Nonenzymatic Wearable Sensor for Electrochemical Analysis of

(6−8) The last several years have witnessed a trend to integrate sensors for ... Figure 1. (a) Perspiration glucose analysis during physical activit...
0 downloads 0 Views 5MB Size
Download PDF
Article pubs.acs.org/acssensors

Cite This: ACS Sens. XXXX, XXX, XXX−XXX

Nonenzymatic Wearable Sensor for Electrochemical Analysis of Perspiration Glucose Xiaofei Zhu, Yinhui Ju, Jian Chen, Deye Liu, and Hong Liu* State Key Laboratory of Bioelectronics, School of Biological Science and Medical Engineering, Southeast University, Nanjing 210096, China S Supporting Information *

ABSTRACT: We report a nonenzymatic wearable sensor for electrochemical analysis of perspiration glucose. Multipotential steps are applied on a Au electrode, including a high negative pretreatment potential step for proton reduction which produces a localized alkaline condition, a moderate potential step for electrocatalytic oxidation of glucose under the alkaline condition, and a positive potential step to clean and reactivate the electrode surface for the next detection. Fluorocarbon-based materials were coated on the Au electrode for improving the selectivity and robustness of the sensor. A fully integrated wristband is developed for continuous real-time monitoring of perspiration glucose during physical activities, and uploading the test result to a smartphone app via Bluetooth. KEYWORDS: wearable sensor, electrochemical analysis, perspiration glucose, nonenzymatic, real-time monitoring

W

biomolecule, the activity of an enzyme can be affected by a range of factors such as temperature, pH, and ionic strength which can hardly be controlled under in situ conditions.31 Second, enzymes can be degraded. Even for glucose oxidase that is highly glycated and thus stable, its activity can gradually decrease with time which affects its shelf life and the ability for long-term wearable monitoring.32−36 Third, the immobilization of enzyme on electrodes involves processes such as covalent cross-linking, polymerization, or sol−gel entrapment on the electrode surface that not only suppresses the activity of the enzyme, but also the immobilized reagents on the electrode may slow down the electron transfer for detection.37 Finally, commercially available enzyme is usually a biosourced agent for in vitro analysis only; it may not be safe for wearable sensing.32,38,39 To solve these problems, it is highly desirable to develop a wearable sensor for nonenzymatic electrochemical detection of glucose. Actually, nonenzymatic electrochemical glucose detection has been investigated for decades. For these sensors, electrocatalytic materials of glucose including bulk metal,40−42 metal oxide,43,44 alloy,45,46 metal nanomaterials,47,48 and carbon nanocomposites were used.49−52 Despite decades of research, one of the most important issues for using them in wearable sensing is that the electrocatalytic oxidation of glucose usually requires an alkaline condition (pH > 11) for acceptable selectivity, sensitivity, and reproducibility.53−58 At physiological pH of body fluid (e.g., perspiration), a dramatic decrease or

earable electronic devices equipped with physical sensors that can monitor physical activity and vital signs have been commercially available for years.1−5 However, these devices only provide limited information regarding human health, and their sensitivity to human physiological states is usually too low for practical application.6−8 The last several years have witnessed a trend to integrate sensors for chemical analysis onto wearable devices that can at the molecular level provide insight into the state of human health for applications such as early diagnostics.9−13 For all kinds of wearable chemical sensors, perspiration has been widely used as a sampling medium so that valuable information such as electrolyte balance, diet, injury, stress, medications, and hydration can be continuously monitored.14,15 For perspiration chemical sensing, an electrochemical detection method is almost exclusively involved because it is simple, sensitive, and intrinsically quantitative and consumes relatively low electrical power.10,16,17 Sensors for perspiration glucose monitoring has attracted much attention and has been reported by several researchers owing to the growing worldwide problem of diabetes.10,18−22 Owing to the relationship between perspiration glucose and blood glucose, the wearable perspiration sensor is promising as an alternative to the invasive, painful, and inconvenient blood glucose test for early screening and even self-monitoring of diabetes.18,19,23,24 Almost all current wearable glucose sensors used an enzyme (e.g., glucose oxidase or dehydrogenase) to convert glucose to a detectable electrochemical signal.25−27 The enzyme is a dispensable reagent for a disposable test strip in a personal blood glucose test.28−30 For wearable sensing however, the use of enzyme leads to several critical problems. First, as a © XXXX American Chemical Society

Received: February 26, 2018 Accepted: May 16, 2018 Published: May 16, 2018 A

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors even complete loss of the electrochemical response was observed.59 Besides, detection of glucose at nonalkaline physiological conditions suffer from serious electrode poisoning or competitive adsorption of chloride, sulfides, and phosphates.54,60−62 Here we report a nonenzymatic wearable sensor for electrochemical perspiration glucose analysis. Multipotential steps are applied on a Au electrode, including a high negative potential step for pretreatment of the sample which produces a localized alkaline condition on the electrode surface, a moderate potential for glucose detection under the alkaline condition, and a positive potential to clean the electrode surface. Fluorocarbon-based materials are coated on the Au electrode for improving the selectivity and robustness of the sensor. A fully integrated wristband is developed for real-time monitoring of perspiration glucose, and uploading the test results to a smartphone app via Bluetooth, as shown in Figure 1. The analytical results are also compared with that obtained using a HPLC-MS.

