Wearable biosensors and applications thereof

Conformable and wearable sensors with integrated on-chip gate for the detection of biomolecules, chemicals, and other substrates and applications thereof are provided. Biosensor chips can be built with In2O3 nanoribbon field-effect transistors. Biosensor chips can conform to features of a human body, enabling ability for individuals to wear a biosensor.

TECHNICAL FIELD

The invention is generally directed to biosensors, and more specifically to wearable biosensors with integrated on-chip gate electrodes.

BACKGROUND

Accordingly, while many commercially available wearable electronics can track users' physical activities, devices that can provide an insightful view of user's health status at molecular level need more development.

SUMMARY

Many aspects of the present invention are directed to wearable biosensors with integrated on-chip gate. More aspects are directed to highly sensitive In2O3nanoribbon transistor biosensors capable of monitoring glucose.

Several aspects are directed to a biosensor chip that a flexible substrate. The biosensor also has at least one parallel pair of flexible extended field effect transistors deposited onto the chip substrate such that each field effect transistor has a source terminal and a drain terminal. The biosensor also has a malleable gate electrode deposited onto the chip substrate for every pair of field effect transistors such that each gate electrode is disposed halfway between and in parallel with each pair of field effect transistors. The biosensor also has a pair of malleable source electrodes deposited onto the chip substrate for every pair of field effect transistors such that the each electrode of each pair of source electrodes is in contact with a source terminal of each field effect transistors of each pair of field effect transistors. The biosensor also has a pair of drain electrodes deposited onto the chip substrate for every pair of field effect transistors such that each electrode of each pair of drain electrodes is in contact with a drain of each field effect transistors of each pair of field effect transistors.

In more aspects, the biosensor also has a glucose oxidase deposited on each source and drain electrode.

In further aspects, the biosensor also has chitosan and single-walled carbon nanotubes deposited with the glucose oxidase on each source and drain electrode.

In even more aspects, the biosensor is capable of detecting glucose in an external body fluid.

In even further aspects, the external body fluid is a fluid selected from the group consisting of sweat, tears, and saliva.

In even further more aspects, the biosensor is able to detect glucose concentrations between 10 nM to 1 mM in a solvent.

In even further more aspects, the biosensor is conformable to a human feature.

In even further more aspects, the biosensor is conformable to human skin.

In even further more aspects, the biosensor is integrated into a skin patch.

In even further more aspects, the biosensor is integrated into a watch.

In even further more aspects, the biosensor is conformable to a human eye.

In even further more aspects, the biosensor is integrated into a contact lens.

In even further more aspects, the biosensor also has a third and fourth extended field effect transistor for each pair of field effect transistors deposited onto the chip substrate such that each third and fourth field effect transistor of each pair of each field effect transistors flank their respective pair of field effect transistors, each on one outer side and situated in parallel to their respective pair of field effect transistors.

In even further more aspects, the substrate is composed of polyethylene terephthalate (PET).

In even further more aspects, the field effect transistor is composed of indium oxide (In2O3).

In even further more aspects, the gate electrode is composed of gold.

In even further more aspects, the source and drain electrodes are composed of gold.

DETAILED DESCRIPTION

The term “substantially,” “generally,” or “about” may be used herein to describe disclosed or claimed embodiments. The term “substantially” may modify a value or relative characteristic disclosed or claimed in the present disclosure. In such instances, “substantially” may signify that the value or relative characteristic it modifies is within ±0%, 0.1%, 0.5%, 1%, 2%, 3%, 4%, 5% or 10% of the value or relative characteristic.

It should also be appreciated that integer ranges explicitly include all intervening integers. For example, the integer range 1-10 explicitly includes 1, 2, 3, 4, 5, 6, 7, 8, 9, and 10. Similarly, the range 1 to 100 includes 1, 2, 3, 4 . . . 97, 98, 99, 100. Similarly, when any range is called for, intervening numbers that are increments of the difference between the upper limit and the lower limit divided by 10 can be taken as alternative upper or lower limits. For example, if the range is 1.1. to 2.1 the following numbers 1.2, 1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.9, and 2.0 can be selected as lower or upper limits. In the specific examples set forth herein, concentrations, temperature, and reaction conditions (e.g. pressure, pH, etc.) can be practiced with plus or minus 50 percent of the values indicated rounded to three significant figures. In a refinement, concentrations, temperature, and reaction conditions (e.g., pressure, pH, etc.) can be practiced with plus or minus 30 percent of the values indicated rounded to three significant figures of the value provided in the examples. In another refinement, concentrations, temperature, and reaction conditions (e.g., pH, etc.) can be practiced with plus or minus 10 percent of the values indicated rounded to three significant figures of the value provided in the examples.

The prefix “nano” as used herein means that the structures described as such have at least one dimension form about 1 to 100 nm. (e.g., at least one dimension less than 100 nm).

