A biosensor system for monitoring biomarkers of body fluids

Abstract

The invention is related to a non-invasive biosensor system, especially wearable system having multiple biosensors for monitoring biomarkers in body fluid.

Description

[0001] A BIOSENSOR SYSTEM FOR MONITORING BIOMARKERS OF BODY FLUIDS

[0002] Technical Field

[0003] The invention is related to a non-invasive biosensor system, especially wearable system having multiple biosensors for monitoring biomarkers in body fluid.

[0004] Prior Art

[0005] Recently, wearable electronic sensors have gained much attention due to their suitability in real-time and continuous monitoring of human body’s physiological biomarkers and their potential for personalized health monitoring and modular therapeutic device applications. These wearable biosensors can be broadly classified into two categories based on the mode of sampling, such as invasive and non-invasive methods. The non-invasive sampling sources can be saliva, sweat, tear, skin interstitial fluid or urine that are ideal choices for wearable biosensing because these samples are readily accessible outside the skin barrier and are on emerging trends in the commercial market. Mostly, fabrication of the wearable biosensing devices requires an alternative to classical fabrication approaches that necessitates requirement of the flexible substrates.

[0006] Developing a non-invasive wearable device to probe physiological signals is crucial for diagnosing and monitoring of diseases, offering sensitive, rapid and cost-effective analysis, especially for chronic conditions like diabetes, and for detecting physiological biomarkers such as glucose, cytokines, or metabolites to enable early health monitoring. Diabetic patients currently require frequent and uncomfortable blood glucose monitoring through painful invasive finger pricking, which raises the risk of insulin errors and sudden changes in levels of glucose. Addressing this pressing global health need, the development of non-invasive systems for continuous vital sign monitoring is essential.

[0007] Inflammatory cytokines, linked to chronic inflammation and age-related diseases, also hold significance as health biomarkers. Elevated levels of interleukin-6 (IL-6) have also been linked to acute stressors and cortisol secretion during psychological stress. Electrolytes are other crucial biomarkers for human health monitoring and have been associated with cystic fibrosis and autonomic neuropathy assessment. Calcium (Ca2+ions) in particular, is an essential mineral for human metabolism and about 1-2% of human body weight. Elevated levels of ionized calcium in bio-fluids have harmful effects on the structure and function of many organs of the human body.

[0008] Existing methods for the monitoring of the physiological biomarkers (e.g., glucose, IL-6 and Ca2+ions) rely on invasive blood sampling procedures. Most common among these methods are blood-based glucose sensor, ELISA tests, and chemical assays, which are not only complex, but also time-consuming.

[0009] There have been few attempts to detect target analytes in sweat, saliva and tears by developing non-invasive electrochemical sensor devices. Wearable saliva sensors are complicated for real-time continuous on-body applications. Similarly, non-invasive health monitoring can be carried out using sensors that utilize tears bio-fluid, which poses several limitations such as the delicate nature of human eyes, as well as evaporation during / after collection of tears.

[0010] Sweat is an excellent bio-fluid for non-invasive health monitoring because of the presence of a variety of biomolecules, such as larger protein to hormones, metabolite and electrolytes in sweat. Progress in colorimetric transducing signal generation has been made using wearable sensors to monitor target biomarkers in sweat while in presence of different indicator dyes. However, these require multi-step processing and external readout analysis components to detect analytes in a biofluid.

[0011] Currently, sweat is used in medical diagnosis of cystic fibrosis, electrolyte imbalance for athletic training, autonomic neuropathy and drug abuse detection. Utilizing sweat as a biofluid for detection and monitoring of vital health markers is still limited. However, possible means by which sweat analytes can be analyzed is by the usefulness of unique nanomaterials in sensing, such as graphene in bio-electronic sensors. Some devices such as those integrated with Near-Field Communication (NFC) components facilitate easy signal / data transfer process for remote health monitoring. Most of the existing non-invasive wearable sensors designed for monitoring vital signs are currently limited to probing heart rate and individual’s physical activity tracking, and these do not meet to the requirements of assessing the levels of an individual’s molecular parameters, nor provide insights into the internal health state. Recently, enzymatic and non-enzymatic methodologies have been discussed in order to develop flexible sweat-based glucose biosensor. Development of non-invasive electrochemical sensors recently began to progress toward utilizing sweat as a non-invasive source for biomarkers monitoring. Most of the reported wearable sensors focused on detecting one analyte at a time. Such wearable sensors have been utilized for detecting metabolites (glucose, lactate, and alcohol), electrolytes (K+ / Ca+), protein biomarkers and heavy metal ions. A few studies also reported on determining hormone and immune responses in sweat that utilizes ZnO transducing flexible platforms. The epidermal wearable sweat sensors function on the basis of iontophoresis and tattoo-based printing methods. However, most epidermal wearable sweat monitoring sensors reported thus far present challenges of mechanical friction that are subject to deformation of implanted device, especially on a human skin.

[0012] Very few sensor devices reported thus far integrate multiple detection modes on a signal sweat sensing platform. These multi-analytes electrochemical sensing technology demonstrated monitoring of a variety of sweat biomarkers, such as glucose, Na+, K+Ca2+and pH. Simultaneous noninvasive multi-analyte sensing is extremely attractive but requires an accurate monitoring system. Most of the reported sensors utilized metal electrodes, carbon fibers, carbon printed electrode for sweat multi-analyte detection that presents several major challenges; specific bio-anchoring without altering sensor electrical properties, temperature effects and resolution, which is an event detection method that measures the smallest unit defining the interfacial event to produce a discernible change in the signal.

[0013] Notable advancement has been made in a recent study which utilized flexible and stretchable graphene electrodes for monitoring sweat glucose during a physical exercise. The developments in flexible wearable biosensors have also been previously reviewed that describe on a variety of nanocomposite-based flexible sweat glucose biosensing. The nanocomposites used in such flexible sensors comprised of materials, such as metal and metal oxides as well as non-metallic materials including graphene. Despite of the advantages of graphene’s flexibility or stretchability, there exists several challenges before its full potential to be exploited, such as limitation with respect to integrating multiple graphene sensor modes for multi-modal capability, sensitivity and signal amplification for accurate multiplexed sensing of sweat biomarkers, and finally reliable monitoring of all parameters in a single device. For example, sweat glucose and electrolytes’ levels are relatively in much smaller concentrations than they are normally found in blood. Therefore, the ability to detect extremely low levels of biomarkers present in sweat is a challenging task due to weak electrochemical and / or bio-affinity of transducing sensor surface. Tedious sweat collection process, long term stability, biocompatibility and on-site signal conditioning for continuous monitoring of sweat biomarkers are other major challenges that hamper their performance.

[0014] The new generation of wearable biosensors can be designed via integrating flexible electrodes made of graphene nanostructures to develop electrochemical sweat sensors, which can enhance the electrochemical biosensing signal and its sensitivity. Biocompatibility of graphene nanostructure combined with its high charge carrier mobility, extraordinary physical and optical properties, large surface area enables homogenous functionalization of chemical / bio-linking. The above graphene features make it one of the most desirable nanomaterials for developing flexible and wearable biosensor. Further, surface modification of graphene’s two-dimensional structure with suitable surface chemistry facilitates easy functionalization of a variety of functional groups with strong affinity without any geometric constrains to coupling of chemical / biomolecules. There is a need to introduce a new dimension to improved graphene application, exploiting its superior properties mainly in sweat biosensing and monitoring applications. Introducing graphene through interfacing with wearable sweat sensing patch could overcome the existing challenges related to the non- invasive sweat multiplexed biomarker monitoring methods.

[0015] As a result, all of the problem mentioned above has made it necessary to provide a novelty in the related field.

[0016] Brief Description and Objects of the Invention

[0017] The main objective of the present invention is to establish a non-invasive biosensor system, especially wearable system having multiple biosensors for monitoring biomarkers in body fluid, especially in sweat, and provide real-time, simultaneous, rapid, cost-effective and sensitive measuring. This invention can work with small sweat volumes (10-30 pL).

[0018] Thus, present invention uses specific design of biosensor which is printed on a flexible substrate and a fluid transfer system to transfer body fluid to mentioned biosensor. The proposed biosensor is type of interdigitated biosensor (IDE) and comprises concentrically provided multiple conductive arcs and a graphene nanostructure interfaced between theses arc and these arcs flanked by a counter and a reference electrode. This design enhances sensitivity of signal by its superior charge-carrier mobility in conjunction with metal electrodes and enlarge electrochemically active surface area allowing sensitive electrochemical signal generation. The invention also proposes a microfluidic system transfer works according to capillary action. Capillary action is the ability of a liquid to flow in narrow spaces without the assistance of, or against, external forces like gravity and does not need any effort to move fluids. In this invention, the microfluidic system provides transfer from the source of biofluid to the biosensors without any support which is important for especially wearable system and prevent evaporating of bioliquids which is important because the body of subject already produce very little biofluids, for example sweat. Moreover, the microfluidics module helps maintain the levels of sweat constituents that reflect vital health signs and also drives sweat to the sensor area using capillary forces, which allows for accurate measurement of sweat biomarker levels. The synergetic effect of improved sensitivity augmented transfer of body fluid to the biosensor, and the expanded electrochemically active surface collectively contribute to superior and accelerated signal measurements. So, the biosensor system of invention can work with small in small sweat volumes (10-30 pL).

[0019] In preferred embodiment of invention, an absorbent inlet pad is used for absorbing more body fluid and enhance measuring quality for by enabling more body fluid to the biosensors.

[0020] Furthermore, an absorbent outlet pad is used in the outlet. This outlet pad provides multiple advantages. The first is the flow is accelerated in system and more sweat transferred to the biosensor by this acceleration and this provides more sweat passes through biosensor and enhance measuring quality. Also, the outlet pad provides absorption of excess sweat from system which also cause flow increasing.