Figure 2. (a) Potential measured from a Au electrode as a function of time for the pH determination on the electrode surface. A constant current of −10 μA was applied. The pH of the solution is indicated. Inset: Potential measured at 200 s as a function of the solution pH. (b) Localized pH measured on the Au electrode surface when a constant potential of −2.0 V was applied for pretreatment as a function of the time measured in solutions with different pHs. The error bar represents standard deviation for three replicated measurements.

pH. This is reasonable because the negative current step induces the reduction of proton on the surface of the electrode; low concentration of protons at high pH resulted in large overpotential for proton reduction.63,64 With the calibration curve shown in Figure 2a, the dynamic pH on the electrode surface during the pretreatment step can be measured using the chronopotentiometric technique. For the pretreatment, a negative potential of −2.0 V, which created an alkaline condition for glucose oxidation, was applied on the working electrode. At −2.0 V, protons were reduced to hydrogen leading to the production of hydroxide ions and a localized increase of pH near the Au surface. As shown in Figure 2b, the pH increased with increasing time and finally approached a maximum value probably owing to equilibrium between the generation of hydroxide ions and diffusion of these ions toward the bulk solution. The pH range of human perspiration is from 4.0 to 7.0.68 We prepared solutions buffered within this pH range, and measured pH change with time during the proton reduction reaction using the chronopotentiometric technique. As shown in Figure 2b, the time required for reaching the equilibrium increased with deceasing pH. After about 20 s, the pH for all of the solutions was about 11. Since the pH has a crucial influence on glucose oxidation on Au, a pretreatment time of 20 s was chosen to ensure the reproducibility for further experiments. Furthermore, the pH change was localized within the diffusion layer

Figure 1. (a) Perspiration glucose analysis during physical activity using the wristband-based electrochemical sensor. (b) Photograph of the electrochemical sensor on the wristband for perspiration glucose analysis, including a Au working electrode (W), a Pt black-coated Pt counter electrode (C), and a polypyrrole-coated Pt as the quasireference electrode (R). (c) Photograph of the smartphone with app for the perspiration analysis. The sensor was connected to the smartphone via Bluetooth.

■

RESULTS AND DISCUSSION Au was used as the working electrode owing to its high electrocatalytic activity toward glucose oxidation under alkaline condition. So the pH near the surface of the electrode would have a dramatic influence on the signal for wearable sensing. In this work, an electrochemical pretreatment was introduced to increase the pH of the sample by proton reduction reaction. To determine the pH of the solution near the surface of the electrode after the pretreatment, a chronopotentiometric measurement was carried out by applying a constant current of −10 μA to trigger the proton reduction and simultaneously measure the potential on the electrode. Solutions buffered at different pHs were prepared, and the potential measured as a function of the pH was shown in Figure 2a. The potential measured was linearly correlated to the pH of the solution. More negative potential was obtained for solutions with higher B

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors near the electrode surface which was much smaller in volume than the bulk solution. So, it should not cause irritation to the skin during the wearable sensing process. The cyclic voltammogram of the Au electrode in 0.10 M PBS (pH 5.0) after pretreatment at −2.0 V for 20 s was shown in Figure 3a. The current obtained from the Au electrode (curve c) increased significantly compared with that without the pretreatment (curve a). An anodic wave at 0.50 V, which was attributed to formation of Au oxide, and a cathodic wave at 0.10