In several embodiments, biosensors are capable of continuous analyte monitoring over a period time. In more embodiments, biosensors detect analytes in an external body fluid. In even more embodiments, biosensors detect analytes in sweat, tears, or saliva. In further embodiments, biosensors do not require breaking of human skin to detect an analyte. In some more embodiments, biosensors are capable of detecting and monitoring glucose levels of an individual. In even more embodiments, biosensors are capable of detection of soluble glucose concentrations between 10 nM to 1 mM.

Embodiments are also directed to flexible, conformable, and wearable body sensors that are fully integrated. Accordingly, in many embodiments, biosensors are manufactured onto a chip composed of flexible material. In some of these embodiments, the flexible material is polyethylene terephthalate (PET). In more embodiments, all electrodes and transistors are deposited directly onto a chip. In even more embodiments, the electrodes and transistors are composed of flexible materials. In some embodiments, electrodes are composed of gold. In some more embodiments, field effect transistors are composed of indium oxide (In2O3). In some particular embodiments, gold gate electrodes are deposited directly onto the chip. In several more embodiments, a biosensor is conformable to a human feature such as human skin or eye. In many further embodiments, a biosensor is integrated into a wearable device, such as a patch designed to adhere to skin, a watch, or a contact lens.

More embodiments are directed to a biosensor design. Accordingly, in several embodiments, biosensors will have a chip substrate with at least one pair or group of 4 of flexible field effect transistors (FET) deposited thereon. In many of these embodiments, each FET will have a source terminal and a drain terminal. In more embodiments, biosensors will have a chip with flexible electrodes deposited thereon. In some embodiments, biosensors will have at least one source electrode, at least one terminal electrode, and at least one gate electrode. In various embodiments, biosensors will have a gate electrode deposited on a chip halfway in between a pair of FETs deposited on said chip. In even more embodiments, biosensors will have a source electrode deposited on a chip that contacts a source terminal of a FET also deposited on said chip. In even further embodiments, biosensors will have a drain electrode deposited on a chip that contacts a drain terminal of a FET also deposited on said chip.

A common problem with classic commercial hand-held analyzers for the detection of glucose or lactate is that most of these devices require blood samples, often necessitating a finger prick or invasive sensor (e.g., needle embedded under skin) (A. J. Bandodkar and J. Wang, Trends Biotechnol., 2014, 32, 363-371, the disclosure of which is herein incorporated by reference). Accordingly, these classical detection devices are undesirable by consumers. Wearable biosensors offer a potential alternative, as they can perform continuous analyte monitoring without undesirable breaking of skin. Continuous analyte monitoring can provide great benefit, considering, for example, optimum diabetes management is best performed with regular glucose monitoring, and glucose level trends are more insightful than temporally sparse collections of data points (P. Makaram, D. Owens, and J. Aceros, Diagnostics, 2014, 4, 27-46, the disclosure of which is herein incorporated by reference). Many other medical real-time detection systems would benefit from wearable biosensors, such as, for example, detection of pathogens to alert onset of pathogenic diseases (M. S. Mannoor, Nat. Commun. 2012, 3, 763, the disclosure of which is herein incorporated by reference).

Although blood is by far the most studied and utilized sample for diagnosis, other biological fluids such as sweat, tears, and saliva, which are more readily accessible, also contain numerous biochemical analytes that can provide valuable analysis (P. Makaram, D. Owens, and J. Aceros, 2017, 2014 cited supra; and C. Liao, Adv. Mater. 2015, 27, 676-681, the disclosure of which is herein incorporated by reference). Although various recent studies suggest a diagnosis system based on the glucose concentration in external body fluids, many challenges still exist (O. Veiseh, et al., Nat. Rev. Drug Descov. 2015, 14, 45-57; O. Olarte, et al., Conf. Proc. IEEE Eng. Med. Biol. Soc. 2013, 2013, 1462-1465; and Q. Yan, et al., Anal. Chem. 2011, 83, 8341-8346; the disclosures of which are herein incorporated by reference), many challenges still exist for the accurate detection (S. K. Vashist, Anal. Chim. acta 2012, 750, 16-27; and M. Tierney, Electroanalysis 2000, 12, 666-671; the disclosures of which are herein incorporated by reference). For example, glucose concentrations in external body fluids are much lower compared with blood (J. Moyer, et al., Diabetes Technol. Ther. 2012, 14, 398-402, the disclosure of which is herein incorporated by reference). Body fluid sensing results can also be negatively affected by ambient temperature changes, mechanical deformation caused by body motion, and the sample collection procedure.