[0021] This invention also aimed at developing a fully integrated non-invasive, label-free graphene- interfaced wearable patch device for real-time monitoring of multi-analytes in sweat. The wearable patch device was designed and fabricated that consist of graphene interfaced electrochemical sensors arrays integrated with microfluidic platform on a suitable flexible, preferably polymeric, substrate and a read-out signal processor unit. The patch device was fabricated to specifically targeting three different model sweat analytes, such as a sweat cytokine, interleukin-6 (IL-6), sweat glucose and sweat Ca2+ions on a single platform in a multiplexed format.

[0022] The fully integrated electrochemical sensor patch device developed in this invention enables real-time and multiplexed sensing of model three sweat analytes. This device not only allows quantification and monitoring of analytes but also enables on-site signal analysis. The functionality of developed sensing patch device relies on the changes in electrochemical properties of graphene (current / voltage) after specific capturing the sweat analytes on sensor transducing surface. Interfacing graphene in between the metal electrode arrays is a crucial step in fabricating a graphene-based bio / nano-sensor system, to improve the charge transfer ability of sensors for sensitive and specific capturing of sweat biomarkers. Insufficient interfacial contact can hinder the synergetic effects and adversely affect the sensor’s performance. Therefore, here a flexible polymeric substrate, such as polyimide (PI) was utilized. PI carries dynamic charges in its chemical potential, which allows for the coating / doping of specific carriers into graphene. Such substrates not only provide flexibility and a large interfacial area in between electrodes on sensor surface, but also offer a high signal-to-noise ratio, making them highly desirable for a wearable device. Additionally, the successful design and application of graphene facilitated the creation of a miniaturized multimodal option for a wearable patch device. This was achieved through a miniaturized sensor patch design that incorporated overlapped graphene interfaces, ensuring the integrity of electric / electronic contact for unbiased and sensitive sensor measurements.

[0023] The developed wearable electrochemical sweat sensor patch design is inexpensive and portable in its characteristics that has the potential for exhibiting outstanding sweat sensing performances. Thus, the developed wearable graphene interfaced sweat sensing patch device ensures its potential application in multi-analyte sweat biomarkers sensing and provides a non-invasive approach for personalized disease monitoring and management applications. The success and outcomes of the invention address the current challenges in the field of invasive clinical diagnosis, which resulted in delivering a wearable technology platform for bio- and chemical sensing. This platform enables personalized non-invasive disease monitoring applicable to health monitoring, sports / athletes, and army personal fitness, and in various other clinical applications. The outcome of this invention contributes to generate tools with socio-economic benefit which enhances competitiveness while contributing to safety, better health and improving the quality of life.

[0024] Description of the Figures of the Invention

[0025] The figures and related descriptions necessary for the subject matter of the invention to be understood better are given below.

[0026] Figure 1. A schematic view of the system that represent a flexible wearable sensor patch patterned with an array of sensors wherein each sensor’s working electrodes are biofunctionalized with specific receptors. Figure la. A schematic view of the biosensor of the invention.

[0027] Figure 2. Another schematic view of the system.

[0028] Figure 2a. A representative cross-section view of the invention.

[0029] Figure 3a. A graphic show that XRD patterns of graphite and GO.

[0030] Figure 3b. A graphic that shows that a Raman spectrum of graphite, GO and rGO.

[0031] Figure 4a. A graphic that shows XRD patterns of polyimide film substrate (PI) that was coated with graphene.

[0032] Figure 4b. A graphic that shows Raman spectra of PI substrate, pure graphene, and graphene coated PI substrate.

[0033] Figure 4c. A graphic that shows thermogravimetric analysis (TGA) analysis of PI substrate and graphene coated PI substrate.

[0034] Figure 5a. A scanning electron microscopic (SEM) images of graphene coated PI substrate.

[0035] Figure 5b. A scanning electron microscopic (SEM) images of graphene coated PI substrate, after 25 washes.

[0036] Figure 5c. A scanning electron microscopic (SEM) images of graphene coated PI substrate, after 50 washes.

[0037] Figure 6a. An optical image of control gold interdigitated electrode arrays sensor on flexible polyimide substrate without graphene.

[0038] Figure 6b. Another optical image of control gold interdigitated electrode arrays sensor on flexible polyimide substrate without graphene.

[0039] Figure 6c. An optical image of single sensor showing graphene-interfaced, gold microelectrode arrays on flexible polyimide substrate.

[0040] Figure 6d. Another optical image of single sensor showing graphene-interfaced, gold microelectrode arrays on flexible polyimide substrate.

[0041] Figure 6e. A SEM image of graphene-interfaced sensor electrodes on flexible polyimide substrate.

[0042] Figure 6f. A SEM image of graphene-interfaced sensor electrodes on flexible polyimide substrate at higher magnification than Fig. 6e. Figure 7a. A graphic that shows FTIR spectra of graphene interfaced sensor patch immobilized with IL-6 antibodies.

[0043] Figure 7b. A graphic that shows FTIR spectra of graphene interfaced sensor patch immobilized with IgG monoclonal antibody and glucose oxidase enzyme.

[0044] Figure 7c. A graphic that shows FTIR spectra of graphene interfaced sensor patch immobilized with Ca2+-ion selective membrane (Ca-ISM).

[0045] Figure 8a. A graphic that shows stability and repeatability response of sensor patch for IL-6 (0 pg / mL, only antibody) on respective sweat sensor patches during 6 successive cycles (column 1), stability of the sensor patch signal over the 1-7 days (column 2), and responses measured from three independent sensors (column 3).

[0046] Figure 8b. A graphic that shows stability and repeatability response of sensor patch for glucose (0.4 mM) on respective sweat sensor patches during 6 successive cycles (column 1), stability of the sensor patch signal over the 1-7 days (column 2), and responses measured from three independent sensors (column 3).

[0047] Figure 8c. A graphic that shows stability and repeatability response of sensor patch for Ca2+(0.75 mM) on respective sweat sensor patches during 6 successive cycles (column 1), stability of the sensor patch signal over the 1-7 days (column 2), and responses measured from three independent sensors (column 3).

[0048] Figure 9a. A graphic that shows CV responses of graphene interfaced electrochemical sensor tested with glucose (0.1 mM) at varying temperatures 25 °C ~ 35 °C).

[0049] Figure 9b. A graphic that shows cyclic voltammogram (CV) response of graphene electrochemical sensor arrays immobilized with Glucose oxidase (GOx) monitored against varying concentrations of glucose at different pH values from 6.4 to 7.4 under applied potentials of-0.1 V to +0.1 V.

[0050] Figure 10. A graphic that shows performance of the microfluidic platform tested using physiologically relevant solution (PBS, pH 7.4) and the flow rate as a function of time.

[0051] Figure Ila. A graphic that shows cyclic voltammogram (CV) response of graphene electrochemical sensor immobilized with anti-IL6 antibodies was monitored against varying concentrations of IL-6 sweat biomarker at applied electric potential of -0.1 V to +0.1 V. Figure 11b. A graphic that shows the linear curve of IL-6 sweat sensor patch at different concentrations of IL-6 sweat biomarker.

[0052] Figure 11c. A graphic that shows CV response of IL-6 sweat sensor patch in the presence of chemical and biologically relevant species of sweat such as glucose, and a control BSA.

[0053] Figure lid. A graphic that shows CV response of graphene electrochemical sensor arrays immobilized with Glucose oxidase (GOx) monitored against varying concentrations of glucose at applied potential of -0.1 V to +0.1 V.

[0054] Figure lie. A graphic that shows the linear response curve from sensor patch against different concentrations of glucose in synthetic sweat.

[0055] Figure Ilf. A graphic that shows CV response from glucose sweat sensor patch in presence of relevant common bio / chemical species in sweat.

[0056] Figure 11g. A graphic that shows CV responses of graphene electrochemical sensor functionalized with Ca2+-ion selective membrane.

[0057] Figure llh. A graphic that shows linear response curve obtained from Ca2+ion sensor patch incubated with synthetic sweat sample containing different concentrations of Ca2+ions.

[0058] Figure Hi. A graphic that shows CV response from Ca2+ion sweat sensor patch tested in the presence of relevant common bio / chemical species of sweat.

[0059] Figure 12a. A graphic that shows linear regression plot of data from gold standard tests carried out using varying concentrations of IL-6 protein in synthetic sweat test samples.

[0060] Figure 12b. A graphic that shows linear regression plot of data from gold standard tests carried out using varying concentrations of glucose in synthetic sweat test samples.

[0061] Figure 12c. A graphic that shows linear regression plot of data from gold standard tests carried out using varying concentrations of Ca2+ions spiked in synthetic sweat test samples.

[0062] Reference Numbers

[0063] The parts and components are given in the figures are referenced for the subject matter of the invention to be understood better.

[0064] 1. Biosensor system

[0065] 10. Flexible substrate

[0066] 11. Sensor reaction chamber 12. Inlet reservoir

[0067] 13. Outlet reservoir

[0068] 14. Transfer channel

[0069] 15. Outlet channel

[0070] 20. Microfluidic system

[0071] 21. Main channel

[0072] 22. Reservoir channel

[0073] 30. Inlet pad

[0074] 40. Biosensor

[0075] 41. Counter electrode

[0076] 42. Reference electrode

[0077] 43. Working electrode

[0078] 44. Graphene coated surface

[0079] 50. Outlet pad

[0080] 60. Upper layer

[0081] Detailed Description of the Invention

[0082] The invention is related to a non-invasive biosensor system (1), especially wearable system having multiple biosensors (40) for monitoring biomarkers in body fluid. The body fluid is preferably saliva, sweat, tear, interstitial fluid, intravascular fluid or urine.

[0083] Referring to Figure 1, la, 2 and 2a; the biosensor system (1) comprises a flexible substrate (10). Preferably, the flexible substrate (10) is a polymer. The polymer is selected between polyimide or polyethylene terephthalate, silicone polymers, polyurethane, or cellulose-based papers. In a preferred embodiment, the flexible substrate (10) is made of polyimide. The flexibility enables use of biosensor system (1) on the skin of subject, for example on the arm. The biosensor system (1) may be a wearable system and may be configured to connect skin of the subject by adhesive surface or connection mean like a strap, etc. On the surface of the flexible substrate (10), multiple reservoirs are formed. These reservoirs include at least sensor reaction chambers (11) and an inlet reservoir (12). The system has at least two sensor reaction chambers (11) and preferably at least three or more.

