is advantageous to measure the analyte electrochemically. On the other hand, if the resulting product is a product that absorbs or emits optically light, then an optical biosensor should be used. If there is a change in pressure, velocity, and strain, it is more appropriate to design a piezoelectric biosensor that converts these changes into electrical charges.

Among these methods, electrochemical biosensor systems are the least susceptible to interference, yet the most practical and cost-effective systems that can provide measurement in almost any type of substance. The low cost and ease of use of electrochemical biosensor systems cause them to be preferred more in scientific studies.

Biosensor systems developed on the basis of bioaffmity are biosensor systems based on the binding kinetics of biomolecule and analyte. Immune system biomolecules, single-stranded DNA, artificial single-stranded DNA (Aptamer) or cell surface receptors can be used as bio recognition agents in bioaffinity-based biosensors. In recent years, highly specific analyzes can be made in the field of health with biosensors. Electrochemically developed affinity biosensors are usually based on measurement by signals from a secondary antibody molecule or a secondary marker molecule specific to the analyte molecule. Generally, electrochemical measurement can be performed as the secondary molecule changes the electric current in a measurable way. Although biosensor systems seem to be very advantageous and sensitive measurement systems due to the specificity of biological sensors to analyte, the use of biological receptors as a sensor molecule limits the analysis when the environmental conditions of the biosensor are considered. Optimum working conditions are required for the efficient operation of biological molecules. These are physical properties such as pH, temperature, pressure, light, ion strength, polarity of the liquid being measured. Even small changes in these properties can affect a measurement. For example, acetyl choline esterase enzyme activity is extremely sensitive to pH changes. Another example is the regeneration of antigen-antibody based biosensors. Disruption of antigen-antibody interactions in this process is achieved by changing the ionic strength, but in the meantime, the molecules can be damaged. Another example is the effects of low ambient temperature. In this case, the activity of the used enzyme may not be observed/observed low. As in these examples, a number of interactions restrict the use of biosensor systems and pose a major obstacle to practical applications.

For the aforementioned reasons, in recent years, more durable sensor materials that are less affected by environmental conditions have been designed. Here, instead of the term biosensor, the term sensor is more appropriate, because synthetic sensors are produced instead of biosensors. The production and application of these artificial receptors is carried out by polymerizing monomers with special functional groups in such a way that they surround the analyte molecule according to its three-dimensional properties. Thus, it is possible to design specialized cavities for certain molecules to enter on these polymers. This artificial receptor creation technology is called molecular imprinting or molecular imprinting technology (MIT). The polymers developed in this type are also called molecular imprinted polymers or molecularly imprinted polymers (MIP). MIT, in brief, is the technology of producing monomers on this mould, which will surround the molecule in accordance with the three- dimensional structure of a mould molecule in complicated and complex solutions. The chemical structure of the monomers of the artificial polymeric receptor to be formed is chosen according to the functional group or groups that will interact with the target molecule to be bound. For example, if there are amine groups or groups that can form a positive charge on the analyte, that is, the target molecule, the monomer molecule must functionally contain negatively charged groups in order for the target molecule to surround/attract. In this way, the "fingerprint" of the target molecule is created on the polymer structure and an artificial receptor containing the cavities that only the target molecule can enter, is obtained Afterwards, these durable polymers, on which specific cavities are formed, can selectively recognize and bind the target molecule.

MIT is a relatively new technology that can perform different tasks such as molecular recognition, catalysis, chromatographic separation, chemical identification in different solvents. These polymers are not easily affected by the physical conditions in which bio receptors are affected and they are more closed to interference. However, their production is also less costly than biologically derived receptors. All these advantages show that the use of MIP instead of biological molecules is a more appropriate choice in sensor systems.

Molecular imprinted polymers (MIPs) are frequently used because of their properties, such as selectivity to the mould molecule, their stable structure, resistance to heat and pressure, resistance to chemicals and reusability. Today, molecular imprinting technique has been successfully combined with various analytical techniques, such as chromatography, sensor etc.