Typical FET-based biosensor platforms have individual sensors with an external Ag/AgCl solution gate electrode, which is used to set the operational point of the sensors to the optimal detection mode. The Ag/AgCl electrode is commonly used as the reference electrode in the electrochemical measurements and biosensing applications due to its ability to provide stable potential and read signals precisely. Integration of Ag/AgCl electrodes into a biosensor chip, however, remains challenging. A stand-alone fully integrated sensor array, in accordance with numerous embodiments, is better suited to build a wearable biosensor platform. Accordingly, in some embodiments, FET-based biosensors are used, wherein the gate electrode only needs to supply stable gate bias to the devices, which can be achieved by an on-chip gate electrode. The source-drain electrodes and the on-chip gate electrodes, in many embodiments, are incorporated into the straightforward 2-step shadow mask fabrication process so that no additional fabrication steps are required. In several embodiments, the integration of In2O3 glucose sensors with wearable electronics generates high impact for diabetes monitoring. In more embodiments, the development of wearable sensors for in-situ, real-time, and low-cost detection of biologically and medically important targets will generate broad impact in many applications involving electronic skin (K. Takei, et al., Nat. Mater. 2010, 9, 821-826, the disclosure of which is herein incorporated by reference), thermal regulation (T. Someya, Proc. Natl. Acad. Sci. U.S.A. 2005, 102, 12321-12325, the disclosure of which is herein incorporated by reference), chemical sensing (A. N. Abbas, et al., ACS Nano 2014, 8, 1538-1546, the disclosure of which is herein incorporated by reference), and the detection of pathogens in body fluids (M. S. Mannoor, et al., 2012, cited supra).

In a number of embodiments, highly sensitive and conformal In2O3nanoribbon FET biosensors with fully integrated on-chip gold gate are described herein, which have been laminated onto various surfaces, such as artificial arms and watches, and have enabled glucose detection in various body fluids, such as sweat and saliva. Many devices, according to various embodiments, are fabricated through two shadow masks. In some embodiments, a first shadow mask is used to define the sputter-coating of In2O3nanoribbons, and a second shadow mask is used for metal deposition of the source, drain and gate. In more embodiments, the source and drain electrodes are modified with the enzyme glucose oxidase (GOx), biocompatible polymer chitosan, and single-walled carbon nanotubes (SWCNT) using ink-jet printing. Gold gated In2O3FETs, in accordance with many embodiments, provide good electrical performance on highly flexible substrates. In even more embodiments, the optimized glucose sensors deliver very wide detection ranges and high sensitivity, spanning at least 5 orders of magnitude and detection limits down to 10 nM. In some more embodiments, the non-invasive sensors are capable of glucose detection in external human body fluids, such as tears and sweat, which is demonstrated on artificial skin and eye replicas in exemplary embodiments below. Accordingly, embodiments of glucose detection platforms as described herein, are highly sensitive for glucose detections and also have many other sensing applications, including, but not limited to, detection of pathogens, chemicals, biologics, and other analytes found in body fluid.

With reference toFIGS.1A,1B,1C,1D, and1E, schematic illustrations of a biosensor having at least one pair of extended field effect transistors are provided. Biosensor10includes a flexible substrate12and at least one field effect transistor assembly13. In a refinement, flexible substrate12is composed of a plastic such as polyethylene terephthalate (PET). Sometimes, flexible substrate12is referred to as a “chip substrate.” Biosensor10includes at least one pair14of flexible extended field effect transistors deposited onto the flexible substrate. Each pair14of flexible extended field effect transistors include a first electrode assembly16and a second electrode assembly18. First electrode assembly16including a first source electrode22, a first drain electrode24, and a first metal oxide channel26. First metal oxide channel26is disposed over and typically contacts flexible substrate14. Mover, first metal oxide channel26contacts first source electrode22and the first drain electrode24. Second electrode assembly18includes second source electrode32, a second drain electrode34, and a second metal oxide channel36. Second metal oxide channel36is disposed over and typically contacts flexible substrate14. Second metal oxide channel36contacts second source electrode32and second drain electrode34. Malleable gate electrode40is disposed over and typically contacts flexible substrate12. Characteristically, malleable gate electrode40is interposed between first electrode assembly16and the second electrode assembly18. In a variation, first metal oxide channel26and second metal oxide channel each independently are composed of or comprise an indium oxide (e.g., In2O3). In a refinement, malleable gate electrode40is composed of or comprises a metal, and in particular a platinum group metal such as gold and platinum. In a further refinement, a first source electrode22, a first drain electrode24, second source electrode32, and second drain electrode34are composed of or comprise an indium oxide (e.g., In2O3). In a refinement, malleable gate electrode40is composed of or comprises a metal, and in particular a platinum group metal such as gold and platinum.

Advantageously, biosensor10can be conformable to a human feature, and in particular, human skin. For example, biosensor10can be integrated into a skin patch and into a watch. When biosensor10is conformable to a human eye, the biosensor can be integrated into a contact lens.

In a variation, first metal oxide channel26includes a first ribbon section44having a first length and a first width where the first length being greater than the first width. Moreover, the first metal oxide channel26defines a first axis a1which is a centerline through the first ribbon section. Similarly, second metal oxide channel36includes second ribbon section46having a second length and a second width where the second length being greater than the second width Second metal oxide channel36defines a second axis a2which is a centerline through the second ribbon section. In a refinement, first axis a1is substantially parallel to the second axis a2. In a further refinement, malleable gate electrode40has a rectangular cross-section that defines a third axis a2through a centerline that portion of the malleable gate electrode40that is substantially parallel to the first axis and second axis.