[0084] The sensor reaction chambers (11) are formed to provide housing for the biosensors (40) and monitoring space for body fluids and the biosensor (40) and the inlet reservoir is formed to enable entrance of the body fluid to the biosensor system (1). The sensor reaction chambers (11) and the inlet reservoir (12) are connected to each other in a such way that the body fluid is transferred from the inlet reservoir (12) to the sensor reaction chamber (11).

[0085] As can be seen in Fig. 2, the sensor reaction chambers (11) and the inlet reservoir (12) are connected to each other by a microfluidic system (20). The microfluidic system (20) is configured to transfer the body fluid by capillary action. Capillary action, also known as capillarity, is the ability of a liquid to flow in narrow spaces without the assistance of, or against, external forces like gravity. This phenomenon is particularly noticeable in thin tubes, capillaries, or porous materials, where the adhesive forces between the liquid and the solid surface pull the liquid upward against the force of gravity.

[0086] In a preferred embodiment, the microfluidic system (20) comprises a main channel (21) and multiple reservoir channels (22). One end of the main channel (21) is connected to the inlet reservoir (12) and the reservoir channels (22) is connected between the main channel (21) and the sensor reaction chambers (11). Other end of the main channel (21) may connect the reservoir channels (22) which connects the last sensor reaction chamber (11). There is one reservoir channel (22) for each the sensor reaction chambers (11).

[0087] Preferably, the diameter of the main channel (21) is larger than the diameter of the sensor reaction chamber channels (22). Diameter difference between the main channel (21) and the sensor reaction chamber channels (22) enables directional flow of body fluid toward sensors by capillary action.

[0088] Furthermore, the reservoir channels (22) have serpentine structure. Preferably, serpentine form comprises parallel section connected to each other by curved section. This form enables to use of longer channels which facilitates mixing, increased surface area, and regulate flow of liquid by capillary action. Referring to Fig. la; the system comprises an interdigitated biosensors (40) to monitor body fluid. The interdigitated biosensors (40) have a reference electrode (41), a counter electrode (42) and an arc-shaped interdigitated working electrode (43). The reference electrode (41) and the counter electrode (42) flank the working electrode (43). All the electrodes are made of metal, preferably gold. Alternatively, platinum, palladium, silver, copper, nickel or aluminum can be used as material of the electrodes. The biosensor (40) is printed over at least part of the sensor reaction chambers (11).

[0089] An alternating and parallel pad connection orientations in biosensor (40) array on the biosensor system (1) can be provided for making convenient electrical contacts. The multimodal electrode design, comprising three biosensors (40), each carry respective reference electrode (41), counter electrode (42) and working electrode (43) that extend with connection pads, not only ensures the integrity of electric and electronic contact but also exhibits adaptability to various applications. This design allows for versatile and precise signal measurements, making it suitable for a wide range of analytical and diagnostic purposes.

[0090] The working electrode (43) features multiple arcs made of metal, which are coaxially positioned. Arc shaped segments extend to interlock with the corresponding arc shaped segments extending from opposite side, forming an interdigitated working electrode (43). Between two sequential arcs a graphene coated surface (44) is provided. The graphene coated surface (44) is made of graphene or its derivatives. Coating is formed as a nanostructure and in a preferred embodiment, nanostructure can be graphene, single-walled or multi-walled carbon nanotubes. To provide such interdigitated biosensors (40), the surface of a flexible substrate (10), at least the sensor reaction chamber is coated with graphene or its derivatives and after that the biosensor (40) is printed over said coated surface.

[0091] Graphene nanostructures (100-200 nm nanosheets) were interfaced between gold interdigitated arc-shaped electrodes (IDE) each of these microelectrodes measure 40-60 pm in width separated by 40 pm gap within 5 mm diameter of a complete working electrode (43) area. The arc design shape of metal / graphene interdigitated sensor electrodes (<40-60 pm in width with a distance between two electrodes of about 40 pm) facilitate liquid sweat droplet (10-30 pL) impingement and hydrodynamic coalescence property allowing better contact and interaction with bio- / chemical receptors (antibodies / enzyme / ions) on sensors that enhances signal -to-noise ratio and sensitivity for electrochemical detection. In a preferred embodiment, at least part of the counter (41) electrode and the reference electrode (42) is circular, more preferably in crescent form.

[0092] In a preferred embodiment, each biosensor (40) made of a circular working electrode (43) measuring the diameter of 5 mm with 18 alternating pm-sized interdigitated gold metal and nanostructured graphene micro-electrodes (IDE) in a total sensor area of ~7 mm x 10 mm on a flexible PI substrate. The working electrode (43) is flanked by crescent shaped reference electrode (RE) and counter electrode (CE).

[0093] Each biosensor (40) is biofunctionalized with a receptor for specific biomarker and according to that, each biosensor captures targeted biomarkers in body fluids transferred to the sensor reaction chamber (11). Capture means, it can be configured to target specific disease, such as diabetes (glucose levels), chronic disease inflammation (IL-6, a cytokine) and kidney disorders or neuromuscular or cystic fibrosis diseases (Ca2+) and for this purpose, the capture means it can be selected between receptors, such as anti-IL-6 antibodies, glucose oxidase enzyme probe and Calcium ion selective membrane, respectively. It should be noted that the capture of targeted biomarkers is not limited to the specified analytes, but this capability also extends to the detection of other analytes. This versatility extends the potential applications to various diseases, including but not limited to diabetes (glucose), chronic inflammatory disease (cancer-immunotherapy) or psychological stress (IL-6), and cystic fibrosis / kidney disease (Ca2+levels), all in a single run. Each biosensor (40) in array was dedicated to a specific analyte through bio-functionalizing with a specific bio- / chemical receptor designed to interact with a specific sweat analyte in a complex mixture without cross-reactivity.

[0094] Each biosensor (40) in the array operates independently and selectively, even when dealing with multiple analytes or modes. This is achieved by utilizing highly specific surface bio- / chemistries and electrically decoupling the operating conditions of each biosensor (40) interface. As a result, each sensor provides a distinct electrochemical signal, creating a unique signature that allows for discriminating between different biomarkers and determining their concentration ratios. All of this can be accomplished using a portable handheld potentiostat unit equipped with a simple profiling algorithm and Bluetooth connectivity. In this invention, the obtained data from biosensors is transferred as an output to a smartphone application or potentiostat. Each biosensor (40) is designed to capture different biomarker in the body fluid for disease monitoring. In preferred embodiment, the biosensor system (1) comprises three biosensors (40) that are bioactivated with specific receptors, and they comprise anti-IL-6 antibodies, glucose oxidase enzyme probe, and calcium ion membrane, respectively.

[0095] Accordingly, the biosensor system (1) comprises:

[0096] - A flexible substrate (10) having at least two sensor reaction chamber (11), an inlet reservoir (12) for receiving the body fluids,

[0097] - A microfluidic system (20) enables to transfer the body fluid from the inlet reservoir (11) to the sensor reaction chamber (11) by capillary action of fluids,

[0098] - Interdigitated biosensors (40) printed in the sensor reaction chambers (11) having,

[0099] A working electrode (43) comprises concentrically provided multiple conductive arcs and a carbon nanostructure interfaced between the arcs and

[0100] A conductive counter electrode (41) and a conductive reference electrode (42) which flanks the working electrode (43) wherein, the nanostructure is made of graphene and its derivatives and

[0101] Capture indicates for each biosensor (40) to capture targeted biomarkers in body fluids transferred to the sensor reaction chamber (11), wherein the capture means each biosensor (40) targets different biomarkers in body fluids.

[0102] In a preferred embodiment, an outlet reservoir (13) is provided. The outlet reservoir is connected to at least one of the sensor reaction chambers (11) to collect the body fluid from the sensor reaction chamber (11). Preferably, each sensor reaction chamber (11) connected sequentially to other sensor reaction chamber (11) by a transfer channel (14) and the last sensor reaction chamber (11) is connected to the outlet reservoir (13) by an outlet channel (15). Alternatively, each sensor reaction chamber (11) may connect to the outlet reservoir (13) by independent outlet channel (15).

[0103] Furthermore, an outlet pad (50) may be positioned in the outlet reservoir (13). In this case the outlet pad (50) is made of absorbent material and absorb the body fluid from the sensor reaction chamber (11). This absorption promotes more flow to pass more sweat through the biosensor (40) and also facilitate disposal of excess sweat from the sensor reaction chamber (H). Moreover, an inlet pad (30) may be positioned in the inlet reservoir (12). The inlet pad (30) is also made of absorbent material and absorb body fluid from sample or skin of the subject. Main function of the inlet pad (30) is to increase the amount of the collected sweat.

[0104] In a preferred embodiment, the biosensor system (1) is designed to contain arrays of interdigitated graphene / metal microelectrodes and a microfluidics system (20), which allows for monitoring of multiple analytes real-time from small sweat volumes as shown in Fig. 1 to 2a. The wearable and flexible multiplexed design carries three independent biosensor (40), each array was made of 18 arc-shaped graphene and gold alternating interdigitated (IDE) microelectrodes that function as working (W) electrode, which is flanked by a reference electrode (42) and a counter electrode (41), on to which sweat is dispensed from integrated microfluidics channel inlets. Here, three electrochemical biosensors (40) enabled multiplexing of three distinct sweat analytes simultaneously. Interfacing of graphene with electronic biosensor (40) arrays promoted the charge transfer ability of sensors to capture sweat biomarkers more sensitively and specifically.

