DISPOSABLE WEARABLE SENSOR FOR CONTINUOUS MONITORING OF BREATH BIOCHEMISTRY

Abstract

An electrochemical method and an electrochemical sensor for breath analysis of single or multiple analytes using a porous, preferably flexible and disposable supporting material is provided, a salt is incorporated, and which can be wetted in contact with the exhaled breath condensate. The electrochemical method acting simultaneously as sampling method, as an electrolyte and as a support for the electrode structures. In some embodiments the salt may be hygroscopic, such that the porous substrate stays wet. To ensure that the obtained signal originates from the analyte, the electrochemical sensor preferably exhibits a differential electrode design, including a sensing (analyte-sensitive) and a blank (analyte-insensitive) electrode in order to isolate and remove the background signals.

Claims

1. An electrochemical sensor for monitoring the presence of an analyte in the breath of a subject, comprising a support comprising a porous substrate material at least one pair of working electrodes as well as at least one counter and/or reference electrode, which are applied onto and/or integrated into said porous supporting material characterized in that the at least two working electrodes comprise an analyte-sensitive sensing electrode and an analyte-insensitive blank electrode and wherein a salt is immobilized in said support, such that upon exhaling onto the sensor a differential electrochemical measurement at said pair of working electrodes allows for monitoring the presence of the analyte in the breath of said subject; and wherein the support is air-permeable such that the breath may at least partially flow through the electrochemical sensor and hygroscopic such that a liquid portion of the breath is captured to allow for dissolution of the salt immobilizes in said support.

2. The electrochemical sensor according to claim 1 characterized in that the support is air-permeable.

3. The electrochemical sensor according to claim 1 characterized in that the pair of working electrodes and the at least one reference or counter electrode are integrated onto or at least partially into the porous support.

4. The electrochemical sensor according to claim 1 characterized in that the analyte is selected from the group consisting of hydrogen peroxide, glucose, lactate, proteins, pathogens, genetic materials, hormones, hydrocarbons, aldehydes, sulfides, ammonia, ethanol, acetone, isoprene, ethane, carbonyl sulfides, carbon dioxides, carbon monoxide, nitrogen monoxide and volatile organic compounds (VOCs).

5. The electrochemical sensor according to claim 1 characterized in that the sensing electrode or porous substrate comprises an analyte-sensitive material, which is a catalyst for an electrochemical reaction of the analyte, the sensing electrode comprises an analyte-sensitive receptor, which causes an electrically measurable signal change in dependence of the analyte concentration and/or the material of the sensing electrode is supplemented with and/or coated with an analyte-sensitive material.

6. The electrochemical sensor according to claim 1 characterized in that the analyte is hydrogen peroxide and the sensing electrode comprises a metal or metal micro/nanoparticles or a mediator, as an analyte-sensitive material.

7. The electrochemical sensor according to claim 1 characterized in that the salt immobilized in said support is hydrophilic to an extent that the humidity of human exhaled breath is sufficient to form a conductive electrolyte, the salt immobilized in said support is hygroscopic to an extent sufficient to keep the porous substrate material wet, the salt immobilized in said support is selected from the group consisting of potassium chloride, sodium chloride, sodium acetate, ammonium acetate, monosodium phosphate and buffer a salt mixture, e.g. phosphate buffer, and/or the salt is immobilized in the support by applying a solution containing the salt on the porous material.

8. The electrochemical sensor according to claim 1 characterized in that the electrochemical sensor comprises at least one pair of working electrodes, at least one counter electrode and optionally, one reference electrode, and/or the electrochemical sensor comprises at least two or more pairs of working (preferably geometrically identical) electrodes targeted at the detection of one or more analytes.

9. The electrochemical sensor according to claim 1 characterized in that the at least two working electrodes are carbon, platinum, gold or silver electrodes, and/or the electrochemical sensor comprises a silver/silver chloride reference electrode and/or a a counter electrode.

10. The electrochemical sensor according to claim 1 characterized in that a structured pattern of a hydrophobic material is applied onto the porous support material and/or wherein the structured pattern may include different compartments inside the electrochemical cell in which the support and/or an electrode are sensitized for the analyte by coating and/or functionalization.

11. The electrochemical sensor according to claim 1 characterized in that the electrochemical sensor additionally comprises a processing unit configured for the reading of electrically measurable signals, of said electrodes and processing of said signals to monitor the presence of the analyte and/or the electrochemical sensor additionally comprises a communication interface for receiving and/or transmitting data to a mobile device.

12. A breath analysis and/or monitoring system comprising an electrochemical sensor according to claim 1, and a filter extension, and/or a respiratory mask, wherein the electrochemical sensor is compatible with the filter extension and/or the respiratory mask.

13. A method for an on-site or clinical monitoring of the presence of an analyte in the breath of a subject comprising providing an electrochemical sensor according to claim 1, positioning said electrochemical sensor in the respiratory flow of said subject, and employing a differential measurement by detecting the differential electrochemical signal at the at least one pair of working electrodes in order to monitor the presence of the analyte.

14. The method according to the claim 13 characterized in that the presence of the analyte is monitored continuously during a single or multiple exhaling and inhaling cycles and/or wherein different segments of the monitored signal are used in order to quantify the presence of the analyte in different regions of a lung and/or airways.

15. The method according to claim 13 characterized in that the signal detected at the analyte-insensitive blank electrode is used for a background correction of non-specific interferences of the signal detected at said analyte-sensitive blank electrode.