In a refinement, each of first metal oxide channel26and second metal oxide channel36have a length l of 200 to 800 μm for their respective ribbon sections, a width w of 10 to 50 μm, and a thickness t of 5 to 25 nm. In further refinement, malleable gate electrode40has a width from 10 to 50 μm and a thickness from 5 to 100 nm. In still further refinements, first source electrode22, first drain electrode24, second source electrode32, and second drain electrode34each independently have a length l1from 50 to 200 μm, width w1from 50 to 200 μm, and a thickness from 5 to 100 nm. Typically, malleable gate electrode40is separated from each of the ribbon sections of first metal oxide channel26and second metal oxide channel36by a distance from about 150 to 600 μm from the axis a1or a2to axis a3.

In another variation as depicted inFIG.1B, biosensor10further includes third electrode assembly50and fourth electrode assembly52that flank pair14of flexible extended field effect transistors. In this regard, third electrode assembly50is positioned such that first electrode assembly16is positioned between the malleable gate electrode40and third electrode assembly50. Similarly, fourth electrode assembly52is positioned such that second electrode assembly18is positioned between malleable gate electrode40and fourth electrode assembly52.

With reference toFIG.1C, at least one of, and typically all of, first source electrode22, first drain electrode24, second source electrode32, and second drain electrode34can be overcoated with layers60including glucose oxidase. In a refinement, first source electrode22, first drain electrode24, second source electrode32, and second drain electrode34can be overcoated with layers60including chitosan and single-walled carbon nanotube. In another refinement, at least one of, and typically all of, first source electrode22, first drain electrode24, second source electrode32, and second drain electrode34is overcoated with layers60of chitosan, single-walled carbon nanotube and optionally glucose oxidase. In a further refinement, biosensor10is capable of detecting glucose in an external body fluid (e.g., sweat, tears, and saliva). In this regard, the biosensor is able to detect glucose concentrations between 10 nM to 1 mM in a solvent.

With reference toFIG.1D, biosensor10can also include microwell64that allows collection and accumulation of body fluids. Typically, microwell64can also serve as a passivation layer to ensure reliable sensing without electrical disturbance that can be introduced by contacting of metal lines with a body and/or body fluids. Typically, microwell64can be formed from a plastic, rubber, silicone or the like. In this regard, polydimethylsiloxane is found to be particularly useful.

With reference toFIG.1D, biosensor10can include one or more additional pairs of flexible of the field effect transistor assemblies13as set forth above. In a refinement, biosensor10includes from 2 to 50 of field effect transistor assemblies13.

In some variations, In2O3nanoribbon devices are fabricated similarly to previously reported shadow mask fabrication technique (Q. Liu, et al, 2016, cited supra), however, side gate patterns were added to the source/drain shadow mask and also a 5 μm ultra-flexible PET substrate was used.FIG.2Billustrates, in accordance with various embodiments, a scheme for fabricating flexible In2O3macroelectronics on PET substrates. As shown, a PET substrate is attached to the first shadow mask using antistatic tape. Then radio frequency (RF) sputtering was used to deposit 16-nm-thick In2O3nanoribbons through the openings on the shadow mask. The second shadow mask was then laminated onto the PET substrate to add a subsequent metal deposition. After using a single mask to define the source, drain, and gate electrodes, the as-made biosensor foil was peeled off from the shadow mask for further electrical characterization. In many previous glucose sensing studies electrochemical sensors with large working electrodes were used with drop casting functionalization (W. Gao, et al, 2016, cited supra; and H. Lee, et al., 2017, cited supra). In several embodiments, an ink-jet printing technique was developed to functionalize the constructed FET In2O3glucose biosensors (SeeFIG.2C). Due to the small dimension (˜25 μm×500 μm) of the nanoribbon biosensors, utilization of the traditional drop-cast deposition method would lead the whole active sensing area to be covered by the chitosan film. In accordance with several embodiments, the channel area is to be kept exposed. Accordingly, several embodiments employ a SonoPlot printer with a 50 μm glass nozzle to print the chitosan ink only on the source and drain pads. The ink was made of chitosan, single-walled carbon nanotube and glucose oxidase, in accordance with various embodiments.

FIG.3Aprovides a photograph of an embodiment of an as-fabricated In2O3biosensor foil having a size of 5 cm×5 cm. An Optical image of a group of In2O3biosensors and two gold gate electrodes, in accordance with more embodiments, are provided inFIG.3B.FIG.3Cprovides a scanning electron microscope (SEM) image of an embodiment displaying the channel region and the gold gate of a biosensor. To further characterize embodiments of In2O3nanoribbons, atomic force microscopy (AFM) and X-ray diffraction (XRD) was used on samples deposited on PET substrate (FIGS.4A and4B). The AFM images show that the nanoribbons are solid and have clear edges. The height profile shows the thickness of In2O3nanoribbons is ˜20 nm. The XRD pattern shown presents only PET peaks, indicating the In2O3is amorphous.FIG.5Aprovides an embodiment of a fabricated In2O3nanoribbon FET foil conformably laminated onto an artificial human hand, indicating the conformability, bendability and wearability of the In2O3nanoribbon biosensors.FIG.5Bexhibits, in accordance with some embodiments, a biosensor foil rolled up with a radius of curvature of ˜1 mm. In more embodiments, the flexible biosensor can be further attached onto the back casing of a watch (FIG.5C), showing the concept that such In2O3transistor biosensors can be integrated with smart watches in the future. Several more embodiments are directed to flexible, lab-on-a-chip, and conformal In2O3nanoribbon electronics for wearable biosensor applications.