V in the reverse scan, which was attributed to reduction of the oxide, were observed. With 8.0 mM glucose in the solution, two distinctive anodic waves were observed (curve d). As previously reported,65−67 the anodic wave at −0.4 V was caused by oxidation of the carbonyl group of glucose, and the anodic wave at 0.2 V was caused by combined oxidation of carbonyl and hydroxyl groups of glucose. The peak of the reverse scan represented further oxidation of the carbonyl and hydroxyl groups, which occurred at a more negative potential, when the passivating oxide layer was removed.54 By holding the potential of the Au electrode at 0.20 V, glucose can be electrochemically oxidized and thus quantitatively determined by measuring the current. However, the products of the glucose oxidation reaction can passivate the electrode.69,70 For continuous wearable sensing, the electrode has to be reactivated after each test. So, after glucose oxidation at 0.20 V, a potential of 1.0 V is applied to clean the electrode surface.71−74 For wearable sensing of perspiration glucose, multipotential steps were now applied to the Au working electrode (Figure S1 in the Supporting Information). The multipotential steps included a potential of −2.0 V for 20 s in order to provide the alkaline condition, a potential of 0.2 V for 5 s for glucose determination, and finally a potential of 1.0 V for 2 s to clean the electrode surface. The corresponding i−t curve of the multipotential steps for detection of glucose with the concentration ranging from 20 to 300 μM was shown in Figure 3b. The current of glucose oxidation was linearly correlated to the glucose concentration (Figure 3c). For analysis of real human perspiration, various compounds in the perspiration can cause interference to the detection. To reduce the interference from the sample matrix, the electrode was covered with a layer of Nafion and then a layer of Kel-F membrane. The Nafion is a cation-exchange polymer membrane, which can selectively exclude anions from the electrode surface.75 The Kel-F membrane is a type of fluorocarbon materials, which can further repel charged molecules (e.g., amino acids and ascorbic acid).76 For wearable sensing, the Kel-F has unusual and outstanding properties. It is chemically stable, which can resist fuming nitric acid, aqua regia, and continued exposure to air at 200 °C. It is insoluble in most solvents. It has high softening points and extremely low brittle points.77 The chemical stability and durability of the Kel-F layer ensured the robustness of the sensor. The effect of interferents in real perspiration such as ascorbic acid, uric acid, lactic acid, and glutamic acid on the detection were investigated. The interferents in large excess (compared with that in human perspiration) were added into the sample. The results of testing the sample showed that the interference was negligible (Figure S2 in the Supporting Information). Glucose was analyzed in a real human perspiration sample using the Nafion and Kel-F modified Au electrode. The sample was obtained from a volunteer in the laboratory and different amounts of glucose were added ranging from 30 to 1100 μM, which covers usual physiological levels of perspiration glucose. As shown in Figure 4a, the current increased with increasing concentration of glucose. The current measured was linearly correlated to the concentration of glucose from 30 to 1100 μM. The sensitivity of the sensor was 114 μA mM−1 cm−2.The limitof-detection, which was calculated as 3 times the standard deviation of the testing results of the blank divided by the slope of the calibration curve, was 15 μM.

Figure 3. (a) Cyclic voltammogram obtained in a 0.10 M PBS (pH 5.0) using a Au electrode without (a and b) and with (c and d) the pretreatment. For the pretreatment, a potential of −2.0 V was applied on the Au electrode for 20 s. For b and d, the solution also contained 8.0 mM glucose. (b) Current measured as a function of time from the Au electrode with multiple potential steps applied: −2.0 V for 20 s to provide the alkaline condition, 0.2 V for 5 s to oxidize the glucose, and finally 1.0 V for 2 s to clean the electrode surface. The initial concentration of glucose in the sample was 0 mM, additional glucose was added after each test so that the resulting glucose concentration in the sample was 20, 50, 100, 150, 200, and 300 μM, respectively. (c) Current due to glucose oxidation as a function of time extracted from (b). Inset: the current measured at 5 s after initiation of the glucose oxidation versus glucose concentration. The error bar represents standard deviation for three replicated measurements. C

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors

Figure 4. (a) Current corresponding to glucose oxidation as a function of time for detection of a real perspiration sample with added glucose. Inset: the current measured at 5 s versus the glucose concentration. (b) Analytical results obtained using the wearable electrochemical sensor versus that obtained using a HPLC-MS. The error bar represents standard deviation for three replicated measurements.

Figure 5. (a) Results of replicated analysis of one perspiration sample using the wristband sensor. (b) Results of replicated analysis of perspiration glucose over 2 weeks using the wearable sensor. The error bars represent standard deviation for five parallel measurements.

■

CONCLUSIONS We have reported the nonenzymatic wearable sensor for electrochemical analysis of perspiration glucose. Multipotential steps were involved for electrocatalytic oxidation of glucose under alkaline condition using a Au electrode and reactivation of the electrode for the next analysis. The selectivity and robustness of the sensor were improved by coating two kinds of fluorocarbon-based materials on the Au electrode. A fully integrated wristband was developed for continuous real-time monitoring of perspiration glucose during physical activity, and uploading the test result to a smartphone app via Bluetooth. The analytical results of the sensor were also compared with that obtained using a HPLC-MS. Therefore, we believe the sensor is promising for early screening of diabetes and, in the future, can be complementary to the invasive, painful, and inconvenient blood glucose test for self-monitoring of diabetes.