[0105] In a preferred embodiment, the designed biosensor system (1) is made of a flexible polyimide (PI), a polymeric film substrate, which carries dynamic charges with its chemical potential accompanied by those charge carriers from coated graphene. Such substrate not only provided flexibility and large interfacial area in between electrodes of sensor surface, but also facilitated high signal-to-noise ratio, making it most desirable for a typical wearable device.

[0106] The signal transduction mechanism of the biosensor system (1) is based on changes in electrochemical properties, which occur due to change in modulation of graphene’s carrier properties combined with protein / metabolite / ions (sweat analytes) upon interaction with its specific capture means / receptors such as antibodies / enzyme / selective membrane present on the biosensor system (1). The independent and selective operation of the individual biosensors (40) in 3-arrays is preserved during multi-analyte and / or multi-modal capabilities by employing highly specific surface bio- / chemistries and electrically decoupling the operating conditions of each sensor interface. The dynamic range of detection established in this study using the wearable sensor patch for IL-6= 0.1-300 pg / mL (cover typical IL-6 concentration in human sweat), glucose= 10 pM-1 mM (covers the typical range found in human sweat for hypoglycemic, hyperglycemic, and healthy individuals) Ca2+ion = 0.1-3 mM (cover range typically found in human body fluids, including sweat sample) and sufficient to diagnose the disease risk.

[0107] This application also covers production method of the biosensor system (1).

[0108] In this method, the surface of a flexible substrate (10), at least the sensor reaction chamber is coated, preferably by spin coating, with graphene and its derivatives and after that the biosensor (40) is printed / patterned over said coated surface. As a patterning, photolithography, electron beam lithography, high-resolution laser-jet, ink-jet printing, screen printing or roll-to- roll printing can be selected. Finally, each biosensor (40) is bio-functionalized in such a way that provide the capture means for each biosensor (40) to capture targeted biomarkers in body fluids transferred to the sensor reaction chamber (11), wherein the capture means of each biosensor (40) targets different biomarkers from each other.

[0109] In below section, characterizations of production method and test carried out are explained.

[0110] The morphology and elemental analysis of sample was carried out using scanning electron microscope (SEM) and energy dispersive spectroscopy (EDS) with LEO Supra 35 VP scanning electron microscope operated at 3 kV. UV-visible and RAMAN spectroscopy analysis was performed using a NanoDrop spectrophotometer, and Renishaw inVia Reflex Raman Microscope, respectively. Fourier transform infrared spectroscopy (FTIR) measurements were performed using a Nicol et iSlO FTIR Spectrometer (Thermo Scientific, USA) (4 cm-1resolution). X-ray diffraction (XRD) analysis was carried out on Bruker D8 DISCOVER X-ray diffractometer (Bruker AXS GmBH, Karlsruhe, Germany) with Cu target using scintillation counter ( = 1.5406A) at 40 kV in the range of 26 = 10-70°. Thermogravimetric analysis (TGA) of the polyimide (PI) substrate and graphene spin-coated substrates was analyzed by TGA analyzer (NETZSCH STA 449 C thermo-microbalance with TG resolution of 0.1 pg) from 25 to 800°C at 10°C / min under N2 flow of 20 mL / min.

[0111] Design and fabrication of wearable graphene interfaced electrochemical sensor patch was carried out on a flexible polymeric substrate by using standard photolithography in a Class 100 cleanroom facility using polyimide (PI) as substrate using following steps.

[0112] Step 1 : Synthesis of graphene: Graphene oxide (GO) was first generated by chemical exfoliation of graphite to synthesize graphene using a previously described method by chemical reduction [1], Briefly, 3 g of graphite and 15 g of KMnO4 (99.35%) was mixed and carefully added into a mixture of 115 mL of sulphuric acid and nitric acid (1 :3) while continuous stirring on ice-water bath (4 °C) in a span of 2 h interval. The resulting mixture tended to change in its color from black to dark green. This mixture was then kept stirring for 24 h on a magnetic stirrer at a room temperature to obtain purple pasty material. Then 500 mL DI water was added slowly spanning an interval of 90 min at room temperature. Then the beaker content was placed in a water bath at 98 °C for 15 min to complete the reaction, and this was followed by addition of 500 mL of water and 50 mL of H2O2 (30-35%) to the mixture, which turned the color of the solution from dark brown to yellow. The resulting solution was homogenized with proper mixing for 30 min in an ultrasonic bath. The solution was then filtered through a filter paper and the residue left behind on the filter paper was collected. This sample residue was divided into two parts and stirred by adding 500 mL of 2M HC1 for 1 h on a magnetic stirrer followed by washing each part with 500 mL distilled water. The solution was centrifuged at 3500 rpm for 40 min and the supernatant was removed, and thus obtained GO pellet was resuspended in fresh DI water and the above cycle process was repeated several times until the sample showed its pH = 5. The GO sample was then purified by washing twice with ethanol and precipitated. The precipitated GO was again dispersed in DI water in 50 mL centrifuge tubes, frozen at -80 °C then dried at 60 °C for 48 h under vacuum. Finally, the pure GO was characterized by XRD and Raman techniques and stored until use.

[0113] Step 2: Synthesis of graphene from GO: As-synthesized GO from Step (1) was subjected to chemical conversion to graphene by reducing with ascorbic acid as a reducing agent. Briefly, 5 mg GO was added to 50 mL DI water and exfoliation of GO was achieved by ultrasonic dispersion. Then 50 mg of ascorbic acid was added to 50 mL of an aqueous dispersion of the GO under vigorous stirring. The pH of the suspension was adjusted to 10 by adding 3M NH4OH (3 mL) solution dropwise to promote colloidal stability through the electrostatic repulsion. Then the solution was stirred for 4 h at 70 °C that changed the solution color from brown to black. The above mixture was centrifuged at 4500 rpm for 30 min then washed with DI water and the process was repeated thrice to finally obtain the precipitated reduced graphene oxide (rGO), which was dried at 60 °C in an oven.

[0114] As synthesized graphene sample was characterized by XRD and Raman spectroscopy. From the XRD analysis, the absence of peak at 20 ~26.5° and appearance of new peak at 20 ~9.6° indicates the conversion of graphite to GO completely (Fig. 3(a)). As-synthesized GO was subjected to chemical conversion to reduced graphene by treating with ascorbic acid and characterized by Raman spectra (Fig. 3b). Notable differences were observed in Raman shift characteristics, where graphite showed its three characteristic peaks at 2718 cm'1, 1580 cm'1, and 1349 cm'1, corresponding to the 2D, G, and D bands, respectively. The relatively minor intensity of the D peak, indicative of low disorder, suggests minimal defects on the graphite substrate. In contrast, pronounced D peaks observed in synthesized GO and rGO samples suggest significant disorder resulting from vigorous oxidative treatment during the graphite oxidation process. The transformation from GO to rGO was evident in the Raman spectra, with the G band shifting from 1580 cm'1to 1592 cm'1, and a broadened D band at around 1325 cm'1, reflecting reduced in plane sp2domain size due to extensive reduction.

[0115] After successful synthesis of graphene, this graphene was coated on PI substrates by spin coating method as follows: first PI substrate was treated with 0.1% sodium dodecyl benzene sulfonate (SDBS) and 0.1% sodium carbonate (w / w) for 30 min at 100 °C and cleaned with distilled water and dried at 60 °C. The polymeric substrates were further cleaned by UV- Ozone for 30 min using ProCleaner™ Plus. For better adherence of graphene on PI, cleaned PI substrate was immersed into a poly(sodium 4-styrenesulfonate) (PSS) solution (3 mg / mL in 0.5 M NaCl) for 20 min followed by rinsing with DI water and dried under N2 flow. PSS coated PI was then immersed into polyehtyleimine (PEI) solution (3.5 mg / mL in 0.5 M NaCl) for 20 min and then rinsed with DI water and dried under N2 flow. The PSS-PEI treated PI substrate was subjected to spin coating of graphene (0.5 mg / mL) at 3000 rpm, 30 s cycle for five cycles to generate 5 layers of graphene. Graphene coated PI was then subjected to curing at 200 °C in a vacuum oven and the final sample was characterized by XRD, RAMAN, TGA and SEM imaging. Structural features, durability and stability of graphene coated polymeric substrates were carried out by SEM imaging before and after several washing cycles (25-50). For this, each 3x3 cm2graphene coated substrate was washed with a 200 mL of SDBS (0.37%) without optical brightener (WOB) solution in a 500 mL beaker containing 10 steel balls with a diameter of 6 mm at 40 °C for 45 min similar to a previously reported method in accordance with the American Association of Textile Chemists and Colorists (AATCC) testing method 61-2013 [2], After successful confirmation of the durability and stability, the graphene coated polymeric substrate was subjected to fabrication of graphene interfaced electrode by photolithography method and the details are described below:

[0116] Graphene coated flexible PI substrates from the above step 1 and step 2 were subjected to patterning with gold microelectrodes to give rise to graphene interfaced array of sensors (electrical chip-based patch) using standard photolithography method. Here, the substrate surface was first cleaned and dried using N2 gas. An image reversal mask patterning was carried out using a photoresist after layering it on substrate followed by baking process. A 10- 12 nm thin titanium layer was deposited to facilitate improved adhesion of gold and a 150-170 nm thick gold layer deposited using direct current sputter deposition. The deposition was carried out in an argon atmosphere with optimized power (150-200 W) for 1 to 3 min. The metal was lifted off using acetone / ion etching procedure. The width of microelectrodes (graphene / gold) varied between 40-50 pm with same gap between the two electrodes for each sensor (-7 mm x 10 mm area). In the fabricated design of graphene interfaced electrochemical sensor arrays, the mi cro / nano- structured graphene interdigitated electrode-based sensor served as a working and two Au electrodes on either side as counter electrodes and reference electrodes all on PI patch substrate. To interpret the changes in sensor responses, the micro- nano-electrodes were also patterned on bare polymeric substrates with no graphene as controls (control sensor). The fabricated micro / nano-structured graphene interfaced electrochemical sensor arrays were characterized for its integrity and suitability for wearable patch sensor by optical / SEM imaging and photographic images. The mechanical resiliency of the fabricated sensor patch was tested against continuous applied mechanical stresses, such as stretching-relaxing, bending, twisting, and marking indentation that were recorded as photographic images for analysis.