The embodiments of the invention will be better understood with the several examples provided within. Many exemplary biosensors are provided that are capable of measuring analytes, such as glucose, in bodily fluids, such as sweat, tears, and saliva. Also provided are various exemplary methods that may be utilized to practice the various embodiments. Exemplary experiments using the biosensors and methods and the resultant data are also described, further clarifying and enabling one to practice the numerous embodiments.

Electrical Characterization

Ag/AgCl electrodes are commonly used as reference electrodes in electrochemical measurements and biosensing applications due to their ability to provide stable potential and read voltage precisely. Integration of the Ag/AgCl electrode onto a biosensor chip, however, renders fabrication difficult and impractical. In accordance with a number of embodiments, gold gates are used in lieu of Ag/AgCl external electrodes to supply gate bias to the devices. In some embodiments, two gold gate electrodes are used in a group of four In2O3FETs. In more embodiments, gold gates are placed in the middle of the four In2O3FETs to supply gate voltage. In even more embodiments, gold gates are placed at the rear to monitor changes in potential on the devices. Performance of devices having gate voltage applied by the external Ag/AgCl electrode or the on-chip gold electrode was compared. The measurements were performed with the device active area immersed into a microwell filled with 300 μL electrolyte solution (0.1× Phosphate Buffered Saline (PBS)).FIGS.6A and6Bprovide family curves of drain current-gate voltage (IDS-VGS) and drain current-drain voltage (IDS-VDS) when the gate voltage was biased through a Ag/AgCl electrode. The schematic diagram of the measurement setup is illustrated in the inset ofFIG.6A. The performance of gold gate controlled In2O3FET is presented inFIGS.6C(IDS-VGS) and6D (IDS-VDS), generated in accordance with an embodiment of the invention. The output and transfer curves of the FET devices demonstrate that In2O3nanoribbon devices can work properly under gate bias supplied by the gold gate. The output characteristics of the FET devices demonstrated Ohmic behavior with a good linear regime in the “on” state, and the drain current got saturated when the bias increased further. All the curves inFIGS.6B and6Dpassing through the origin point indicate the minimal contribution of the gate leakage current to the drain current. The field-effect mobility of the In2O3FET, in accordance with various embodiments, is extracted to be 22.34±1.44 cm2V−1s−1using the following equation:

gm=dIDdVGS=WL⁢CDL⁢μFE⁢VDEq.⁢No.⁢1
where W is the channel width, L is the channel length, and CDLis the electrical double layer capacitance per unit area in 0.1 M ionic strength aqueous solution (25.52 μF cm−2) (See S. Park, et al, 2015, cited supra). Maximum transconductance 5.69 μS was observed at a drain voltage of 0.2 V and a gate voltage of 0.527 V (FIG.7). To further confirm the gate control of the on-chip gate electrode, and in accordance with several embodiments, one electrode was used as the gate bias supplier and another as a reference electrode to monitor the actual change of potential on the devices, as the scheme shows in the inset ofFIG.8A. InFIG.8A, the reference voltage (VREF) was plotted against the gold gate voltage (VGS) with different distances between those two electrodes, 150 μm, 750 μm, and 1350 μm, respectively. The plot provides that VREFis almost identical to VGSregardless of the distance. Drain current versus gate bias applied through the gold gate at difference distances was also plotted (FIG.8B), revealing negligible differences between gate-to-channel distances.

A statistical study of key electrical properties for 50 In2O3nanoribbon devices comparing gate biased through the Ag/AgCl electrode and the gold gate was conducted.FIG.9provides that the Ag/AgCl electrode and the gold gate devices performed nearly identically in assays assessing mobility (μ), threshold voltage (Vth), on/off ration, and on-state current. These data imply that the gold gate and the Ag/AgCl gate can have analogous gating effects. These data provide support that the on-chip gate electrode has a great control over the nanoribbon transistors in the aqueous environment, in accordance with a number of embodiments of the invention.

Flexibility

In order to characterize the flexibility of the wearable In2O3FETs, various bending tests were carried out. As shown inFIG.10A, fabricated In2O3foil was tightly wrapped around a cylinder. The electrical performance of the devices under tensile strain was measured.FIG.10Bcompares the transfer characteristics of a representative In2O3nanoribbon FET in three conditions: relaxed status, bent with a radius of curvature of ˜3 mm, and after 100 bending cycles. The devices, in accordance with numerous embodiments, exhibited n-type behavior in all three conditions without any perceptible change of their performances. In order to verify the reliability of various devices when deformed, flexibility tests were performed on In2O3FETs functionalized with a gel film containing chitosan, SWCNT, and glucose oxidase.