The sensing system, including the three electrodes and a miniaturized potentiostat with Bluetooth connection to an Android-based smartphone app was integrated in a wristband fabricated using a three-dimensional printer for wearable sensing, as shown in Figure 1.The three electrodes included the Au electrode as the working electrode, a Pt black coated Pt electrode as the counter electrode, and a Pt electrode with electropolymerized polypyrrole as the quasi-reference electrode. The three electrodes were included in an electrochemical cell so that they were kept away from direct contact with skin to avoid irritation. Three volunteers wore the wristband and cycled for 30 min. After perspiration, a test button was pressed. After a test time of 31 s, the test result was displayed on the screen of the smartphone. The glucose in the perspiration sample was also analyzed using HPLC-MS. The results were in agreement with that obtained using our sensor, as shown in Figure 4b. The reproducibility of the wearable sensor was evaluated by replicated determination of one perspiration sample for 7 times. The relative standard deviation (RSD) was about 7.5% (Figure 5a). To evaluate the long-term stability of the sensor, it was used to detect perspiration sample over a period of 2 weeks (Figure 5b). During the long-term wearable sensing process, no irritation to skin was observed by any of the volunteers. Results indicated that during the 2 week period of time, the RSD for perspiration analysis was about 4.6%, which demonstrated the reproducibility and practical applicability of the sensor.

■

EXPERIMENTAL SECTION

Chemicals and Apparatus. Disodium hydrogen phosphate dodecahydrate (Na2HPO4·3H2O), potassium dihydrogen phosphate (KH2PO4), hexachloroplatinic acid (H2PtCl4·6H2O), sodium hydroxide (NaOH), anhydrous citric acid, sodium chloride (NaCl), potassium chloride (KCl), acetonitrile, tetrabutylammonium hexafluorophosphate (98%), Nafion perfluorinated solution (20 wt % in mixture of lower aliphatic alcohols and water, contained 34% water), and Kel-F oil were obtained from Sigma-Aldrich. Pyrrole (99%) was purchased from Macklin. Anhydrous glucose was obtained from Sinopharm Chemical Reagent (China). All reagents were of analytical grade unless specifically indicated. Au rod (99.9%, 2 mm diameter), Pt foil (99.9%), and Pt rod (99.9%, 2 mm diameter) were obtained from Alfa Aesar. A three-dimensional printer (M-Jewelry) was used to print the wristband. CHI 8520d workstation from Chenhua (Shanghai, D

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors China) was used for electrochemical measurements. Ultrapure water was used in all of our experiments. All reagents were used as received without further purification. Electrochemical pH Measurement. The electrochemical measurements were carried out using a CHI 8520d workstation at 25 °C. Citrate buffered solution with different pHs, including 2.2, 3.0, 4.0, 5.0, 6.0, 7.0, and 8.0, were prepared. Three-electrode system was employed to carry out electrochemical experiments, including a Au disk electrode as the working electrode, a Pt wire as the counter electrode, and a Ag/ AgCl electrode as the reference electrode. The pH measurement was conducted using chronopotentiometric technique with a sampling interval of 0.1 s. A current of −10 μA was applied for 200 s. For pretreatment that changed the pH of the solution, a negative potential of −2.0 V was applied for 20 s before the chronopotentiometric pH measurement as mentioned above. Glucose Detection. Glucose detection in 0.10 M PBS (pH 5.0) was conducted by applying multipotential steps to the working electrode. Three potential steps were involved, including a high negative potential of −2.0 V for 20 s, a moderate potential of 0.2 V for 5 s, and a positive potential of 1.0 V for 2 s. Preparation of Pseudo-Reference Electrode. For wearable sensing, a polypyrrole pseudoreference electrode was prepared. A Pt rod (2 mm diameter) was polished with abrasive paper (3000 mesh) and chamois leather, respectively, and used as the working electrode. Electrochemical polymerization of pyrrole was carried out in a threeelectrode cell with a Ag/AgCl reference electrode and a carbon counter electrode. An acetonitrile solution containing 0.10 M tetrabutylammonium hexafluorophosphate and 0.01 M pyrrole was used as the electrolytic solution. The electrodeposition of polypyrrole was carried out by sweeping the potential at a scan rate of 0.10 V/s from −0.8 to 1.2 V for 50 cycles, as previously reported (Figure S3 in the Supporting Information).78 After a few cycles, the black polypyrrole film was observable on the Pt surface. For the final scan, the scan was stopped at a potential of 0.4 V versus Ag/AgCl so that the polypyrrole was partially oxidized. After the electropolymerization, the electrode was rinsed with ultrapure water. Fabrication of Wristband Sensor. A gold rod (2 mm diameter) was polished with abrasive paper (3000 mesh) and chamois leather, respectively, followed by ultrasonic cleaning with ultrapure water, ethanol, and ultrapure water each for 5 min, respectively. The rod was finally dried in air and used as the working electrode (Figure S4 (a)). The electrode surface was modified with Nafion by dipping the electrode into 5% Nafion in low aliphatic alcohols and water and drying. Then, a 5% Nafion and 8% Kel-F oil mixture was mixed together in a 1:2 rato, and was dropcast on the electrode. The electrode was dried in an oven for 5 h at 60 °C.76 A Pt electrode coated with Pt black was used as counter electrode. Briefly, a Pt rod (2 mm diameter) was polished with abrasive paper (3000 mesh) and chamois leather, respectively. Pt black was electrochemically deposited on the surface with a constant potential of −0.25 V in a 5.0 mL solution containing 1.0 mM H2PtCl4 and 0.50 M H2SO4.79 For wearable sensing, a three-electrode system, including the polypyrrole-coated Pt as the quasi-reference electrode, the gold rod modified with selective membrane as the working electrode, and the Pt black-coated Pt as the counter electrode, was assembled into a threedimensional printed wristband shell (Figure S4 (b)). A printed circuit board acted as interlayer in order to connect the sensor with miniaturized potentiostat (Figure S4 (c, d, and f)). A Bluetooth low energy chipset (NRF51822) was used for wireless connection to a smartphone. A Rainsun 2.45 GHz chip antena (AN2051) and impedance matched Johanson Technology balun (2450BM14E0003) were employed for wireless transmission. A lithium battery (3.7 V, 1000 mAh) regulated via an ADP151_3v3 low dropout voltage regulator were utilized as the power source. The printed circuit board for the system was shown in Figure S4 (e). The averaged total energy consumption of the system was 30 mW during the analysis. Glucose Detection in Persipiration. Perspiration samples were collected from volunteers in the laboratory. The concentration of glucose in the perspiration samples were measured by HPLC-MS. The perspiration sample containing glucose was introduced into the