[0117] The present sensor patch design fabrication method is not limited only to a photolithography process, but it can also be fabricated using different methods, including electron beam lithography, high-resolution laser-jet and ink-jet printing, screen printing, and roll-to-roll printing.

[0118] Here, the polymeric film substrate is not only restricted to PI, but also other flexible and stretchable substrates can be used including other polymers, such as polyethylene terephthalate (PET), silicone polymers, polyurethane, or cellulose-based papers. Within the circular design shape of the metal / graphene interdigitated sensor electrodes, the selection of carbon nanomaterials is not limited to graphene, but also its derivatives, including single- / multi-walled carbon nanotubes. The metal microelectrodes are not limited to gold, but also applicable to other metals, such as platinum, palladium, silver, copper, nickel, or aluminum.

[0119] Bio-functionalization of antibody specific to sweat inflammatory / chronic diseases cytokine biomarkers, such as IL-6 was carried out using covalent coupling method through crosslinking IL-6 antibodies with 1 -pyrenebutanoic acid succinimidyl ester (PASE) on graphene and mercaptopropionic acid (MPA) on interdigitated gold electrodes. First, gold electrode arrays were cleaned and treated with 20 mM of MPA in methanol for 2 h to generate free -COOH groups. The activated sensor patch was washed with pure methanol followed by water and dried. Then, graphene electrode arrays were treated with 5 mM of PASE for 2 h to generate a free NHS-ester groups. Freshly prepared 200 mM of l-ethyl-3(3- dimethylaminopropyl) carbodiimide hydrochloride (EDC) and 100 mM of N- hyroxysuccinimide (NHS) rection mixture was incubated on sensor patch. Each sensors from the activated sensor patch was then incubated overnight with 20 pL of anti-IL-6 antibodies (10 pg / mL) at 4 °C. The sensor patch was thoroughly rinsed thrice with PBS solution (pH 7.4). Finally, the antibody immobilized sensor patch was incubated with 5% BSA in PBS solution to block free or unreacted functional groups on sensor patch. Detection of IL-6, its specificity and selectivity of the sensors was carried out within a window of 0.1-300 pg / mL in synthetic sweat. For glucose sweat biomarker sensing, the capture probe IgG monoclonal antibody specific to glucose oxidase enzyme was functionalized on cleaned graphene interdigitated electrodes arrays by chemical cross-linking chemistry using a modified method previously reported [3, 4], First, gold electrode arrays of sensor patch were cleaned and treated with 20 mM of MPA in methanol for 2 h to generate free -COOH group. The activated sensor patch was washed with methanol followed by water and dried. The graphene interdigitated electrodes arrays were treated with 5 mM of PASE. Freshly prepared 200 mM of EDC and 100 mM of NHS reaction mixture was incubated on sensor patch. The PASE and EDC / NHS-activated sensors were coupled with a-glucose oxidase specific IgG antibody (5- 10 pL of 10 pg / mL) and incubated for 15 to 20 min. The electrode arrays were thoroughly washed with aqueous solution. Glucose oxidase (GOx) enzyme solution (5 mg / mL) in PBS (pH 7.4) was then incubated on sensor at 4 °C for 5 h and washed with PBS, pH 7.4 and stored at 4 °C. Electron mediators, such as either 3 mM K3Fe(CN)e or Prussian blue (2.5 mM FeC13, 100 mM KC1, 2.5 mM K3Fe(CN)6 and 100 mM HC1) can be directly mixed with GOx. The antibody immobilized on sensor was confirmed by FTIR analysis. Detection of glucose was optimized to a concentration range 10 pM -1 mM in synthetic sweat and the specificity and selectivity were carried out.

[0120] For selective sweat Ca2+ion sensing, the surface modification of graphene electrochemical sensor was carried out with a calcium ion selective membrane (ISM) cocktail using a simple and easy ion-selective electrode (SC-ISE) solid-contact method. For this, Ca2+ ISM (Ca- ISM) cocktail was prepared using the following composition containing 1.0% (w / w) calcium ionophore N,N,N',N'-tetracyclohexyl-3-oxapentanediamide (ETH 129), 0.2% (w / w) potassium tetrakis[3,5-bis(trifluoromethyl)phenyl]borate (KTFPB), 65.8% (w / w) o- nitrophenyl octyl ether (o-NPOE), and 33.0% (w / w) poly(vinyl chloride) (PVC) dissolved in tetrahydrofuran (THF) solvent [5], For the surface modification of graphene sensor, Ca-ISM, 100 pL of cocktail deposited onto the working graphene electrode arrays surface. The functionalization of Ca-ISM on sensor surface was confirmed by FTIR spectroscopy.

[0121] The cyclic voltammetry (CV) responses of sensor were measured using a handheld potentiostat by applying potential of -0.1 V to +0.1V and measured the signal as a function of current change (pAmp) with synthetic sweat for many cycles. Sensor calibration response was analyzed using several replicates (n = 3~4) and statistical errors were calculated for electrochemical signal generation from sweat sensor patch.

[0122] All experiments were carried out in biological and / or technical replicates (n = 3~4), and the average and standard deviation (SD) values were used for plotting. The inter and intra assay coefficient of variations (%CV) were calculated, and the results that showed below 10% CV were only considered qualified. Long-term stability of sensor patch was evaluated by measuring its electrochemical performance over time (after 1, 4 and 7 days) for stability, reproducibility, and statistical analysis.

[0123] The fabricated sensor patch with a specific concentration of glucose was tested at varying pH of 6.4 to 7.4, and these pH values are close to the physiological pH of sweat samples for variation in pH for glucose sensing.

[0124] The sensors patch was tested in a successive repetition with appropriate concentration / s of selected sweat biomarker / s in artificial sweat and relative standard deviation (RSD%) was calculated for repeatability and reproducibility. The reproducibility of sensor patch performance was evaluated by replicating electrochemical signal measurements (n = 3-4) with different fabricated electrochemical sensors arrays. The results that showed RSD=<5~10% were only regarded as valid data and those exceed over 10% were not considered valid in this work.

[0125] For fabrication of artificial skin and setup for on-artificial skin testing, artificial skin was fabricated by using a previously reported method [6, 7], Briefly, 0.3267 g NaCl and 1.743 g agar mixed well in 56.25 g DI water. Here, NaCl adjusts conductivity and physiological salinity, while agar provides skin-like texture self-shaping and prevents water from evaporating. The mixture was then heated on a hotplate under slow stirring and removed after observing an appropriate increase in viscosity. Following this step, 5.625 g PET powder and 1.382 g TX-151 were mixed and added to maintain the viscosity of agar and PET powder mixture and mixed thoroughly. The mixture is then transferred into a glass petri dish to obtain fabricated skin phantom. Pilot experiments were undertaken for the detection of known specific spiked concentration (0.1-0.6 mM glucose) of model sweat biomarkers on artificial skin setup. The influence of skin temperature on sensor patch for sweat biomarkers detection was tested by measuring CV signal at a potential range -0.1 V to 0.1 V with a scan rate of 50 mVs-1 from 25-35 °C.

[0126] For designing and integration of microfluidics and sensor arrays patch, First, a microfluidic design layout suitable for patterning pchannels on sensor patch was drawn with the help of AutoCAD and printed high-resolution opaque black ink on an acetate mask. This mask was used as a template on silicon wafer for patterning microfluidic channels using SU8-2050 resin using photolithography. About 1.0 mL of SU8-2050 resin was dispensed on silicon substrate and programmed to coat in a two-step process, (i) by spin coating at 500 rpm for 15 s with an acceleration of 100 rpm / s, and (ii) at 1000 rpm for 30 s with an acceleration of 300 rpm / s. The wafer was then subjected to soft baking at 65 °C for 7min followed by 95 °C for 30-45 min (hard baking at 65 °C for 5 min followed by 95 °C for 12-15 min). The wafer was removed and cooled for 10 min at room temperature and the wafer with thin polymer coating was subjected to UV-exposure using a template acetate mask in a mask-aligner (Midas). The UV exposure time, energy / power required was calculated using equation E = time x power, where power = 14.7 J, energy density for 170-225 pm thickness = 260-350 mJ / mc2. This process was optimized to yield microfluidic channel thickness of 170-220 pm. Finally, the SU8 mold was developed in a SU8 developer until clear pchannels appeared on the wafer and washed with isopropanol and dried.

[0127] This mold was placed in a clean Petri dish and the Slygard 184 silicone elastomer and hardener mixture (10: 1 ratio) was poured until the mold was fully submerged. The entire content was placed under a vacuum desiccator to remove all the air bubbles. The Petri dish contents were later placed in a hot-air oven for curing at 75 °C for 1 h. The cured silicone (polydimethylsiloxane, PDMS) layer was peeled out of the mold that left the microfluidic design imprints on PDMS. This layer was subjected to bonding on to a sensor patch made of PI film carrying pre-patterned metal-graphene microelectrodes. Before assembling the microfluidic device on sensor patch, the sensor patch was first affixed on a clean glass support and treated with 15 s Corona plasma treatment on both the surface of patch and silicone imprints that were overlayed and firmly clamped. The entire assembly was immediately placed in an oven overnight at 90 °C for the thermal adhesion fusion bonding between the substrate and microfluidics mold for the adhesion of microfluidics pattern onto the sensor patch.

[0128] To facilitate the absorption of excess sweat, a thin absorbent rayon pad is positioned at the punctured outlet chamber of the microfluidics module where the sample is collected. This accelerates the sweat absorption or the flow of the sweat sample through the sensor patch. Alternatively, if there is collected sweat in the inlet reservoir, it can be manually driven through the sensor by compressing the closed-inlet reservoir chamber with a thumb, allowing the excess sweat to drain thorough the open outlet chamber.