16. The electrochemical sensor according to claim 1 characterized in that wherein the support is selected from the group consisting of a cellulose based material, a ceramic, a hydrogel and hydrophilic polymer.

17. The electrochemical sensor according to claim 5 characterized in that wherein the analyte-sensitive material is selected from the group consisting of metal, metal oxide or semiconducting micro- or nanoparticles, enzymes, selective membranes and conductive polymers.

18. The electrochemical sensor according to claim 6 characterized in that the salt is potassium chloride.

19. The method according to claim 15 characterized in that the background correction method is able to compensate current variations caused by the respiratory movement and environmental conditions.

Description

FIGURES

[0156] The present invention is further described by reference to the following figures. The figures exemplify non-limiting and potentially preferred embodiments, presented for further illustration of the invention.

Description of the Figures

[0157] FIG. 1 (A) Schematics of chip fabrication steps including the wax isolation and the screen printing of the Ag/AgCl, the carbon and the PB-mediated electrodes, (B) CAD drawing of the electrochemical sensor with PMMA carrier, (C) SolidWorks™ model of a filter extension for respiratory mask, including the paper based hydrogen peroxide sensor and (D) image of respiratory mask with the commercial filter extension with customized sidewalls, containing the sensor chip.

[0158] FIG. 2. (A) Calibration curve of the paper based H.sub.2O.sub.2 sensors with different hydrogen peroxide concentrations: 5 to 320 μM H.sub.2O.sub.2 in 1 M KCl solution. Herein, the frontside of the chip was insulated with an adhesive tape since the sensor is placed into the filter with the backside towards the patient and thus, the frontside of the electrodes has no direct contact with the exhaled breath. Error bars represent ±standard deviation (SD) of n=7 replicates. (B) Scheme of measurement setup for simulation of respiration, including lung simulator, humidifier, H.sub.2O.sub.2 evaporator and filter housing with integrated H.sub.2O.sub.2 sensor. (C) Cyclic voltammograms of a dry chip with a PB coated working electrode, pre-treated with 1 M KCl, in vapor after 9 (grey), 24 (red), 70 (blue), 185 (green), 195 (orange) and 198 (black, dashed) breaths at a scan rate of 100 mV s.sup.−1.

[0159] FIG. 3. Signals of sensing (black) and blank (red) electrodes of an amperometric measurement at different respiration (A) frequencies and (B) volumes, (C) Current density of a calibration measurement with 5 to 320 μM H.sub.2O.sub.2 in vapor, (D) calibration curve of the aqueous and vaporous hydrogen peroxide in artificial breath. Error bars represent ±SD of n=3 replicates.

[0160] FIG. 4. CAD drawing of structures used for resolution testing of screen-printing process, containing lines with different widths and distances (0.05 to 3 mm/0.05 to 1.5 mm) and arrays of circles and squares with different diameters and edge lengths (0.05 to 3 mm).

[0161] FIG. 5. Scheme of structure and configuration for the resistance measurement with 4-point probes method to find the minimum width possible for the conducting paths. The current was applied to the outer legs and the voltage was measured at the inner legs of the structure.

[0162] FIG. 6. CAD drawings of two tested electrode designs with the same 2D area. Electrode design 1 (A) with a smaller edge area, compared to design 2 (B). The paper window constitutes the electrochemical cell, where the electrolyte droplet is placed for the measurements. The wax isolation prevents the electrolyte to spread all over the sensor.

[0163] FIG. 7. Results of multi-step amperometry for (A) carbon, (B) PB and (C) CP mediated carbon electrodes in 0.1 M PBS and 35 μM H.sub.2O.sub.2 at voltages in the range from −0.2 to 0.45 V vs. Ag/AgCl in 0.05 V steps. For PB and CP mediated electrodes, the highest signal difference between PBS and measured H.sub.2O.sub.2 concentration was observed at a voltage of 0.0 and 0.4 V, respectively. In the case of the carbon electrode, there was no significant signal change for H.sub.2O.sub.2 at the voltages of interest.

[0164] FIG. 8. Calibration curves of different electrode designs with (A) CP and (B) PB mediated paste for different H.sub.2O.sub.2 concentrations. For the CP mediated paste, design 1 had the higher sensitivity, compared to design 2. Overall, the best results were achieved with PB, where design 2 had the highest sensitivity. Error bars represent ±SD of n=5 replicates.

[0165] FIG. 9. Calibration curve of differential electrode design for H.sub.2O.sub.2 concentrations between 5 and 160 μM in 1 M KCl solution applied to the front of the sensor chip. For this calibration, the whole 3D area of the working electrode is in contact with the sample solution. Error bars represent ±SD of n=7 replicates.

[0166] FIG. 10. (A) Image of respiratory mask with extension with customized 3D printed sidewalls, containing the paper based H.sub.2O.sub.2 sensor and (B) CAD drawing of differential electrode design with a hydrogen peroxide sensing working electrode (WE), consisting of PB-mediated carbon, a carbon blank electrode (Blank) to subtract background, a silver/silver chloride reference electrode (RE), a carbon counter electrode (CE) and a PMMA cover for stabilization and isolation of conducting tracks from humidity.

[0167] FIG. 11. Plot of mean peak current over square root of scan rate for the screen-printed electrodes on paper and foil in order to determine the electrochemically active electrode area. Error bars represent ±SD of n=4 replicates.