However, selectivity is an important problem in sensors developed with MIP known in the art. In the applied methods, when a different electrical stress is applied, other molecules other than glucose can undergo electrochemical transformation, so they can generate signals. For these reasons, there is a need to develop a glucose-selective sensor that is not affected by other components in the liquid sample in order to measure blood glucose accurately. Brief Description and Aims of the Invention

The invention relates to an impedimetric/capacitive electrochemical sensor modified with molecular imprinted polymers (MIPs) that can quickly detect the amount of sugar (glucose, analyte) in liquid samples, such as blood, plasma, urine, etc. based on a single frequency. In the sensor of the invention, glucose is bound to a unique cavity located on the sensor surface. Since the glucose molecule entering into the glucose-specific cavity with this sensor will only generate a binding-separation signal and there is no electrochemical reaction, it is highly selective since the impedimetric/capacitive method is used as the method only to measure the binding. With the invention, an electrochemical sensor has been developed that is not affected by environmental conditions, can be used multiple times and does not interfere with substances other than sugar.

With the sensor of the invention, the interaction of the bioreceptor-analyte molecule can be determined by measurement of surface capacitance or impedance without the use of secondary molecules. Since impedance and capacitance allow examining the surface characteristics of the electrodes electrically, binding of only the bioreceptor-analyte is sufficient for the measurement.

Within the scope of the invention, new generation sensors have been designed by diol formation to contain artificial recognition agents to be used for glucose measurement, that is, molecularly imprinted polymers and boronic acid derivatives that can interact with hydroxyl groups on glucose entering these imprinted cavities. These polymers were synthesized electrochemically (in situ) using molecular imprinting technology and placed on the sensor that contains strips, that is, glucose-imprinted polymers. Thus, the new generation strips were developed using an artificial receptor that would directly bind to the target molecule glucose on an affinity basis, not an indirect recognition agent (enzyme).

The measurement of the binding of glucose to the MIP placed on the graphene layer on the gold and platinum electrodes was measured by single frequency impedance spectroscopy. Since EIS can measure the thickness of the electrode surface or the charge distribution (capacitance), it is a sensitive method to determine even small changes that change the electrical charge distribution on the electrode surface. With EIS, electrode surface capacitance can be measured using a potentiostat without any biochemical reaction. Single frequency impedance is called a non-electrochemical method by measuring glucose binding at a frequency that does not change in terms of time-dependent resistance, i.e. impedance.

In the invention, the binding characteristic of glucose to the strip surface is observed, and surface impedance and capacitance are used as the measurement method. A glucose sensor modified with glucose binding MIPs formed on the graphene layer between a platinum electrode (electron source) and a gold electrode (electron acceptor) has been developed. With the frequency potential applied between these electrodes, the glucose amount can be determined by impedimetric and capacitive measurement.

With the sensor of the invention, glucose concentrations in different samples can be determined by measuring the increase in the impedance of the strip/electrode and the decrease in capacitance in unit of time, caused by the glucose specifically bound to these polymers in the new generation strips, which are designed as modified with MIP on the sensor surface and contain glucose-imprinted polymers.

Since the change in the electrical charge distribution of the surface of the inventive sensor electrodes reduces the surface capacitance (C) together with the impedance, it is important to monitor glucose binding not only by impedance measurements but also by capacitance measurements. In addition, since these measurements will follow the correlated changes in unit of time, the developed glucose measurement basis is in chrono-impedimetric and chrono- capacitive properties.

With the invention, a strip modified with glucose binding MIPs formed on the graphene layer between a platinum electrode (electron source) and a gold electrode (electron acceptor) has been developed. With the frequency potential applied between electrodes, the glucose amount can be determined by impedimetric and capacitive measurement.

Definitions of Drawings Illustrating the Invention

The figures prepared for a better understanding of the device developed with this invention are explained below.

Figure 1: Side view and parts of the glucose sensor Figure 2:Top view of the glucose sensor Figure 3: Side integrated view of the glucose sensor Definitions of Elements/Pieces/Parts Forming the Invention

In order to explain better the blood glucose measurement sensor developed with this invention, the parts/pieces/elements in the figures are numbered separately and the explanation of each number is given below.