FIG.11provides plots of the mobility, the on-off ratio, and the threshold voltage averaged over 9 devices bent with a radius of curvature of infinity (relaxed), 3, 10, and 15 mm, respectively. Exemplary transfer curves of the devices under the different bending conditions are plotted inFIG.12.

Tensile strain of the various bent In2O3foils were calculated using the formula:

ɛ=1R×ds+df2×χ·γ2+2·χ·γ+1χ·γ2+χ·γ+γ+1Eq.⁢No.⁢2
where R is the bending radius, dsis the thickness of the substrate, and dfis the thickness of In2O3nanoribbon transistor (TFT). γ=df/dsand χ=Yf/Ys, where Yfand Ysare the Young's modulus of In2O3FET and the substrate, respectively. Accordingly, Yf=Ysis assumed and Eq. No. 2 can be further simplified:

ɛ=1R×ds+df2
The thickness of the substrate is 15 μm and the total thickness of the TFT is less than 100 nm. With the bending radius of 3 mm, the tensile strain is calculated to be ˜0.25%. The mobility as a function of tensile strain is plotted inFIG.13.

With a radius of curvature of ˜3 mm, a tensile strain of ˜0.25%, was applied to In2O3FETs parallel to the drain-to-source current direction. There was no significant change of the electrical performance of the In2O3FETs when the devices were in different bending conditions, as the mobility only showed small variation between 22.15±1.68 cm2V−1s−1and 22.70±1.65 cm2V−1s−1, the threshold voltage only showed variation between 0.273±0.028 V and 0.280±0.027 V, and the logarithm on-off ratio showed variation between 4.71±0.13 and 4.84±0.12.

FIG.14plots the mobility, the threshold voltage, and the on-off ratio of the devices without bending and after 5, 10, 50, and 100 bending cycles. As can be seen in the figure, the changes in device performance were negligible. The mobility varied in the range of 22.98±1.34 cm2V−1s−1and 23.78±1.87 cm2V−1s−1, the threshold voltage varied between 0.273±0.005 V and 0.266±0.016 V, and the logarithm on-off ratio varied between 4.98±0.17 to 4.96±0.14. On the basis of the test results, In2O3nanoribbon FETs after bending tests still maintained excellent performance, confirming that the platform is reliable under mechanical deformation. Exemplary transfer curves of the devices after 5, 10, 50, and 100 bending cycles are plotted inFIG.15.

pH and Glucose Sensing

Due to low volume of external body fluid, the ability to detect in a small amount of liquid is crucial to wearable sensors. A polydimethylsiloxane (PDMS) stamp was adapted to be used as a microwell to accumulate body fluids (FIG.16A). It can also serve as a passivation layer to ensure reliable sensing without electrical disturbance that can be introduced by contacting of metal lines with a body and/or body fluids. A mixture of curing agent and PDMS at a ratio of 1:10 was first spin-coated onto a silicon wafer before thermally cured at 80° C. for 1 h. After punching a hole with a diameter of 3 mm, the PDMS stamp was laminated onto the biosensor substrate utilizing van der Waals force. To guarantee the biosensor can work properly in a limited amount of liquid, the PDMS microwell was filled with 10 μL solution and electrical measurements were performed using a gold side gate electrode.FIGS.16B and16Cprovide transfer curves and output curves of the In2O3FETs measured with a gold gate in the electrolyte of ˜10 μL 0.1×PBS, generated in accordance with various embodiments. The electrical performance measured in a small amount of liquid is comparable to the results provide inFIGS.6C and6D(measured in 300 μL 0.1×PBS). Accordingly, several embodiments of the described biosensing platforms can efficiently work in the liquid with amount as small as 10 μL, which is a 30-fold decrease from that previous reports (Q. Liu, et al., 2016, cited supra).

To further establish the sensing ability of the described biosensor platforms, pH sensing experiments were conducted to test the ionic sensitivity of biosensor chips in response to commercial pH solutions.FIG.16Dprovides a comparison of the pH sensing responses (ΔI/I0) with gate bias supplied using either a gold electrodes, in accordance with embodiments, or Ag/AgCl electrodes. The responses are plotted into black upward pointing triangles and red downward pointing triangles for devices gated with a Ag/AgCl external liquid electrode and a gold electrode, respectively. The baseline current I0was obtained using 0.1×PBS (pH=7.4) to stabilize the device, and then the PBS was sequentially changed to commercial pH buffer solutions ranging from pH 10 to pH 5. Both gold and Ag/AgCl electrode devices increased in conduction when the pH value of the solution decreased, as hydroxyl groups on the nanoribbon surface were protonated due to more H+ions in the solution, resulting in the positive gating effect on the channel area of the n-type In2O3nanoribbon transistor. As observed, the sensing results from the gold and Ag/AgCl gate electrodes are almost identical. They both are exponentially dependent on pH changes, and the drain current increased ˜2.4 times when the pH value increased by 1. Representative real-time pH sensing results are provided inFIG.17.