[0168] FIG. 12. Cyclic voltammograms performed with chips under dry and wet condition and with 160 μM H.sub.2O.sub.2 at a scan rate of 100 mV s.sup.−1 in (A) 1 M KCl, (B) 0.1 M PBS and (C) 10×PBS. Please note that the results are given in current values, instead of current density, since the electrochemically active surface area of the dry sensor is undefined.

[0169] FIG. 13. (A) Image of humidifier, used for hydrogen peroxide evaporation, with heater element, intake tube and vapor outlet and (B) resulting current densities of the stability tests with 80 μM hydrogen peroxide diluted in DI water and 1 M KCl after 10 and 90 min at 0.0 V versus Ag/AgCl, where for DI water a significantly higher decrease can be observed, than for potassium chloride.

[0170] FIG. 14. Offset corrected current density of a calibration measurement using 5 to 320 μM hydrogen peroxide in vapor.

[0171] FIG. 15. The detailed plot showing the correlation of the calibration curves of the aqueous and vaporous hydrogen peroxide measurement in artificial breath. Error bars represent ±SD of n=3 replicates.

[0172] FIG. 16. Image of measurement setup employed for exhaled breath analysis, including the lung simulator, the H.sub.2O.sub.2 evaporator, humidifier, heated inspiration and expiration tubes, the housing with the sensor and a cooling trap.

[0173] FIG. 17: Chip design and fabrication showing the main elements of a preferred paper-based glucose sensor: including a wax isolation (lime green), electrodes from left to right: sensing, reference, blank and counter electrodes.

[0174] FIG. 18: Proof-of-principle study of non-invasive glucose monitoring using a paper-based electrochemical sensor. Sensor (black), blank (grey) and differential (green) current density signals (at −0.2 V vs. 1 M Ag/AgCl) during a series of applications of glucose containing aerosol puffs. For each given concentration, three consecutive puffs were applied (red triangles and dotted lines).

[0175] FIGS. 19 and 20: Images illustrating the integration of a preferred paper-based sensor into a conventional respiratory mask.

EXAMPLES

[0176] The invention is further described by the following examples. These are not intended to limit the scope of the invention, but represent preferred embodiments of aspects of the invention provided for greater illustration of the invention described herein.

[0177] The following examples report a low-cost approach for the continuous, real-time and on-site surveillance of the concentration of H.sub.2O.sub.2 in exhaled breath. The wearable system developed employs a paper based electrochemical sensor (abbreviated in the following as paper sensor) comprising a differential electrode design with a Prussian Blue (PB)-mediated carbon electrode for H.sub.2O.sub.2 detection and carbon blank electrode for subtracting the background signals. A silver/silver chloride (Ag/AgCl) reference and carbon counter electrode are used to complete the electrolytic cell. The signal detection is achieved as H.sub.2O.sub.2 oxidizes the PB, contained in the sensing electrode, which is subsequently reduced at the electrode and results in a detectable cathodic current signal. This decrease in the amperometric signal increases with increasing H.sub.2O.sub.2 concentration. For the compatibility with a standardized respiratory mask, the developed paper based H.sub.2O.sub.2 sensor is integrated into the housing of a commercially available airway filter mainly used in anaesthetic applications.

[0178] Materials and Methods Used in the Examples:

[0179] Chemical Components and Reagents

[0180] The chemicals and methods for the experiments are listed below. Unless otherwise stated, all chemicals were purchased from Sigma Aldrich, Germany. [0181] Humectants and electrolytes for sensor preparation [0182] 1 M potassium chloride (KCl) [0183] 0.1 M phosphate buffered saline (PBS) containing 0.1 M sodium chloride (NaCl) [0184] 10×PBS: 1.37 M NaCl, 27 mM KCl in 0.1 M PBS [0185] Hydrogen peroxide (30 wt %, Merck KGaA, Germany) [0186] 1 mM ferrocenemethanol for the electrochemical characterization of the paper based sensors

[0187] All electrochemical measurements in this work were performed with a potentiostat EmStat3 with an eight-channel multiplexer MUX8 and the corresponding software PSTrace 5.4 (PalmSens, The Netherlands).

[0188] Resolution of Screen-Printing

[0189] For assessing the limitation of screen printing, a mask with different structures for resolution testing was designed with CleWin (WieWeb software, The Netherlands) and ordered from Beta Layout GmbH (Germany). The test structures comprise lines with different widths (0.05 to 3 mm) and distances (0.05 to 1.5 mm), arrays of 3×3 circles with different diameters (0.05 to 3 mm) and squares of different edge lengths (0.05 to 3 mm), as illustrated in FIG. 4. These structures were screen printed onto a paper substrate, by utilizing carbon paste purchased from Gwent Group (UK) and a squeegee. With this, the minimum width, realizable with this procedure, was determined. The smallest structures screen-printable were 100 μm thin lines, but their outcome was very inconsistent. In addition, 200 μm circles and squares, showed a uniform result, except that single lines of the array did not work for the 200 and 300 μm structures. These deficiencies might also stem from irregularities in the mask for such small structures. Therefore, structures with a width less than 300 μm were not considered further for electrode design.

[0190] Resistance Measurements

[0191] To determine the minimum width with acceptable values for the conducting paths, resistance measurements were performed. As the voltage dependent current is gauged by amperometry, the resistance of the electrode structures has an impact on the sensor performance. Therefore, structures with different widths, as shown in FIG. 5, were screen printed and their resistance was measured with the 4-point probes method.