[0129] For estimating microfluidic platform integrated patch sensor performance, First, the sweat flow through the microfluidic channels on the sensors were characterized using PBS solutions. The flow rate experiments were performed using a syringe pump apparatus at a constant sweat rate of 15 pL / min. The thermal adhesion fusion bonding method applied for bonding substrate and microfluidics mold was tested for possible evaporation, which was assessed using a dye fluid in the microfluidic channel at different temperatures (25-40 °C) and relative humidity (RH%) rates over specific time intervals. Here, the degree of evaporation of fluid in the microfluidic channel was determined by the shifting of colored dye in the channel over a specific time interval (over -60 min) and the results were analyzed. For off-body / on-artificial skin sweat sensing analysis of signal-read out and signal processing unit to test microfluidic integrated sensors arrays, Sensor electrodes extending from the patch were connected through a customized adaptor to either a portable pStat-i 400 (Bi) potentiostat (DropSens), or with a portable handheld adaptor unit of a EmStat Pico module (Analog Devices) controlled by a PSTouch Android smartphone application for signal-read out and signal processing as well as to display output on a smartphone for all experiments using fully integrated sensors arrays on flexible patch.

[0130] Sample preparation for fully integrated electrochemical sensing patch testing for targeted sweat analytes / biomarkers and validation is disclosed below:

[0131] Typically, IL-6 concentrations in human sweat varies from 5-15 pg / mL [8, 9], Therefore, IL-6 biomarker concentrations were prepared at a range 0.1-300 pg / mL, which represents below normal to elevated / diseased levels. The samples were tested by serially dispensing in 10- 30 pL volumes onto the sensor patch with increasing concentration gradient from 0.1 pg / mL to 300 pg / mL followed by incubation for 15 to 20 min. Electrochemical parameters of IL-6 binding on sensors were measured and the specificity of sensors was tested with non-specific molecules (negative control / s), such as bovine serum albumin (50 pg / mL) and glucose (50, 100 pg / mL). All the above non-specific molecules were tested on anti-IL6 functionalized graphene sensor surface for specificity performance.

[0132] Artificial sweat samples containing varying glucose concentrations were prepared from lower 10 pM to a maximum 1 mM, which is a typical glucose concentration range reported in human sweat

[0010] , Different concentrations of glucose (in 10-30 pL) were dispensed and incubated for 15 min on the sensor surfaces prior to measuring electrochemical parameters. The selectivity of glucose sensing was carried out by testing them in the presence of other non-specific biomolecules (negative controls), such as fructose and uric acid.

[0133] The graphene electrochemical patch sensor was tested against varying Ca2+ions (0.25-3 mM) in 10 mM acetate buffer (pH 4.6). Here, Ca2+ions concentration range is chosen from 0.1- 1.5 mM and the Ca2+levels body fluids commonly vary between 0.5 mM to 3 mM including sweat samples [11, 12], Kinetic parameters for dynamic electrochemical response with Ca2+were carried by exposing low / high concentrations of Ca2+ions. The interference with other cations were tested, such as by addition of chloride forms of Mg2+(1 mM), and K+ (8 mM) ions. Electrochemical measurements were done within a 1 min waiting period window, and paused while changing solutions.

[0134] The electrochemical kinetic signals / parameters (CV) were measured at potential range -0.1 V to 0.1 V at a scan rate of 50 mVs'1before and after the incubation of known concentrations of model sweat biomarkers (IL-6, glucose and Ca2+) in synthetic sweat on sensor patch. First, the CV response was measured sequentially as follows; (1) testing sensors immobilized with specific receptor elements (anti-IL antibodies, anti-glucose oxidase and Ca-ISM), and (2) sensors incubated with varying levels of sweat biomarkers, specifically IL-6 (0.1-300 pg / mL), glucose (10 pM -1 mM) and Ca2+ions (0.25- 1.5 mM), respectively. The signal responses from sensor patch were measured using a portable Bi-potentiostat. The results of above step (1) were considered as zero-dose / baseline response for determining the sensor signals from specific biomarkers. Sensor responses were calibrated by considering replicate test sample measurements (n = 3~4) along with positive and negative controls and statistically validated the sensor responses.

[0135] The limit of detection (LOD) for each specific lowest sweat biomarker concentration (IL-6, glucose, Ca2+) was calculated using previously reported method

[0013] , The sensor response for spiked sweat biomarkers in synthetic sweat was fitted using polynomial curve fitting and regression analysis coefficient (R2) were determined.

[0136] Validation of model synthetic sweat using gold standard methods is explained below:

[0137] The validation of Ca2+ions in model synthetic sweat using Fluo-4 DirectTM Calcium Assay Kit (Thermo Fisher Scientific) as a gold standard method. Here first, 2X Fluo-4 DirectTM calcium reagent loading solution (50 pL) was added to each well of a black 96-well microplate with transparent bottom and mixed with 50 pL of each test Ca2+ions samples (0.25 to 3 mM) in synthetic sweat that were used for sensor patch testing. The plate was incubated at 37 °C for 60 min under constant shaking and measured the fluorescence using a 485 nm excitation filter and 528 nm emission filter using Synergy HTX-multimode microplate reader (Biotek). For IL-6, an in house developed IL-6 ELISA kit was used as a gold standard assay for validating IL-6 levels in the synthetic sweat used for sensor patch measurements. Validation of patch sensor with synthetic sweat sample was tested using glucometer tests with Accu Chek Performa Nano Roche Glucose Meter (4015630066063) according to the manufacturer’s instructions. Here, synthetic sweat sample containing different concentrations of glucose from 10 pM to 1 mM were prepared. First the glucose strip is inserted in the glucose meter. The edge of the sensor strip was loaded with a 10 pL sample drop and read the glucose level (mg / dL) reading on display. For each tested glucose concentration, a new activation strip was used to read the glucose level.

[0138] Graphene coated on flexible polymeric PI substrate was fabricated as described in experimental above section. The graphene-coated film substrate was characterized by XRD, Raman, TGA analysis and SEM examinations (Fig. 4a-5c). XRD-pattems of bare PI film, graphene nanoplatelets and graphene coated PI substrate are shown in Fig. 4a. A high-intense diffraction peak at 29 = 25.6° and tiny peaks at 43.59°, 53.0° and 63.9° was identified for PI semi-crystalline structure (Fig. 4a). XRD-pattem for pure graphene showed a single highly intense diffraction peak found at 26.15° as 29, with a 3.35 A inter-spacing. This peak corresponds to the crystalline plane (002) of the graphene nanoplatelet. Peaks found at 42.69°, 44.53° and 55.00° as 29 were assigned to the (100), (101) and (004) planes, respectively, which are consistent to a previous study

[0014] , The XRD patterns of graphene coated PI substrate exhibited merging of diffraction peaks, indicating a mixture of diffraction peaks from both PI and pure graphene within the 29 = 20-30° (25.6°, 26.15°). Additional peaks were observed at 29 = 5-60° including smaller peaks at 43.59°, 44.53°, and 63.70°, corresponding to (002) (100), and (101), respectively (Fig. 4a). These results indicate the partial crystallinity of pure graphene and confirm the successful coating of graphene on PI substrate. The intensity of the diffraction pattern from PI substrate is lower than that of coated graphene due to the interference of the PI substrate in the overall XRD data.

[0139] Fig. 4b displays the Raman spectra of PI, graphene, and graphene coated PI. The Raman spectra of graphene exhibited D, G, and 2D bands at 1341 cm-1, 1570 cm-1and 2685 cm-1, respectively. The low intense D band signifies the presence of hybridized carbon atoms with few defects, while the high-intensity G band corresponds to the vibrations of the sp2 hybridized carbon atoms in the basal graphene plane. In the Raman spectra of graphene coated PI substrate, we observed both the D and G bands on PI surface, and the Raman signal from the graphene nanoplatelets encapsulated within PI substrate were easily detectable. Interestingly, the G band of the graphene coated PI showed a slight shift from 1570 cm-1to 1575 cm-1compared to the G band of pure graphene. Additionally, the G band of graphene- coated PI overlapped with the other two bands of PI located at 1510 and 1612 cm'1. These findings demonstrated the successful coating of graphene nanoplatelets on PI substrate. TGA analysis of cleaned PI substrate and graphene coated PI sensor substrates, as shown in Fig. 4c. The results revealed that the residual weight of graphene coated PI was higher than that of control uncoated PI substrate. This increase in residual weight indicates that graphene confers higher thermal stability to the coated PI substrate. This aligns with the well-known property of graphene being thermally stable and resistant to decomposition at high temperatures [2], Hence, the graphene coated PI substrate achieved enhanced thermal stability which enabled thermal treatment (baking) and annealing processes during photolithography patterning of microelectrodes.

[0140] Structural features, durability and stability of graphene coated polymeric substrate was carried out by SEM imaging before and after several washing cycles (25-50) (Fig. 5a-5c). For this, each 3 x 3 cm2graphene coated substrate was washed with a 200 mL of SDBS detergent (0.37%) without optical brightener (WOB) solution. SEM images of graphene coated PI substrate showed partial detachment of non-adhered graphene after 50 cycles of washing. SEM images revealed that graphene coated PI substrate could resist washing for 50 cycles. After successful confirmation of the durability and stability, the graphene coated polymeric substrate was subjected to fabrication of graphene-interfaced microelectrodes by photolithography method in the following section.

[0141] Fabrication of gold interdigitated microelectrodes on graphene coated flexible polymeric film substrates was carried out using standard photolithography technique toward developing sensor patches. To interpret the changes in sensor responses, the micro-nano-electrodes were also patterned on bare polymeric substrates without graphene, serving as control sensors. The fabricated micro / nano-structured graphene-interfaced electrochemical sensor on a flexible polymeric substrate was characterized by optical and SEM imaging (Fig. 6a-6f). The optical images of flexible sensor patch with control gold microelectrodes (without graphene) and graphene interfaced sensor patch are shown in Figure 6a-6d. The control sensor clearly showed interdigitated rings on a smooth polymeric film surface without graphene in between the microelectrodes. Whereas the graphene interfaced electrochemical flexible sensor showed rough surface of the film between the gold microelectrodes (Fig.6a-6b). The dimension of each microelectrode (gold / graphene) measured was <40-60 pm that are separated by 40 pm within a sensor area of ~7 mm x 10 mm. The SEM images of graphene-interfaced electrochemical sensor patch are shown in Fig.6e-6f. In the fabricated design of sensor arrays, the nano-structured graphene (nanosized platelets) homogenously distributed in between the gold interdigitated (GID) microelectrode arrays.