electrochemical cell of the wearable sensor for glucose determination. For real-time glucose monitoring, the wristband was connected to the smartphone app via Bluetooth. Volunteers were asked to wear the wristband and cycled for half an hour. The test button was then pressed down to initiate the analysis and results were displayed on the screen of smartphone after completion of the analysis.

■

ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acssensors.8b00168. Information about the fabrication and experiments (PDF)

■

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. ORCID

Hong Liu: 0000-0002-9841-1603 Notes

The authors declare no competing financial interest.

■

ACKNOWLEDGMENTS We gratefully acknowledge financial support from Chinese Recruitment Program of Global Experts, Innovative and Entrepreneurial Talent Recruitment Program of Jiangsu Province, the National Natural Science Foundation of China (21635001), State Key Project of Research and Development (2016YFF0100802), the Science and Technology Development Program of Suzhou (ZXY201439), the Project of Special Funds of Jiangsu Province for the Transformation of Scientific and Technological Achievements (BA2015067), the Fundamental Research Funds for the Central Universities (2242017K41015).

■

REFERENCES

(1) Kim, D. H.; Lu, N.; Ma, R.; Kim, Y. S.; Kim, R. H.; Wang, S.; Wu, J.; Won, S. M.; Tao, H.; Islam, A.; et al. Epidermal Electronics. Science 2011, 333, 838−843. (2) Yang, T.; Jiang, X.; Zhong, Y.; Zhao, X.; Lin, S.; Li, J.; Li, X.; Xu, J.; Li, Z.; Zhu, H. A Wearable and Highly Sensitive Graphene Strain Sensor for Precise Home-Based Pulse Wave Monitoring. ACS Sens. 2017, 2, 967−974. (3) Güder, F.; Ainla, A.; Redston, J.; Mosadegh, B.; Glavan, A.; Martin, T. J.; Whitesides, G. M. Paper-Based Electrical Respiration Sensor. Angew. Chem., Int. Ed. 2016, 55, 5727−5732. (4) Pang, C.; Koo, J. H.; Nguyen, A.; Caves, J. M.; Kim, M.-G.; Chortos, A.; Kim, K.; Wang, P. J.; Tok, J. B.-H.; Bao, Z. Highly SkinConformal Microhairy Sensor for Pulse Signal Amplification. Adv. Mater. 2015, 27, 634−640. (5) Gong, S.; Schwalb, W.; Wang, Y.; Chen, Y.; Tang, Y.; Si, J.; Shirinzadeh, B.; Cheng, W. A Wearable and Highly Sensitive Pressure Sensor with Ultrathin Gold Nanowires. Nat. Commun. 2014, 5, 3132− 3139. (6) Ho, M. D.; Ling, Y.; Yap, L. W.; Wang, Y.; Dong, D.; Zhao, Y.; Cheng, W. Percolating Network of Ultrathin Gold Nanowires and Silver Nanowires toward “Invisible” Wearable Sensors for Detecting Emotional Expression and Apexcardiogram. Adv. Funct. Mater. 2017, 27, 1700845. (7) Wang, Y.; Gong, S.; Wang, S. J.; Simon, G. P.; Cheng, W. L. Volume-Invariant Ionic Liquid Microbands as Highly Durable Wearable Biomedical Sensors. Mater. Horiz. 2016, 3, 208−213. (8) Takei, K.; Honda, W.; Harada, S.; Arie, T.; Akita, S. Toward Flexible and Wearable Human-Interactive Health-Monitoring Devices. Adv. Healthcare Mater. 2015, 4, 487−500.