[0192] Herein, carbon paste on paper and foil substrates, as well as the Prussian Blue (PB) and cobalt phthalocyanine (CP) mediated carbon pastes on paper were tested. In addition, the resistance of structures with silver/silver chloride beneath the carbon on paper were determined. The width of the measured structures ranged between 3 to 0.5 mm. The resulting resistances for structures with different materials and different widths are summarized in Table 1.

[0193] As expected, the resistance of the structures increases with increasing width due to:

[00001]$R=\rho .Math.\frac{l}{A}$

[0194] With the electrical resistivity ρ, the length of the conductor l and the cross-sectional area A (Marinescu, M. and Winter, J., Grundlagenwissen Elektrotechnik: Gleich-, Wechsel- and Drehstrom. Vieweg+Teubner Verlag, 2011), which in composed of the width and the height h of the structure:

A=w.Math.h

[0195] Due to the high resistivity of the carbon pastes, the resistance was fairly high for the width of 1 mm preferred for the final chip design. Additional silver/silver chloride tracks were printed beneath the carbon tracks in order to decrease the resulting resistance. This width was chosen due to good results of the resolution test and as it offers an optimal size for a compact chip design.

[0196] Evaluation of Different Electrode Designs

[0197] Two different electrode shapes, with the same 2D area, but different edge lengths, resulting in a different 3D area, were designed and fabricated. The idea was to assess the influence of the edge area on the current signal, as a larger overall area should lead to a higher signal. The two chip designs with differently shaped working electrodes are depicted in FIG. 6. The bend electrode (design 2, FIG. 6B) has a larger edge area, which is 1.55-times bigger than the round electrode (design 1, FIG. 6A). Please note that the electrode height on the paper is presumed to be the same on a rigid substrate and taken from the product datasheet of the manufacturer.

[0198] To determine a suitable voltage for the amperometric signal readout, previously multi-step amperometry was performed in 0.1 M PBS and 35.28 μM H.sub.2O.sub.2 in the range between −0.2 and 0.9 V with 50 mV steps. These results are illustrated in FIG. 7A-C and show that the carbon paste shows no reaction to H.sub.2O.sub.2 in a potential range from −0.1 to 0.45 V versus Ag/AgCl (0.1 M PBS). For PB, a voltage of 0 V and for CP, 0.4 V were chosen as at these potentials highest current signals for H.sub.2O.sub.2, compared to PBS, were obtained.

[0199] To compare the two designs, calibration curves of H.sub.2O.sub.2 were taken by means of amperometry using CP- and PB-mediated electrodes. Herein, the current signals for different H.sub.2O.sub.2 concentrations at a constant voltage of 0.0 V for PB and 0.4 V for CP were recorded. The CP paste proved to deliver lower current densities than the PB paste. The sensitivity for the CP paste was 0.053 and 0.041 nA μM.sup.−1 mm.sup.−2 with correlation coefficients of 0.99 for design 1 and 2, respectively. In the case of the PB mediated paste, the sensitivities were 0.12 and 0.16 nA μM.sup.−1 mm.sup.−2 with correlation coefficients of 0.99 for design 1 and 2, vice versa. The calibration curves with the resulting mean current densities are illustrated in FIG. 8. In the case of PB, the curved structure of the working electrode results in higher current densities as the diffusion of the analyte is enhanced at the edges by using such an electrode design compared to the circular one. Surprisingly, this was not the case for the CP paste which might be caused by the inappropriate assumption of the electrode height on the paper substrate. For the final chip design, the PB-mediated carbon paste and the curved electrode (design 2) were chosen, since the obtained current densities for this combination delivered the highest signals.

[0200] Differential Electrode Design

[0201] The differential sensor design comprises two working electrodes on a single chip. One of these is the sensing electrode, containing PB as mediator and the other one consists of carbon paste without mediator serving as blank electrode to filter the background noise. Due to the similar resistance values of the carbon paste and the PB mediated paste, the current signal was expected to behave likewise and thus, signals coming from other sources than the oxidation of hydrogen peroxide could be easily excluded.

[0202] For the preferred paper based sensor with the differential electrode design, a calibration curve of H.sub.2O.sub.2, as illustrated in FIG. 9, was carried out by amperometric measurements at a voltage of 0 V versus screen-printed Ag/AgCl. Here, H.sub.2O.sub.2 concentrations in a range between 5 and 160 μM delivered a linear current response with a sensitivity of 0.23 nA μM.sup.−1 mm.sup.−2 and a correlation coefficient of 0.99.

[0203] System Integration

[0204] To enable a comfortable use of the developed H.sub.2O.sub.2 sensor in on-site or clinical breath monitoring, it is beneficial to be compatible with a common respiratory mask. For this purpose, the housing of a commercially available filter for anaesthetic applications (Ultipor® 25, Pall corporation, US) was modified. Herein, the filter was removed from the housing and the sidewalls were replaced by customized 3D printed sidewalls. These sidewalls were designed with SolidWorks 2017 (Dassault Systémes, France), so that the chip fits airtight into the housing and the contact pads of the sensor are located on the outside. They were manufactured via 3D printing with the Ultimaker 3 Extended (Geldermalsen, The Netherlands).