1: Insulating Support Layer

2: Plastic Insulating Layer

3: MIP with Glucose-Selective Cavities

4: Gold Electrode

5: Graphene Connection Layer

6: Platinum Electrode

7: Molecular Glucose-Imprinted Polymer Layer 8: Sample Chamber 9: Copper Conductor Wires

Detailed Description of the Invention

The invention relates to an electrochemical sensor modified with molecular imprinted polymers (MIPs) that can quickly detect the amount of sugar (glucose, analyte) in liquid samples, such as blood, plasma, urine, etc. based on a single frequency

Figure 1 shows the side view and parts of the inventive sensor.

The inventive sensor that detects the glucose amount in liquid samples includes the insulating support layer (1) used as the top layer coating and increasing the durability of the sensor surface, the plastic insulating layer (2) containing the electrodes (4, 6), the MIP (3) with glucose selective cavities, the gold electrode (4), which interacts electrically with the platinum electrode (6) through the graphene connection layer (5), the graphene connection layer (5) which is located on the plastic insulating layer (2) between the gold and platinum electrodes (4,6) and conducts the electric current between these two electrodes, the platinum electrode (6) that interacts electrically with the gold electrode (4) through the graphene connection layer (5), molecular glucose imprinted polymer layer (7) specific to glucose, which is the analyte desired to be determined, placed on the graphene connection layer (5), the sample chamber (8), which can receive at least 50 pL of sample, which is defined to bring the sample in which the sugar amount is to be determined to contact the sensor, copper conductive wires (9) used for the external connection of gold (4) and platinum (6) electrodes.

The Production of Molecular Glucose Imprinted Polymer (Artificial Recognition Agent or Artificial Receptor)

Two types of monomers are used in the production of molecular imprinted polymers. In the range of 0.2-1 mg, preferably 1 mg of acrylamidophenyl boronic acid (AAPBA), used as the primary monomer, is dissolved in a solvent with a pH of 7 and containing 50 mM of dihydrogen phosphate. Then, in the range of 0.1-5 mg, preferably 1 mg of glucose (Glc) is added into this mixture and 0.01-0.5 mg, preferably 0.5 mg of pyrrole is added to this mixture as secondary monomer and it is waited at room temperature (~25 °C) for a maximum of 4 hours until the sensor is prepared. This mixture, which is necessary for the production of molecularly imprinted polymers (3) with glucose selective cavities, is added on the graphene connection layer (5) and left for a while for the substances that need to dissolve to self- organize and spread. Then, a constant potential in the range of 600 mV to 900 mv is applied between the gold (4) and platinum (6) electrodes and on the graphene connection layer (5) for at most 20 minutes. Thus, on the graphene connection layer (5), MIP (3), which contains glucose-selective cavities that can be bound to the three-dimensional structure of glucose and hydroxyl groups on glucose is produced. These polymers were synthesized electrochemically (in situ) using MIT and formed on the graphene connection layer on the surface of the strip (graphene layer coated with MIP). Thus, the new generation strips were developed using an artificial receptor that would directly bind to the target molecule glucose on an affinity basis, not an indirect recognition agent (enzyme). The capacity of the polymer increases due to the ability of AAPBA to form diol with glucose and to polymerize pyrrole electrochemically. These monomers have polymerized around glucose, with this polymerization, cavities were obtained where only glucose can enter, that is, cavities arranged in accordance with the three- dimensional structure of glucose and on which glucose can enter. These cavities are formed by arranging the boronic acid derivatives that can bind to both the three-dimensional structure of glucose and also to the hydroxyl groups on glucose accordingly. Thus, this cavity was oriented like a key-lock model so that only glucose can enter. Sugar Determination in Liquid Samples with Electrochemical Sensor of the Invention

After the MIP (3) with glucose selective cavities is obtained, the liquid samples whose glucose amount is to be determined are dropped into the sample chamber (8) by the sensor of the invention. A maximum potential of 200 mV is applied to the sensor with a frequency in the range of 100-150 Hz. The binding of glucose in the liquid sample to the glucose selective MIP cavities on the graphene layer increases impedance and decreases capacitance. Converting the rate of increase in impedance to ratio of glucose confirms the selectivity of glucose by showing a decrease in the decrease in capacitance due to glucose binding. Thus, the amount of sugar in liquid samples can be detected with the sensor of the invention.