In2O3nanoribbon biosensors were also tested for their ability to detect D-glucose.FIG.18Aprovides a schematic diagram depicting the working principle of the glucose determination using In2O3nanoribbon biosensors, in accordance with various embodiments. The surfaces of source and drain electrodes are functionalized with chitosan/carbon nanotube/glucose oxidase, which can be performed using ink-jet printing. In some embodiments, chitosan is the immobilization layer, which may be beneficial because it is a biocompatible polymeric matrix with good film-forming ability and high water permeability (H. Tang, et al, 2011, cited supra). Carbon nanotubes have been reported as efficient routes for increasing the sensitivity for many types of sensors, owing to their good electrocatalytic property and capacity for biomolecule immobilization (H. Tang, et al, 2011, cited supra; and S. Hrapovic, et al.,Anal. Chem.2004, 76, 1083-1088; and J. Wang, M. Musameh, and Y. LinJ. Am. Chem. Soc.2003, 125, 2408-2409, the disclosures of which are herein incorporated by reference). After immobilized onto the chitosan film and carbon nanotubes, glucose oxidase can accept electrons when interacting with glucose in the solution. The accepted electrons thereafter transfer to molecular oxygen to produce hydrogen peroxide (H2O2), which will be oxidized under a bias voltage. The reactions are as follows:

The generation of H+depends on the concentration of glucose. Decreasing of the pH leads to protonation the OH groups on the In2O3surface and results in changes in the local FET electric field, and ultimately causes changes in the conductance and current.FIG.18Bprovides a plot detailing continuous monitoring of sensing signals in response to increasing glucose concentrations. The channel current increases as the concentration of glucose increases, with a detection limit of about 10 nM (˜2.2% of the baseline current), in accordance with a number of embodiments. Accordingly, embodiments of glucose sensors can detect glucose in the concentration range between 10 nM to 1 mM, which covers typical glucose concentrations in human body fluids, such as, for example, sweat diabetes patients and healthy people (P. Makaram, D. Owens, and J. Aceros, 2014, cited supra). The detection limit observed with the described embodiments of In2O3nanoribbon biosensors is much lower than a typical electrochemical amperometric glucose sensor (W. Gao, et al., 2016, cited supra; and H. Lee, et al., 2017, cited supra). Fabricated sensors merely lacking glucose oxidase did not respond to glucose (FIG.19).

Wearable In2O3nanoribbon glucose sensors are further analyzed in external human body fluid, such as tears, sweat and saliva, which have much lower glucose concentrations than blood. While normal blood glucose levels range between 70 mg/dL (3.9 mM) and 140 mg/dL (7.8 mM) or higher, by contrast, tear glucose levels are on the order of 0.1-0.6 mM (E. R. Berman, BIOCHEMISTRY OF THEEYE, Springer Science & Business Media: 2013; H. Yao, et al.,Biosens. Bioelectron.2011, 26, 3290-3296; and H. Yao, et al.,J. Micromech. Microeng.2012, 22, 075007; the disclosures of which are herein incorporated by reference), sweat glucose has been reported to be 5 to 20 mg/dL (0.277 mM-1.11 mM) (J. Moyer, et al, 2012, cited supra), and saliva glucose concentrations are around 0.51-2.32 mg/dL (28.3 μM-0.129 mM) (P. Abikshyeet, V. Ramesh, and N. OzaDiabetes Metab. Syndr. Obes.2012, 5, 149; W. Zhang, Y. Du, and M. L. Wang,Sens. Biosens. Res.2015, 4, 23-29; and W. Zhang, Y. Du, and M. L. Wang,Sens. Biosens. Res.2015, 4, 96-102; the disclosures of which are herein incorporated by reference).FIG.20provides representative current responses to increasing glucose concentrations in artificial human tears, artificial human sweat, and saliva. Initially, wearable In2O3nanoribbon devices, in accordance with various embodiments, were submerged in 0.1×PBS to obtain the baseline current (FIG.20). When the electrolyte was changed from 0.1×PBS to artificial tears at 150 s, the sensing signal bumped up a little bit, due to the pH difference between the fluids. Signal Noise in artificial tears were higher than the results in PBS (compareFIG.18BandFIG.20). This increased amount of noise comes derives from the weaker buffer of the artificial tears, resulting in a decrease in the signal-to-noise ratio and affecting detection limit. The relationships between the glucose concentration and the saturated current response from the real-time sensing data in PBS solution, artificial tears, sweat, and saliva were extracted and plotted (FIG.20). The high correlation between the data with PBS and the data with artificial tears indicates that the detected signals from tears are attributed to mainly glucose and not other nonspecific proteins. In the cases of artificial sweat and saliva, even though the sensing signals are slightly lower than the responses from PBS, which may be due to their different ionic strengths and complex ingredients, the sensors can differentiate the glucose concentration as low as 0.1 μM. This sensitivity is sufficient to detect glucose in both sweat and saliva, in accordance with many embodiments.