[0205] FIG. 10 shows an image of the respiratory mask with the extension containing the paper based H.sub.2O.sub.2 sensor. As paper itself is vulnerable, it is beneficial to stabilize the paper before the integration into the housing for the demonstrator version. Therefore, 1 mm thick poly(methyl methacrylate) (PMMA) sheets with a double sided adhesive film were lasered and the paper based sensors were placed in between two PMMA sheets. With these carriers, not only the stabilization of the paper based sensors is achieved, but also the conducting tracks are isolated against humidity, while an opening ensures that the electrodes are exposed to the breath.

[0206] Electrochemically Active Electrode Area

[0207] Due to the roughness of the paper substrate, the active surface area of the screen-printed electrodes could deviate largely from the geometric area. To examine the electrochemically active surface area, cyclic voltammograms of screen-printed carbon electrodes on both, paper and foil, were performed at different scan rates, between 25 and 200 mV s.sup.−1. First, the capacitive contribution was determined in 50 mM KCl. Then, CVs in 1 mM ferrocenemethanol were recorded to identify the peak currents I.sub.p for the reduction peaks at the different scan rates v. After subtracting the capacitive current signals, the mean values of the peak currents (n=4) were plotted over the square root of the scan rate and the slope was determined to calculate the electrochemically active electrode area A. The results are shown in FIG. 11. The relation between these variables is described in the rearranged Randles-Sevcik equation:

[00002]$A=\frac{I_p}{v^{\frac{1}{2}}}.Math.{\left(2.69.Math.{10}^5.Math.n^{\frac{3}{2}}.Math.D^{\frac{1}{2}}.Math.c_0\right)}^{-1}$

with the number of transferred electrons n, the diffusion coefficient D of the electroactive species and the bulk concentration of the redox molecules c.sub.0.

[0208] The ratio between the electrode areas on paper and foil was calculated by considering that the paper electrodes have a larger surface area. The assumption is that the foil blocks one whole side of the electrode and therefore, the area of the foil electrode only amounts 51% of the paper electrode area. However, taking this into account, the resulting experimental ratio of 0.959 implies, that the electrochemically active area on paper is insignificantly smaller than the one on foil. A possible reason might be that paper fibers block a part of the electrode surface and therewith, reduce its availability for the redox active species.

[0209] Study of Different Electrolytes

[0210] Since the paper itself is not well-conductive, it is necessary to treat the paper with an electrolyte prior to an experiment. This was done by placing an electrolyte droplet onto the paper and allowing it to dry, before measuring in vapor. For the first measurements, 0.1 M PBS was used, but during the measurement the sensor got dry quickly and it was not possible to assure constant conditions. To overcome this problem, three different electrolytes were tested, regarding their sensing performance under dry and wet conditions. The used solutions include 1 M potassium chloride, 0.1 M PBS, 10×PBS. Each of these was applied to a sensor (100 μl) and left to dry for one day. Subsequently, CVs were recorded, first with the dry sensors, second the sensor was wetted with 50 μl of DI water and finally, a droplet of 200 μl 160 μM H.sub.2O.sub.2 was added.

[0211] It turned out the tested solutions different in their capabilities to wet the paper or interfere with a hydrogen peroxide signal. CVs in 1 M KCl delivered better characteristics and results than 0.1 M PBS, as shown in FIG. 12. Furthermore, the amperometric measurement of H.sub.2O.sub.2 in 1 M KCl provided a better sensitivity than in 0.1 M PBS. Therefore, 1 M KCl was chosen as electrolyte for further experiments.

[0212] Stability of Hydrogen Peroxide Solution

[0213] For evaporating hydrogen peroxide, a commercially available humidifier HME-BOOSTER® (Medisize, The Netherlands), consisting of a heater element, an intake for injecting the solution and an outlet for the vapor, was employed as shown in FIG. 13. During evaporation of hydrogen peroxide, diluted in a KCl solution, salt crystals formed, which blocked the pores of the humidifier. Therefore, from then on hydrogen peroxide was diluted in deionized water (DI water). This led to the problem, that lower current signals were observed. For this reason, the stability of H.sub.2O.sub.2 in DI water was studied and compared to H.sub.2O.sub.2 stability in 1 M KCl. Herein, amperometric measurements were performed with a fresh solution of H.sub.2O.sub.2 (10 minutes) and again after 90 minutes at 0 V (FIG. 13). As the stability de facto was worse in DI water than in 1 M KCl, but potassium chloride was crystallized in the evaporator, the compromise was to prepare the stock solution with 1 M KCl and then dilute the stock solution with DI water to the desired concentration immediately before evaporating.

[0214] Offset Correction of Measured Current Signals

[0215] The measured current densities of blank and signal electrodes do not have the same baseline.

[0216] This can be corrected by setting these signals to “zero” using an offset value prior to the calibration measurement.

[0217] Correlation of Vaporous and Aqueous H.sub.2O.sub.2 Measurement

[0218] By division with a constant factor, a parallel linear plot (FIG. 3D) can be obtained which is very close to the calibration curve in FIG. 2A. Thus, the sensitivity of the sensor is reproduced under the measuring conditions in the water-saturated vapor. Therefore, it can be assumend that the paper sensor shows the same response with the same sensitivity as in solution. By adding a factor to the x values, the measured values in vapor can be superposed perfectly with the calibration curve. After this operation, the ratio between the original H.sub.2O.sub.2 concentrations of the prepared solutions used for the artificial breath and the obtained correlated concentrations are not completely conserved. For example, the ratio is 320:160=2 in aqueous, but only 42.91:22.73=1.89 in vapor. However, this may be caused by inaccuracies in the supply by the perfusor, and the evaporation would cause concentration fluctuations in the vapor.