After the sample is dropped, 200 mV potential and a frequency between 100 and 150Hz will be applied. During this period, an increase in impedance will be observed per unit time and depending on this increase, the amount of sugar can be measured. Capacitance measurement is used here as a control mechanism. When there is binding to the surface, that is, to the molecularly imprinted polymers on the graphene layer, the increase in capacitance is independent of the glucose concentration, but if there is binding to the cavities, in other words, if glucose is bound to the cavities, the change in capacitance correlates with the glucose concentration. With glucose binding to the cavities, an increaseis observed over time and then, no increase is observed. However, if it accumulates on the surface, this increase continues, so when both the impedance increase and the capacitance increase are correlated with each other, the correct amount of glucose is measured.

Glucose measurement is carried out by glucose binding to selective cavities produced for allowing glucose on MIPs coated on the graphene layer to enter. This binding is achieved through the formation of diol of the boronic acid ends (B-OH) on AAPBA and the hydroxyl (OH) ends on glucose, while the aldehyde (C=0) group on glucose is achieved by forming a hydrogen bond with the imino(-NH-) group on the pyrrole by interaction. The interacting boronic acid and imino groups are polymerized after self-arrangement around glucose, and the cavities are thus formed in accordance with the three-dimensional structure of glucose. Since this hollow glucose is suitable for its three-dimensional structure, only glucose can enter these cavities. This can be thought as a key-lock match.

With the sensor of the invention, glucose concentrations in different samples can be determined by measuring the impedance/capacitance changes of strip/electrode and the decrease in capacitance in unit of time, caused by the glucose specifically bound to these polymers in the new generation strips, which are designed as modified with MIT on its surface and contain glucose-imprinted polymers. Since the change in the electrical charge distribution of the surface of the electrodes changes the surface capacitance (C) together with the impedance, it is important to monitor glucose binding not only by impedance measurements but also by capacitance measurements. In addition, since these measurements will follow the correlated changes in unit of time, the developed glucose measurement basis is in chronoimpedimetric and chrono-capacitive properties.

In this invention, the binding characteristic of glucose to the surface is observed, and surface impedance and capacitance are used as the measurement method. Impedance (Z or R) or electrochemical impedance spectroscopy (EIS) is an effective measurement technique used in examining the electrolyte-electrode interface, in measuring mass transfer rates and investigating electrode reactions. With this measurement, non-electroactive large mass proteins and antigens can be measured at very low detection ranges and very low detection limits.

Not requiring a secondary molecule, it is more economical than other electrochemical methods in terms of only measuring the interaction between two molecules. This measurement is simply a technique that measures the surface composition and surface resistance of the electrodes. With the binding of glucose in the sample to the surface of the graphene layer coated with MIP, a change in surface resistance occurred depending on the glucose concentration, and this change and the binding rate were measured impedimetrically. Since there will be a change in time-dependent impedance in this measurement, one aspect of the measurement basis is chronoimpedimetric. Especially for the interpretation of glucose binding as a method, glucose measurement is performed with the algorithm created with capacitive measurement. This algorithm was calculated to be 1 millimeter square of graphene coated MIP area and derived as follows: with binding of glucose to the surface, an increase in capacitance and impedance is observed, this increase continues for 3 seconds, after this second there is no increase in capacitance, but increase in impedance continues and an increase is achieved in impedance with glucose concentration. If the capacitance increases and the impedance does not change after 3 seconds, it means that non-glucose molecules accumulate on the surface. This time increases as the graphene-coated area increases. In order for the sensor to be used again after the measurement, the electrode is put into ethanol-water mixture and kept for 5 minutes. The sensor can be reused after the Glc has been removed (washing)

The measuring range of glucose in a liquid sample is in the range of 20 mg/dL and 800 mg/dL with this sensor.

In the invention, there can be preferably a very thin cellulose membrane that protects the MIP at the point where this liquid sample contacts. This membrane has the ability to hold the shaped elements originating from blood and it was made of nitrocellulose material of 10 pm thickness that allows the passage of molecules that are below 1000 Daltons.