Wearable In2O3biosensors, in accordance with multiple embodiments, can be comfortably attached onto an artificial eyeball and an artificial arm (FIG.21). To ensure the on-body sensing ability, the data collection on an artificial eyeball with the biosensor facing out was imitated.FIG.22provides ex-situ glucose sensing results using artificial tears. Indium wires were used to connect the bonding pads to the measurement unit, and artificial tears were constantly flowing through the sensing area (see inset ofFIG.22). After obtaining a stable baseline current, artificial tears were sequentially flowing, spiked with 0.01, 0.1, 1, 10, 100, and 1000 μM glucose. The sensing results demonstrate that the wearable glucose sensing platform, in accordance to several embodiments, can be utilized in conjunction with contact lenses when embedded with various sensors described herein.

Glucose sensing, according to a number of embodiments, was also performed on an artificial arm, but with the sensor facing the skin. Sensing results are provided inFIG.22, which demonstrates that In2O3biosensors, in accordance with numerous embodiments, can work as sweat patch for glucose monitoring. To further confirm that the sensing platform can be utilized as wearable sweat analyzer, sweat samples was collect from human subjects' foreheads during exercise. The sweat was spiked with different concentrations of glucose and sensing was performed as described herein and in accordance with many embodiments.FIG.23provides sensing results with real sweat. The sensing signal shows a large increase after the PBS was replaced with sweat due to the changes in pH and intrinsic glucose concentration. Good sensitivity was observed ranging from 0.1 μM to 1 mM, indicating that the sensing platform described herein can be used for wearable sweat analysis.

Sweat glucose levels were also measured before and after meal of an individual with no observable health deficiencies. Sweat samples were collected 30 min before and 30 min after intake of a glucose-rich beverage. The sensing results are provided inFIG.24. The inset figure provides the device transfer curve measured of sweat samples acquired before and after glucose intake. For comparison, the subject's blood sugar level before and after glucose intake was also recorded using a commercial glucose meter, which recorded glucose concentrations 79 mg/dL and 118 mg/dL, respectively.

To determine storability of In2O3biosensor functionalized with chitosan/CNT/GOx, glucose measurements were performed using a single sensor at intermittent time points over two weeks. The device was used to measure glucose every day and kept stored at 4° C. in between measurements. Results of glucose sensing are provided inFIG.25. Over the first 4 days, there was little no loss of detected signal. Furthermore, after two weeks, detection of glucose concentrations between 10 μM and 100 μM glucose in PBS decreased only about 25% and 30% (FIG.25). The decrease in the glucose detection ability can be attributed to the deactivation of the glucose oxidase and/or the loss of enzyme during washing steps. Despite this loss of detection due to repeated uses, it should be noted, that enzyme degradation would be mitigated in a single (or couple) usage regime. Accordingly, embodiments are directed biosensors utilizing low-cost and/or disposable devices, such as devices with In2O3FET with gold gate electrodes, as described herein.

A PET substrate was first cleaned with acetone and isopropyl alcohol, and then went through ultra violet treatment before the fabrication process. After the cleaning process, the first shadow mask was attached to the PET substrate to define the channel area. Then the In2O3nanoribbons were deposited by RF sputtering (Denton Discovery 550 sputtering system). By simply detaching the shadow mask, well patterned nanoribbons were formed. The source, drain, and gold electrodes were then defined by the second shadow mask, and followed with electron beam evaporation of 1 nm Ti and 50 nm Au. After deposition, the shadow mask was removed.

Optical microscopy images were taken with an Olympus microscope. The SEM images were taken with a Hitachi S-4800 field emission scanning electron microscope. Electrical characteristics and sensing results were measured with an Agilent 1500B semiconductor analyzer.

1 weight % (wt %) chitosan powder was first dissolved in 2 wt % acetic acid aqueous solution. Next, the chitosan solution was mixed with single-walled carbon nanotubes (SWCNT) (2 mg ml−1in 1×PBS) using ultrasonication for over 30 min. The chitosan/SWCNT solution was mixed with glucose oxidase solution (10 mg ml−1in 1×PBS) in the volume ratio 2:1. The mixed solution was then ink-jet printed onto the source and drain electrode, and dried under ambient conditions.

Human Body Fluid Samples.

Artificial human tear was bought from Walgreens. Artificial human sweat was prepared by mixing 22 mM urea, 5.5 mM lactic acid, 3 mM NH4+, 100 mM Na+, 10 mM K+, 0.4 mM Ca2+, 50 μM Mg2+and 25 μM uric acid with varying glucose concentrations. Real sweat samples were collected from human by scratching their foreheads with micro tubes.

While exemplary embodiments are described above, it is not intended that these embodiments describe all possible forms of the invention. Rather, the words used in the specification are words of description rather than limitation, and it is understood that various changes may be made without departing from the spirit and scope of the invention. Additionally, the features of various implementing embodiments may be combined to form further embodiments of the invention. Accordingly, the scope of the invention should be determined not by the embodiments illustrated, but by the appended claims and their equivalents.