[0219] Measurement Setup for Exhaled Breath Analysis

[0220] For the exhaled breath analysis, a measurement setup was installed and human respiration, as well as H.sub.2O.sub.2 containing breath were simulated. An image of this setup is illustrated in FIG. 16.

Example 1: Fabrication Procedure for the Electrochemical Sensor

[0221] The fabrication procedure for the paper sensors is illustrated schematically in FIG. 1A. First, wax patterns are printed on chromatography paper (grade 1 CHR, 200×200 mm.sup.2, Whatman, UK) using a commercially available wax printer (ColorQube 8580, Xerox corporation, US) and baked for 10 minutes at 120° C. in a conventional oven. When heated, the layer of wax printed on the surface of paper wicks through the bulk of the substrate and forms a hydrophobic barrier, defining the electrolytic cell. The wax barrier plays two important roles: i) It prevents wicking of any droplets of water condensed during exhalation to the contact pads during operation. ii) The wax pattern contains a solution of electrolyte before the water is evaporated from the substrate to form a solid-electrolyte. Next, the Ag/AgCl reference electrode (RE) and conducting tracks are screen-printed and baked for 10 minutes at 80° C. Finally, the carbon counter (CE), blank and PB-mediated sensing electrodes are screen-printed and baked for 15 minutes at 80° C.

[0222] The paper based sensor chip is placed inside a wearable respiratory mask, such that the patient is breathing directly onto the sensor. For integration into the ventilation mask, the paper chip is glued between two PMMA sheets with an opening for the electrodes, as depicted in FIG. 1B. With this, the sensor is mechanically stabilized and, at the same time, the conducting tracks are isolated from potential shorts due to water droplets originating from exhaled breath. Furthermore, the housing of a commercial filter is modified (FIG. 1C) by replacing the sidewalls with custom-made 3D printed parts to mount the paper based sensor into the housing which allows to place it directly in the respiratory flow. With this approach, moisture from the breath is captured by the paper sensor to humidify the paper substrate, forming an electrochemical cell, which is crucial for the operability of the sensor. The entire system is illustrated in FIG. 1D.

[0223] Because paper itself is not ionically conductive, a droplet of electrolyte is placed on the paper and dried, before measuring analytes from exhaled breath. For the first measurements, 0.1 M phosphate buffered saline (PBS) is used as electrolyte, but during the measurement the sensor dries more quickly making it more difficult to maintain constant conditions. To solve this problem, three different electrolytes were tested (see FIG. 12). Cyclic voltammograms (CVs) have been performed in dry (after one day) and wet (DI water added) condition and finally, with a droplet of 160 μM H.sub.2O.sub.2. The compounds tested exhibited difference in their ability to keep the paper wet and a possible interference with the detection of H.sub.2O.sub.2. According to the test results, 1 M potassium chloride (KCl) provides the best characteristics for the CVs and sensitivity for H.sub.2O.sub.2 in amperometric measurements. It has been, thus, chosen as an electrolyte salt for further experiments.

Example 2: Calibration and Amperometric Measurements Using the Sensor

[0224] For the calibration of the paper based H.sub.2O.sub.2 sensor, the current behaviour over time is recorded for different hydrogen peroxide concentrations. Amperometry at a constant potential of 0.0 V versus Ag/AgCl (screen-printed RE electrode) is carried out using different paper chips (n=7). The frontside of the electrodes is isolated using an adhesive tape, as the paper sensors are positioned in the respiratory mask with the backside facing the user, hence, the frontside of the electrode structures has no direct contact with the exhaled breath. First, a droplet of 1 M KCl solution is placed on the electrolytic cell of the paper chip, and then, measurements with increasing the H.sub.2O.sub.2 concentration are performed. The obtained calibration curve is shown in FIG. 2A. Here, a linear measurement range between 5 to 320 μM hydrogen peroxide is achieved with a sensitivity of 0.19 nA μM.sup.−1 mm.sup.−2 and a correlation coefficient of 0.99.

[0225] To mimic the human respiration, it is necessary to create a periodic air flow generating a warm and humid gas flow using a lung simulator, as the human exhaled breath contains ˜100% RH at a temperature of around 34° C. Using a customized LabVIEW software (National Instruments, USA), the lung simulator pumps a desired volume of air with a predefined frequency. RH and temperature are adjusted using a commercially available humidifier (HumiCare® 200, Gründler Medical, Germany) that contains heated tubing. To introduce different concentrations of H.sub.2O.sub.2, an evaporator with a heating element is placed in between the lung simulator and the paper sensor. A scheme of this setup is illustrated in FIG. 2B.

[0226] Since the moisture content of paper is varying with changing RH during inhalation and exhalation, to study the effect of RH on the redox characteristics of the PB-mediated carbon electrode, CV measurements using a dry chip, pre-treated with 1 M KCl, in H.sub.2O.sub.2-free simulated breath were performed. In all experiments (except the tests of respiration frequency and volume), the lung simulator is set to generate a tidal volume of 500 ml and a frequency of 15 breaths per minute, which are realistic values for a healthy adult. As it can be observed in FIG. 2C, the initially dry sensor can be wetted only by the respiratory stream itself, assuring a high and more reliable electrochemical signal. After 195 breaths (13 minutes), the measured current signals do not alter anymore and exhibit a typical PB-CV of a wet sensor in 1 M KCl.