E

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors

(29) Heller, A.; Feldman, B. Electrochemical glucose sensors and their applications in diabetes management. Chem. Rev. 2008, 108, 2482−2505. (30) Green, M. J.; Hilditch, P. I. Disposable single-use sensors. Analyst 1991, 116, 1217. (31) Wilson, R.; Turner, A. P. F. Glucose oxidase: an ideal enzyme. Biosens. Bioelectron. 1992, 7, 165−185. (32) Li, J.; Lin, X. Glucose biosensor based on immobilization of glucose oxidase in poly (o-aminophenol) film on polypyrrole-Pt nanocomposite modified glassy carbon electrode. Biosens. Bioelectron. 2007, 22, 2898−2905. (33) Wu, B.; Zhang, G.; Shuang, S.; Choi, M. M. F. Biosensors for determination of glucose with glucose oxidase immobilized on an eggshell membrane. Talanta 2004, 64, 546−553. (34) Han, K.; Wu, Z.; Lee, J.; Ahn, I.-A.; Park, J. W.; Min, B. R. Activity of glucose oxidase entrapped in mesoporous gels. Biochem. Eng. J. 2005, 22, 161−166. (35) Heller, A.; Feldman, B. J. Electrochemistry in diabetes management. Acc. Chem. Res. 2010, 43, 963−973. (36) Heller, A.; Feldman, B. J.; Say, J.; Vreeke, M. S. Small volume in vitro analyte sensor. US Patent No 6,143,164. (37) Wang, G. F.; He, X. P.; Wang, L. L.; Gu, A. X.; Huang, Y.; Fang, B.; Geng, B. Y.; Zhang, X. J. Microchim. Acta 2013, 180, 161−186. (38) Wilson, R.; Turner, A. P. F. Glucose oxidase: an ideal enzyme. Biosens. Bioelectron. 1992, 7, 165−185. (39) Toghill, K. E.; Compton, R. G. Electrochemical non-enzymatic glucose sensors: a perspective and an evaluation. Int. J. Electrochem. Sci. 2010, 5, 1246−1301. (40) Ernst, S.; Heitbaum, J.; Hamann, C. H. The electrooxidation of glucose in phosphate buffer solutions: Part I. Reactivity and kinetics below 350 mV/RHE. J. Electroanal. Chem. Interfacial Electrochem. 1979, 100, 173−183. (41) Luo, P.; Prabhu, S. V.; Baldwin, R. P. Comparison of metallic electrodes for constant-potential amperometric detection of carbohydrates, amino acids and related compounds in flow systems. Anal. Chim. Acta 1991, 244, 169−178. (42) Nagy, L.; Nagy, G.; Hajós, P. Copper electrode based amperometric detector cell for sugar and organic acid measurements. Sens. Actuators, B 2001, 76, 494−499. (43) Reitz, E.; Jia, W.; Gentile, M.; Wang, Y.; Lei, Y. CuO nanospheres based nonenzymatic glucose sensor. Electroanalysis 2008, 20, 2482−2486. (44) Cheng, X.; Zhang, S.; Zhang, H. Y.; Wang, Q. J.; He, P. G.; Fang, Y. Z. Determination of carbohydrates by capillary zone electrophoresis with amperometric detection at a nano-nickel oxide modified carbon paste electrode. Food Chem. 2008, 106, 830−835. (45) Marioli, J. M.; Kuwana, T. Electrochemical detection of carbohydrates at nickel-copper and nickel-chromium-iron alloy electrodes. Electroanalysis 1993, 5, 11−15. (46) Marioli, J.; Luo, P. F.; Kuwana, T. Nickelchromium alloy electrode as a carbohydrate detector for liquid chromatography. Anal. Chim. Acta 1993, 282, 571−580. (47) Kurniawan, F.; Tsakova, V.; Mirsky, V. M. Gold nanoparticles in nonenzymatic electrochemical detection of sugars. Electroanalysis 2006, 18, 1937−1942. (48) Feng, D.; Wang, F.; Chen, Z. Electrochemical glucose sensor based on one-step construction of gold nanoparticle−chitosan composite film. Sens. Actuators, B 2009, 138, 539−544. (49) Ye, J. S.; Wen, Y.; Zhang, W. D.; Gan, L. M.; Xu, G. Q.; Sheu, F. S. Nonenzymatic glucose detection using multi-walled carbon nanotube electrodes. Electrochem. Commun. 2004, 6, 66−70. (50) Chen, J.; Zhang, W. D.; Ye, J. S. Nonenzymatic electrochemical glucose sensor based on MnO2/MWNTs nanocomposite. Electrochem. Commun. 2008, 10, 1268−1271. (51) Luo, J.; Jiang, S.; Zhang, H.; Jiang, J.; Liu, X. A novel nonenzymatic glucose sensor based on Cu nanoparticle modified graphene sheets electrode. Anal. Chim. Acta 2012, 709, 47−53. (52) Lu, L. M.; Li, H. B.; Qu, F.; Zhang, X. B.; Shen, G. L.; Yu, R. Q. In situ synthesis of palladium nanoparticle−graphene nanohybrids and