[0142] Referring to Fig. 7a to 7c: The FTIR spectra of flexible PI substrate showed the main five spectral regions of polyimide C-C skeletal modes, such as C-N-C stretching, C-H, carbonyl groups (C=O), and N-H stretching

[0015] (Fig. 7a-7c). The C-N-C vibration featured at 710- 1000 cm'1, while CHz groups emerged between 1000-1500 cm'1and carbonyl groups (C=O) were observed between 1540-1720 cm'1. The C-H stretching vibration of methylene groups were observed at 2865 and 2935 cm'1, while amide group (N-H) appeared at 3311 cm'1. The graphene interfaced flexible sensor showed C-0 and C=O groups of reduced graphene oxide (rGO) at 1070 cm-1and 1760 cm'1, respectively

[0016] (Fig. 7a-7c).

[0143] IL-6 antibody coupling on PASE functionalized graphene interfaced sensor surface was confirmed by the appearance of two new bands of amide-II and amide-I at 1510 cm'1and 1658 cm'1, respectively (Fig. 7a). The first type is the characteristic amide bands that emerged in the region between 1200-1700 cm'1and these bands corresponded to linking segment to amino acids [3], The second type is associated with IL-6 antibody functionalized on graphene interfaced sensor surface, where amide-II at 1510 cm'1corresponded to bending vibration of the N-H bond of the amide group and N-C stretching and C=O bending. The amide-I group emerged at 1658 cm'1that corresponds to C=O stretching mode of amide with contribution of C-C-N and C-N stretching vibrations. The O-H stretching band is derived from the N-H2 bond that appeared at 3380 cm'1(Fig. 7a). The emergence of these characteristics of IL-6 antibody functionalized on graphene interfaced sensor clearly showed coverage of C=O and N-H bonds originating from protein functional groups and confirmed the immobilization of IL-6 specific antibodies on the graphene interfaced sensor surface.

[0144] Glucose oxidase coupling on IgG antibody (specific to glucose oxidase) functionalized graphene sensor surface was confirmed by the appearance of two new bands of amide II and amide I at 1520 cm'1and 1640 cm'1, respectively (Fig. 7b). The FTIR spectra of glucose oxidase (GOx) immobilized on flexible graphene sensor surface showed strong N-H stretching at 3280 cm'1, while the characteristic peaks of N-H in plane bending and C-N stretching modes of the polypeptide chains of glucose oxidase are assigned at 1640 cm'1(amide-I) and amide-II assigned at 1520 cm'1, respectively (Fig. 7b)

[0017] , The slight shift of amide-I and amide-II modes in the glucose oxidase functionalized sensor surface in comparison with only IgG antibody surface was attributed to the intermolecular interaction between glucose oxidase enzyme and IgG antibody, which favor the stability of glucose oxidase on the graphene interfaced sensor surface.

[0145] The FTIR spectra of surface modified graphene interfaced electrode with Ca-ISM cocktail is shown in Fig. 7c. The calcium ionophore bands were observed at 3321 cm'1(N-H, secondary amine), 2925 cm'1(O-H stretching), 1647 cm'1(C=O stretching) and 1000 cm'1(C-0 stretching). The KTFPB, o-NPOE, and PVC related bands were observed at 2867 cm-1 (CEE stretching), 1620 cm'1and 1540 cm'1(N-H primary amine), 1456 cm'1(C-C stretching) and 1000 cm'1(C-Cl stretching). The presence of these functional groups confirmed the functionalization of Ca-ISM sensor electrode.

[0146] Referring to Fig. 8a to 8c: The sensor patch pre-activated with receptors for IL6, glucose oxidase or Ca2+-specific membrane was tested with synthetic sweat samples for stability and reproducibility in its CV response signals (Fig. 8a-c). After 6 successive cycles of measurements, the sensor responses against specific concentration of each of the analytes tested (eg., 0 pg IL-6, 0.4 mM glucose, and 0.75 mM Ca2+) was recorded that showed consistent and reproducible signals (Figs. 8a(i), 8b(i), 8c(i)). Each of these sensors was further subjected to testing after 7-days of storage at standard 4 °C temperature and the results are shown in (Figs. 8a(ii), 8b(ii), 8c(ii)). We observed that storage conditions did not affect the sensor patch responses indicating that the sensors have shelf-life of at least over a week under ideal 4 °C storage conditions. Further, fabrication processes, such as batch-to-batch variations in fabricated sensor patch responses were tested to ascertain for any unintended sensor patch behavior. Here, three independent sensor patches fabricated from three different batches were tested in triplicates for their electrochemical responses (designated as S1-S3) under optimized conditions as described above with each type of sensors against IL-6, glucose, and Ca2+in synthetic sweat sample, respectively (Figs. 8a(iii), 8b(iii), 8c(iii)). The sensor patch exhibited consistent current ( Al) responses as shown for IL-6 and glucose, respectively. However, batch-to-batch variation in the sensor responses was prone to occur with Ca2+ion sweat sensors, mostly due to the inconsistent thickness of the ion-membrane layer formed on sensors during pre-activation process, which can be acceptable because this variation occurred whose percent relative standard deviation (%RSD) was less than 10%. Referring to Fig. 9a and Fig. 9b; The electrochemical graphene interfaced electrode sensor patch functionalized with specific receptor elements, such as anti-glucose oxidase receptors was fixed on an artificial skin set up. An absorbent pad was fixed on artificial skin to accelerate the sweat access to the sensor surface. Electrochemical signal form the device measured against a single concentration of glucose (0.1 mM) at three distinct temperatures (25 °C, 30 °C and 35 °C) at pH 6.0 is as showing in Fig. 9a. The peak current values against ±5 °C increase in temperature did show a slight increase in electrochemical response curves, where sensor patch exhibited a mild increase in the peak current response at 35 °C as compared to the sensor patch tested at a room temperature (25 °C) (Fig. 9a). The fabricated sensor patch at a specific concentration of glucose as a modal sweat biomarker at varying pH between 6.4 to 7.4 was tested against same 0.1 mM glucose concentration, and these pH values are within the physiologically varying pH levels of the sweat. Therefore, effect of pH on the electrochemical performance of the fabricated graphene interfaced electrochemical patch was measured in artificial sweat spiked with glucose (0.1 mM) at varying pH ranging from 6.4 to 7.4 (Fig. 9b). The peak current values showed minor changes with pH 7 and 7.4. However, at lower pH values of <6.4, increased cathodic peak current was observed, which confirmed that the GOx functionalized sensor responses has sensitivity towards acidic pH value (Fig. 9b).

[0147] The above results indicate that the current responses of fabricated flexible sensor patches were found to be stable at both at room temperature and above room temperatures, which is most desired for electrochemical sensing of sweat biomarkers and for real-time glucose monitoring applications.

[0148] Referring to Fig. 10; The sweat flow through the microfluidic channels on the sensors was first tested using physiologically relevant solution (PBS pH 7.4) at different flow rates (15-35 pL / min). This allowed understanding the risks of leakage, microfluidics channel clogging with bubbles, molecules or salts form PBS through micro channels under controlled conditions using a syringe pump. Fig. 10 shows the measured flow-rate and the time taken to saturate with fluid in the pchannels that was dependent on time. The loss of fluid sample imbalances optimal local temperature and relative humidity in sensor patch system that may affect the overall sensor patch system performance. Therefore, the effect of local temperature and rate of evaporation was tested by passing a rhodamine dye fluid into the microfluidic system that was placed at different temperatures (25-40 °C) and relative humidity (RH%) are determined.

[0149] The degree of evaporation of fluid in the microfluidic channel was determined by observing for any shift in the level of a colored dye in the channels over a specific time interval (0-1 h). It was observed that there was no significant loss of fluid in the microfluidics system at room temperature. However, the fluid loss tended to occur only when the temperature rises to above 40 °C. Further, the fully integrated sensors with microfluidic platform were analyzed for determining the rate of evaporation of dyed fluid. These experiments demonstrated that there was no significant dye fluid evaporation occurred at different time intervals tested at varying physiologically relevant temperatures (25-40 °C), suggesting that the microfluidic integrated sensor patch meets the requirement for testing with sweat samples.

[0150] A series of standard sweat biomarkers were spiked in synthetic sweat, such as IL-6 (0.1-300 pg / mL), glucose (10 pM -1 mM) and Ca2+ions (0.25- 3 mM) as described in the experimental section. Before sensor patch response, these standard samples were subjected to confirmation for their concentrations using respective gold-standard tests, such as ELISA test for IL-6, Folin Wu method by copper reduction for glucose, and Fluorescence based Ca2+ion specific assay kit for measuring Ca2+ion. The samples that were qualified using gold-standard methods were then used for testing with fabricated sweat sensor patch. Each of these tests were carried out using independent replicates using different sensor patch devices (n = 3). Each sensor patch device contained 3 separate sensors arranged in an array each sensor is dedicated to detecting a specific biomarker analyte on a single platform.