[0227] In FIG. 3A, the current signal of an amperometric measurement at different respiration frequencies is illustrated. It is noticeable, that slower breathing results in a lower frequency of the signal measured and vice versa, while no significant signal change is observed by a volume change (see FIG. 3B). During one breathing period, the water content in the paper changes periodically, as the air stream is drier during inhalation and reaches a RH of ˜100% during exhaling. Accordingly, the ionic conductivity of the paper fibers changes.sup.21,23. Even though variations in conductivity are less important in an amperometric setup, the signal might decrease if the paper becomes too dry (see FIG. 2C). However, as the blank electrodes without PB show a similar response, we conclude that the periodic variations must be mostly attributed to capacitive currents due to the humidity dependent changes of the dielectric properties of the paper.sup.24. Probably only fibers in direct contact with the electrode surface contribute to this effect. Hence, this capacitive part of the current is probably quite sensitive to the surface morphology of every individual electrode and is expected to reach its maximum at the reversal points of the respiratory movement.

[0228] In order to obtain a calibration curve for hydrogen peroxide in the vapor of the artificial breath, H.sub.2O.sub.2 solutions of different concentrations are evaporated and the current signal over time is recorded continuously. A typical measurement is shown in FIG. 3C. As soon as a steady-state current is reached, the next higher concentration is added, as indicated with the arrows labelled with the corresponding concentration. The response time to obtain a steady-state current depends on the peroxide concentration in the vapor. The reason for this behaviour is probably due to the time required for the concentration of H.sub.2O.sub.2 in vapor to equilibrate with its dissolved form in water (i.e. dissolved in the moisture within paper). At higher H.sub.2O.sub.2 concentrations in the vapor, the gradient between vapor and “paper electrolyte” is higher and thus, a steeper current increase, as well as a higher limiting current are expected which is almost in line with our observations.

[0229] The behaviour of the blank (background) electrode can be also observed in FIG. 3A-C. The measured blank signals do not settle at the same baseline currents as the sensing electrode. However, these different baseline currents can be aligned for the evaluation by setting an offset value (see FIG. 14).

[0230] For the construction of the calibration curve, the current densities of the blank curve are first subtracted from those of the sensor electrode. After averaging and baseline subtraction, a measurement value is taken for each hydrogen peroxide concentration at a point on the timeline shortly before the next higher concentration is introduced. The baseline value of the sensor is taken right before addition of the first hydrogen peroxide concentration of 40 μM. The mean values for the calibration curve presented in FIG. 3D are obtained from three independent measurements.

[0231] From these results, it can be concluded that H.sub.2O.sub.2 concentrations in the range between 40 and 320 μM give rise to a response with a sensitivity of 0.02 nA μM.sup.−1 mm.sup.−2 and a correlation coefficient of 0.99. It is crucial to note, however, that the resulting current signals for the respective H.sub.2O.sub.2 concentrations are significantly lower than of the former calibration in aqueous solutions (FIG. 2A). This may be due to the heating of the H.sub.2O.sub.2 in the evaporator and the poor stability of H.sub.2O.sub.2 in DI water (see FIG. 13). By correlating the obtained current densities with those of the previous calibration in solution, the real H.sub.2O.sub.2 concentrations in the vapor of the artificial breath can be estimated to lie between 5 and 40 μM. This means that, after evaporation, the H.sub.2O.sub.2 concentration may be decreasing to approximately ⅛ of its original value while the same sensitivity as in solution is maintained, i.e. 0.19 nA μM.sup.−1 mm.sup.−2 (FIG. 3D and FIG. 15). The humid air from the humidifier needs may also be diluting the analyte yielding a smaller concentration. Nevertheless, after accounting for all these factors, a reliable quantification of different hydrogen peroxide concentrations in vapor is achieved and, the proof-of-concept for on-site H.sub.2O.sub.2 analysis in exhaled breath is successfully demonstrated.

[0232] In summary, this example describes a differential electrochemical method using low-cost porous materials (for example, a low-cost cellulose paper) for on-site monitoring of hydrogen peroxide in exhaled breath. For compatibility with standardized ventilation masks, the sensor developed may be integrated into the housing of a commercially available airway filter for anaesthetic applications. Under realistic conditions by simulating human respiration with authentic lung volume and respiration rate, the proof-of-principle of the hydrogen peroxide measurements in exhaled breath are successfully shown for the first time. With further modifications and improvements, this sensor model can be employed in a large variety of applications, including clinical or wearable monitoring of exhaled breath.

[0233] As evident from the data described herein the claimed method and sensor have the following advantages: (i) Because of differential measurements, the influence of various interfering substances and/or environmental conditions (for example, temperature and humidity) are eliminated, hence, the system always produces reliable results. (ii) By changing or modifying and/or coating the material of the substrate or the sensing electrode (for instance, with metals, metal oxide- or semiconducting micro- and nanoparticles, enzymes, selective membranes or conducting polymers), the sensor model presented can be extended for the analysis of other compounds from exhaled breath. (iii) A flexible and hygroscopic porous support, like paper, acts as a “solid electrolyte” eliminating the need for additional membranes (containing the electrolyte) and at the same time as a substrate for the electrodes. (iv) Flexible and porous substrates can be shaped and patterned in a way that the sensing surface as well as the collection volume can be considerably increased. (v) The orientation and porosity of the sensing surface can be tuned to minimize breathing resistance and to improve signal quality (i.e. signal-to-noise ratio).