The sensor of the invention is a lateral flow sensor. It is designed as a 3.5mm headphone jack that can be connected to a device, mobile phone, analyzer or a platform by means of copper conductive wires (9) connected separately to the gold (4) and platinum (6) electrodes of this sensor.

There is a distance of 1 mm between the gold conductor and the platinum conductor (graphene connection layer length). This distance can change, with the change of this distance, the amount of graphene, AAPBA, Pyrrole must be multiplied by the percentage increase in distance.

Claims

1. An electrochemical sensor that detects the amount of glucose in liquid samples; comprising

• An insulating support layer (1) used as the top layer coating in the sensor and increasing the durability of the sensor surface,

• A plastic insulating layer (2) containing the electrodes (4, 6),

• MIP (3) with glucose selective cavities,

• A gold electrode (4), which interacts electrically with the platinum electrode (6) through the graphene connection layer (5),

• A graphene connection layer (5) which is located on the plastic insulating layer (2) between the gold and platinum electrodes (4,6) and conducts the electric current between these two electrodes,

• A platinum electrode (6), which interacts electrically with the gold electrode (4) through the graphene connection layer (5),

• Molecular glucose imprinted polymer layer (7) specific to glucose, which is the analyte desired to be determined, placed on the graphene connection layer (5),

• A sample chamber (8), which can receive at least 50 uL of sample, which is defined to bring the sample in which the sugar amount is to be determined to contact the sensor,

• Copper conductive wires (9) used for the external connection of gold (4) and platinum (6) electrodes.

2. The production method of the sensor according to claim 1; characterized in that the production of MIP (3) with glucose selective cavities includes the steps of

• Dissolving 0.2-1 mg of aery 1 ami dophenyl boronic acid (AAPBA), which is the primary monomer, in a solvent with a pH of 7 and containing 50 mM of dihydrogen phosphate,

• Adding 0.1-5 mg of glucose,

• Adding pyrrole (Py) in the range of 0.01-0.5 mg as the secondary monomer and leaving it at room temperature for a maximum of 4 hours, • Adding the mixture on the graphene connection layer (5) and holding the substances that need to dissolve in it for self-arrangement and spreading,

• Applying a constant potential in the range of 600 mV to 900 mv between the gold (4) and platinum (6) electrodes and on the graphene connection layer (5) for at most 20 minutes,

• Forming diol of the boronic acid ends (B-OH) on AAPBA and the hydroxyl (OH) ends on glucose, obtaining MIP (3) on the graphene connection layer (5), which contains glucose-selective cavities that can be bound to the three-dimensional structure of glucose and hydroxyl groups on glucose by forming a hydrogen bond with aldehyde (C=0) on glucose with the imino (-NH-) group on the pyrrole, polymerization of boronic acid and imino groups around glucose.

3. A method of determining the glucose amount in liquid samples of the sensor according to Claim 1, characterized by comprising the steps of;

• Dripping the liquid samples whose sugar amount is desired to be determined into the sample chamber (8),

• Applying a maximum potential of 200 mV to the sensor with a frequency in the range of 100- 150 Hz,

• The binding of glucose in the liquid sample to the glucose selective MIP cavities on the graphene layer,

• Glucose bound to glucose selective MIP cavities on the graphene layer increases impedance and decreases capacitance,

• Converting the rate of increase in impedance to ratio of glucose, confirming the selectivity of glucose by showing a decrease in the decrease in capacitance due to glucose binding

4. A sensor according to claim 1; characterized in that it comprises a cellulose membrane which protects the MIP at the point of contact of this liquid sample.

5. A sensor according to claim 4; characterized in that said cellulose membrane allows the passage of molecules that are below 1000 Daltons.

6. A sensor according to claim 4; characterized in that said cellulose membrane is nitrocellulose having a 10 pm thickness.

7. A sensor according to claim 1; characterized in that it can be connected to a device, mobile phone, analyzer or a platform by means of copper conductive wires (9) that are connected separately to the gold (4) and platinum (6) electrodes.

8. A sensor according to claim 1; characterized in that the measuring range of glucose is in the range of 20 mg/dL to 800 mg/dL.