[0151] The flexible graphene-interfaced IL-6 sweat sensor exhibited a concentration-dependent response to applied electric potentials ranging from -0.1 V to +0.1 V (refer to Fig. I la). Values extrapolated from cyclic voltammogram at a specific potential V displayed a linear curve demonstrating the increasing signal responses of sensor patch to increasing levels of IL- 6 in sweat (Fig. 1 lb). These findings highlight the remarkably sensitive cyclic voltammetry (CV) signals originating from the sensors patch, specifically in relation to IL-6 concentrations within the range of 0.1 pg / mL and 300 pg / mL (typical IL-6 concentration in human sweat). The linear response of IL-6 sweat sensor patch with increasing concentration of 0.1-300 pg / mL, and this range is a representative of typical IL-6 concentrations found in human sweat. The limit-of-detection (LoD) as determined from the linear relationship was 10 pg / mL. The sweat sensor patch showed selective or discerning response to IL-6 in presence of other chemically and biologically relevant compounds in sweat, such as glucose and BSA (Fig. 11c).

[0152] Fig. l ld-f display the electrochemical signals acquired from interdigitated electrodes integrated with graphene and immobilized with GOx (glucose oxidase) to create a glucose sweat sensor patch. These sensor patch respond to varying glucose concentrations spanning 10 pM to 1 mM. The selected glucose concentration range effectively encompasses levels typically observed in human sweat, encompassing hypoglycemic, hyperglycemic, and healthy ranges. The recorded monitoring responses of sweat patch for glucose exhibit a linear relationship characterized by an incremental change in sensor response signal until reaching to a saturation point at 1 mM glucose as depicted in Figs, l ld-e. Employing linear regression analysis, the LoD extrapolated from the data was determined to be 6 pM glucose. Fig. I lf underscores the versatile capability of the flexible sensor patch to discriminate glucose (Glu) from other prominent bio / chemical compounds that may be present in sweat, such as uric acid (UA), as well as a structurally analogous compound, fructose (FRA).

[0153] Fig. l lg-i shows the electrochemical responses to Ca2+ions in synthetic sweat using sweat sensor patch preactivated with Ca-ISM. The electrochemical signals from the sensor patch against varying concentrations of Ca2+ ions ranging from 0.25 mM to 3 mM, was measured at applied potentials from -0.1 V to +0.1 V and the results are shown in Fig. 11g. The tested Ca2+ion levels were within the normal range for human body fluids, including sweat. The linear calibration curve, shown in Fig. l lh, demonstrates concentration-dependent increasing signals that correspond to the Ca2+ions present in the test samples (Fig. l lg-h). The LoD as calculated through the linear regression analysis was found to be 26 pM Ca2+ion. Fig. Hi shows the selective response of the Ca2+ion sensor patch to Ca2+ions, even in the presence of the of other common sweat ion species such as Mg2+and K+.

[0154] The outcomes derived from the responses of the sensor patch distinctly demonstrate the establishment of minimum detectable or interaction levels for sweat biomarkers on the wearable sensor patch when operating under optimized conditions.

[0155] Sensor patch assay validation was carried out with gold standard assays such as commercial glucometer for validating glucose response and determining Ca2+ions using Fluo-4 DirectTM Calcium Assay Kit as gold standard method. Here, same synthetic sweat aliquots samples spiked with known concentrations and acceptable ranges of a specific sweat, glucose 10 pM - 1 mM and Ca2+ions 0.25- 1.5 mM) used for testing sensor patch was applied for validation.

[0156] The synthetic sweat samples containing different concentrations of spiked IL-6 protein (0.1 - 300 pg / mL), glucose (10 pM -1 mM) and Ca2+ions (0.25- 1.5 mM) were tested using standard assay, and the results of these assays were compared with biosensor patch device responses. Fig. 12a-c shows linear regression plots using data obtained from spiked IL-6 protein, glucose and Ca2+ions in synthetic sweat. These linear responses demonstrated that the biosensor patch signals were comparable to standard in vitro assays. The sensor patch response showed consistent values for IL-6 protein that were comparable to the IL-6 protein levels determined using gold standard IL-6 ELISA tests (Fig. 12a). The flexible biosensor patch device responses against micro molar (pM) range levels of glucose, however commercial glucose meter responded from 0.6 mM (millimolar range) of glucose levels (Fig. 12b). This clearly indicated that the developed flexible biosensor patch is sensitive towards the detection of pM levels of glucose in sweat. The commercial glucose meter is designed to detect the glucose levels in blood, where blood glucose is present in millimolar levels and therefore commercial glucose sensor (glucometer) did not respond to pM levels of glucose in sweat. Similarly, gold standard assay results for the detection of Ca2+ions (0.25- 1.5 mM) were comparable with biosensor patch device responses (Fig.12c). It was observed that the higher levels of Ca2+ions (above 1 mM), the gold standard assay response for this analyte was saturated, and it is challenging to accurately quantify possibly due to the handling of reaction mixture and reaction time and environmental conditions. We noticed that the quantitative analysis of sweat biomarkers in synthetic sweat developed flexible biosensor patch device was more suitable, rapid, sensitive and generate distinguish signal compared to the standard assays (Table 1), and gold standard assays lack in options for lower levels detection of sweat biomarker signal (glucose) in given settings.

[0157] Analyte Linear regression analysis

[0158] Detection range

[0159] Glucose Gold standard assay 0.6-2 mM

[0160] Wearable sensor patch signal 0.01-1.0 mM

[0161] Ca2+ionsGold standard assay 0.25-1.5 mM

[0162] Wearable sensor patch signal 0.25-3.0 mM

[0163] IL-6 Gold standard assay 10-400 pg / mL

[0164] Wearable sensor patch signal 10-300 pg / mL

[0165] Table 1. Summary of detection range of sweat biomarkers in synthetic sweat samples by gold standard tests and comparison of detection carried out using flexible biosensor patch device using same samples.

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Claims

CLAIMS1. A biosensor system (1) for monitoring multiple biomarkers in body fluids, characterized by- A flexible substrate (10) having at least one sensor reaction chamber (11), an inlet reservoir (12) for receiving the body fluids,- A microfluidic system (20) enables to transfer the body fluid from the inlet reservoir (30) to the sensor reaction chamber (11) by capillary action of fluids,- Interdigitated biosensors (40) patterned in the sensor reaction chambers (11) having,A working electrode (43) comprises concentrically provided multiple conductive arcs and a nanostructure interfaced between the arcs wherein, the nanostructure is made of graphene or its derivatives andA conductive counter electrode (41) and a conductive reference electrode (42) which flanks the working electrode (43) andCapture means for each biosensor (40) to capture targeted biomarkers in body fluids transferred to the sensor reaction chamber (11), wherein the capture means of each biosensor (40) targets different biomarkers from each other.

2. A biosensor system (1) according to Claim 1, characterized by further comprising an outlet reservoir (13) connected to at least one of the sensor reaction chambers (11) to receive the body fluid from the sensor reaction chamber (11).

3. A biosensor system (1) according to Claim 2, characterized by further comprising an outlet pad (50) positioned in the outlet reservoir (13) to absorb the body fluid from the sensor reaction chambers (11).

4. A biosensor system (1) according to Claim 1, characterized by further comprising an inlet pad (30) positioned in the inlet reservoir (12) to absorb the body fluid.

5. A biosensor system (1) according to Claim 1, characterized by the microfluidic system (20) comprises a main channel (21) connected to the inlet reservoir (12) and reservoir channels (22) connected between the main channel (21) and the sensor reaction chambers (11).

6. A biosensor system (1) according to Claim 5, characterized by the diameter of the main channel (21) is bigger than the diameter of the reservoir channels (22).

7. A biosensor system (1) according to Claim 1, characterized by at least two sensor reaction chambers (11).

8. A biosensor system (1) according to Claim 1, characterized by at least three sensor reaction chambers (11).

9. A biosensor system (1) according to Claim 1,7 or 8, wherein the capture means is selected between anti-IL-6 antibodies, glucose oxidase enzyme probe and Calcium ion selective membrane.

10. A biosensor system (1) according to Claim 1 wherein by the counter electrode (41) and the reference electrode (42) are in form of crescent.

11. A biosensor system (1) according to Claim 1, characterized by the working electrode (43) or the counter electrode (41) or the reference electrode (42) are made of gold.

12. A biosensor system (1) according to Claim 1, characterized by the working electrode (43) or the counter electrode (41) or the reference electrode (42) are made of platinum, palladium, silver, copper, nickel or aluminum.

13. A biosensor system (1) according to Claim 1, characterized by further comprising a connection element for sending output data based on the monitoring.

14. A biosensor system (1) according to Claim 1314, characterized by further comprising a potentiostat and connection element is configured to connect potentiostat.

15. A biosensor system (1) according to Claim 14, characterized by a connection element for an adapter which transfers data to a mobile device having a wireless communication module.

16. A biosensor system (1) according to Claim 15, wherein the adaptor is configured to connect to the mobile device by a wired or a wireless communication module.

17. A biosensor system (1) according to Claim 1, characterized by the flexible substrate (10) is made of polyimide, polyethylene terephthalate (PET), silicone polymers, polyurethane or cellulose-based papers.

18. A biosensor system (1) according to Claim 1, wherein the nanostructure is made of single-, or multi-walled carbon nanotubes.

19. A biosensor system (1) according to Claim 1, wherein the microfluidic system (20) and the biosensor configured to transfer and monitoring sweat, respectively.

20. A biosensor system (1) any of preceding claims, characterized by being wearable.

21. A biosensor system (1) according to Claim 20, characterized by adhesive surface to stick skin of the user.

22. A production method for the biosensor system (1) of Claim 1, characterized byCoating at least part of the sensor reaction chamber (11) of the flexible substrate (10) by graphene or its derivatives,Patterning the counter electrode (41), the reference electrode (42) and working electrode (43) of the biosensor (40) on nanostructure coated part,Bio-functionalizing each the biosensor (40) in such a way that providing the capture means for each biosensor (40) to capture targeted biomarkers in body fluids transferred to the sensor reaction chamber (11), wherein the capture means of each biosensor (40) targets different biomarkers from each other23. A method according to Claim 22, characterized by the coating by spin coating.

24. A method according to Claim 22, characterized by the patterning by photolithography, electron beam lithography, high-resolution laser-jet, ink-jet printing, screen printing or roll-to-roll printing.