(9) Chen, Y.; Lu, S.; Zhang, S.; Li, Y.; Qu, Z.; Chen, Y.; Lu, B.; Wang, X.; Feng, X. Skin-like biosensor system via electrochemical channels for noninvasive blood glucose monitoring. Sci. Adv. 2017, 3, e1701629. (10) Gao, W.; Emaminejad, S.; Nyein, H. Y. Y.; Challa, S.; Chen, K.; Peck, A.; Fahad, H. M.; Ota, H.; Shiraki, H.; Kiriya, D.; Lien, D.-H.; Brooks, G. A.; Davis, R. W.; Javey, A. Fully integrated wearable sensor arrays for multiplexed in situ perspiration analysis. Nature 2016, 529, 509−514. (11) Kim, J.; Imani, S.; de Araujo, W. R.; Warchall, J.; ValdesRamírez, G.; Paixao, T. R. L. C.; Mercier, P.; Wang, J. Wearable salivary uric acid mouthguard biosensor with integrated wireless electronics. Biosens. Bioelectron. 2015, 74, 1061−1068. (12) Bandodkar, A. J.; Jia, W.; Yardımcı, C.; Wang, X.; Ramirez, J.; Wang, J. Tattoo-based noninvasive glucose monitoring: a proof-ofconcept study. Anal. Chem. 2015, 87, 394−398. (13) Yao, H.; Liao, Y.; Lingley, A. R.; Afanasiev, A.; Lahdesmaki, I.; Otis, B. P.; Parviz, B. A. A contact lens with integrated telecommunication circuit and sensors for wireless and continuous tear glucose monitoring. J. Micromech. Microeng. 2012, 22, 075007. (14) Coyle, S.; Curto, V. F.; Benito-Lopez, F.; Florea, L.; Diamond, D. Wearable bio and chemical sensors. Wearable sensors 2014, 65−83. (15) Heikenfeld, J. Let them see you sweat. IEEE Spectrum 2014, 51, 46−63. (16) Munje, R. D.; Muthukumar, S.; Panneer Selvam, A.; Prasad, S. Flexible nanoporous tunable electrical double layer biosensors for sweat diagnostics. Sci. Rep. 2015, 5, 14586. (17) Kinnamon, D.; Ghanta, R.; Lin, K. C.; Muthukumar, S.; Prasad, S. Portable biosensor for monitoring cortisol in low-volume perspired human sweat. Sci. Rep. 2017, 7, 13312. (18) Moyer, J.; Wilson, D.; Finkelshtein, I.; Wong, B.; Potts, R. Correlation between sweat glucose and blood glucose in subjects with diabetes. Diabetes Technol. Ther. 2012, 14 (5), 398−402. (19) Lee, H.; Choi, T. K.; Lee, Y. B.; Cho, H. R.; Ghaffari, R.; Wang, L.; Choi, H. J.; Chung, T. D.; Lu, N.; Hyeon, T.; Choi, S. H.; Kim, D.H. A graphene-based electrochemical device with thermoresponsive microneedles for diabetes monitoring and therapy. Nat. Nanotechnol. 2016, 11, 566−572. (20) Munje, R.; Muthukumar, S.; Prasad, S. Lancet-free and label-free diagnostics of glucose in sweat using Zinc Oxide based flexible bioelectronics. Sens. Actuators, B 2017, 238, 482−490. (21) Abellán-Llobrega, A.; Jeerapan, I.; Bandodkar, A.; Vidal, L.; Canals, A.; Wang, J.; Morallón, E. A stretchable and screen-printed electrochemical sensor for glucose determination in human perspiration. Biosens. Bioelectron. 2017, 91, 885−891. (22) Witkowska-Nery, E.; Kundys, M.; Jeleń, P. S.; Jönssoń M. Electrochemical glucose sensing: is there still room Niedziołka, for improvement? Anal. Chem. 2016, 88 (23), 11271−11282. (23) Olarte, O.; Chilo, J.; Pelegri-Sebastia, J.; Barbe, K.; Van Moer, W. Glucose detection in human sweat using an electronic nose. Conf. Proc. IEEE Eng. Med. Biol. Soc. 2013, 1462−1465. (24) Sakaguchi, K.; Hirota, Y.; Hashimoto, N.; Ogawa, W.; Hamaguchi, T.; Matsuo, T.; Miyagawa, J. I.; Namba, M.; Sato, T.; Okada, S.; Tomita, K.; Matsuhisa, M.; Kaneto, H.; Kosugi, K.; Maegawa, H.; Nakajima; Kashiwagi, H. A. Evaluation of a Minimally Invasive System for Measuring Glucose Area under the Curve during Oral Glucose Tolerance Tests: Usefulness of Sweat Monitoring for Precise Measurement. J. Diabetes Sci. Technol. 2013, 7, 678−688. (25) Bandodkar, A. J.; Jeerapan, I.; Wang, J. Wearable chemical sensors: Present challenges and future prospects. ACS Sens. 2016, 1, 464−482. (26) Oliver, N. S.; Toumazou, C.; Cass, A. E. G.; Johnston, D. G. Glucose sensors: a review of current and emerging technology. Diabetic Med. 2009, 26, 197−210. (27) McCaul, M.; Glennon, T.; Diamond, D. Challenges and opportunities in wearable technology for biochemical analysis in sweat. Curr. Opin. Electrochem. 2017, 3, 46−50. (28) Williams, D. L.; Doig, A. R., Jr.; Korosi, A. Electrochemicalenzymatic analysis of blood glucose and lactate. Anal. Chem. 1970, 42, 118−121. F