[0234] The performance of this method and sensor can be further enhanced by: (i) screening for further humectants as possible electrolytes to facilitate the handling and signal processing, for example, by keeping the porous substrate wet and ensuring that the sensor does not need to adsorb any humidity from the breath. (ii) PB-mediated carbon paste with different PB contents and modification procedures may be further tested in order to further increase the H.sub.2O.sub.2 sensitivity. Alternatively, hydrophilic metal electrodes (especially Pt), realized by metallization of fabrics, may be employed.sup.25. Moreover, the implemented sensor system may be extended with a compact and low-power wearable signal readout unit along with a smartphone app to enable on-site monitoring.

Example 3: Design and Proof-of-Principle for a Paper-Based Glucose Sensor

[0235] Chip Design and Fabrication for a Paper-Based Glucose Sensor

[0236] The design and fabrication procedure for the paper-based glucose sensors were carried out according the methods described above in relation a paper-based hydrogen peroxide sensor and shown schematically in FIG. 16.

[0237] The only difference in the chip design is the use of two identical compartments, separated with wax, but still employing a common reference and a counter electrode. The fabrication starts with printing wax patterns on the chromatography paper (grade 1 CHR, 200×200 mm.sup.2, Whatman, U.K.) by means of a commercial wax printer (ColorQube 8580, Xerox corporation, USA). This is followed by a 10-min bake at 120° C. in an oven which results in the wicking of wax printed through the paper substrate and thus, defines a hydrophilic area for the electrolytic cell. At the next step, the reference electrode (RE) and conducting tracks are screen-printed with silver/silver chloride (Ag/AgCl) paste (C2040308P2, Gwent Group, U.K.) and baked for 10 min at 80° C. Last, the carbon counter (CE) and PB-mediated working electrodes are screen-printed using the carbon and mediated carbon pastes (C2030519P4 and C2070424P2, Gwent Group, U.K.).

[0238] On-Paper Functionalization of Glucose Oxidase

[0239] Glucose oxidase (GOx) solved in 1 M potassium chloride (KCl) is adsorbed into the compartment surrounding the sensing electrode. Alternatively, GOx can be either immobilized covalently (for example, by using glutaraldehyde), encapsulated (by polyethylenimine), or entrapped in a gel (such as hydrogel) into the paper substrate.sup.32.

[0240] GOx catalyses the oxidation of glucose into hydrogen peroxide (H.sub.2O.sub.2) which can be reduced at the screen-printed Prussian Blue (PB)-mediated carbon electrode. The measured current relates directly to the glucose concentration of the sample. A second, identical cell, but only treated with 1 M KCl (without GOx), enables to subtract background signals and periodic variations caused by the respiratory movement.

[0241] Results of a Proof of Principle Study for the Non-Invasive Glucose Sensing Approach

[0242] To demonstrate the proof-of-principle of the described non-invasive glucose sensing approach, the sensor is exposed to aerosols of different glucose concentrations (5 μM to 10 mM) using a deodorant nebulizer. Within the physiological range, a stepwise increase of the differential current signals (see FIG. 18) at consecutive aerosol administrations is observed. At higher concentrations, a peak followed by a current decay is noticed, possibly due to limited oxygen supply and thus, a limitation of enzyme activity.

[0243] Glucose entrapped in aerosols can be cumulatively sampled and directly measured with paper-based sensors at concentrations of less than 5 μM. Compared to EBC analysis, our approach minimizes the risk of analyte degradation, while considerably reducing acquisition time and system's complexity. It also allows continuous glucose monitoring.

[0244] As illustrated in FIGS. 19 and 20 advantageously the paper-based sensor may be integrated into a conventional respiratory mask in a straightforward manner. To this end the paper-based sensor can be easily applied directly onto the mask, e.g. using by isolating it with a flexible tape. Via a suitable chip connector and wires the sensor may be connected to a potentiostat, which is connected to a mobile device such as a smart phone.

[0245] In conclusion a proof-of-principle of a facile, inexpensive and non-invasive approach for the simultaneous sampling and measurement of exhaled glucose could be demonstrated, for the first time. Further improvement may include the characterization and optimization of the developed system in simulated breath, followed by a further clinical validation.

TABLE-US-00001 TABLE 1 Mean values of resistances for screen-printed structures with different widths and materials. Mean value of resistance in Ω Width in mm 3 2.5 2 1.5 1 0.9 0.2 Carbon 261.2 ± 45 296.8 ± 55 385.8 ± 52 477.4 ± 56  765.2 ± 127 843.2 ± 108 1477.0 ± 285 n = 5 on paper Carbon 228.3 ± 6  276.7 ± 16 375.7 ± 42 475.0 ± 58 682.0 ± 60 769.0 ± 83  1372.0 ± 48  n = 3 on foil PB 294.3 ± 46 313.7 ± 29 358.0 ± 17 465.7 ± 28 691.0 ± 35 820.0 ± 135 1345.0 ± 75  n = 3 mediated carbon on paper CP 242.7 ± 23 267.7 ± 6  306.3 ± 11 420.3 ± 32 662.7 ± 52 818.3 ± 147 1414.3 ± 320 n = 3 mediated carbon on paper Ag/AgCl 0.07 ± 0   0.08 ± 0.01  0.1 ± 0   0.12 ± 0.02   0.19 ± 0.02 .sup. 0.20 ± 0.01   0.50 ± 0.06 n = 3 tracks beneath carbon on paper

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