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX

Article

ACS Sensors their application in nonenzymatic glucose biosensors. Biosens. Bioelectron. 2011, 26, 3500−3504. (53) Skou, E. The electrochemical oxidation of glucose on platinumI. The oxidation in 1 M H2SO4. Electrochim. Acta 1977, 22, 313−318. (54) Vassilyev, Y. B.; Khazova, O. A.; Nikolaeva, N. N. Kinetics and mechanism of glucose electrooxidation on different electrode-catalysts: Part I. Adsorption and oxidation on platinum. J. Electroanal. Chem. Interfacial Electrochem. 1985, 196, 105−125. (55) Ernst, X.; Heitbaum, J.; Hamann, C. H. The electrooxidation of glucose in phosphate buffer solutions: Part I. Reactivity and kinetics below 350 mV/RHE. J. Electroanal. Chem. Interfacial Electrochem. 1979, 100, 173−183. (56) Rao, M. L. B.; Drake, R. F. Studies of electrooxidation of dextrose in neutral media. J. Electrochem. Soc. 1969, 116, 334−337. (57) Toghill, K. E.; Xiao, L.; Phillips, M. A.; Compton, R. G. The non-enzymatic determination of glucose using an electrolytically fabricated nickel microparticle modified boron-doped diamond electrode or nickel foil electrode. Sens. Actuators, B 2010, 147, 642− 652. (58) Chen, Z. − L.; Hibbert, D. B. Simultaneous amperometric and potentiometric detection of sugars, polyols and carboxylic acids in flow systems using copper wire electrodes. J. Chromatogr. A 1997, 766, 27− 33. (59) Zadeii, J. M.; Marioli, J.; Kuwana, T. Electrochemical detector for liquid chromatographic determination of carbohydrates. Anal. Chem. 1991, 63, 649−653. (60) De Mele, M. F. L.; Videla, H. A.; Arvia, A. J. Potentiodynamic Study of Glucose Electro-Oxidation at Bright Platinum Electrodes. J. Electrochem. Soc. 1982, 129, 2207−2213. (61) Vassilyev, Y. B.; Khazova, O. A.; Nikolaeva, N. N. Kinetics and mechanism of glucose electrooxidation on different electrode-catalysts: Part II. Effect of the nature of the electrode and the electrooxidation mechanism. J. Electroanal. Chem. Interfacial Electrochem. 1985, 196, 127−144. (62) Adzic, R. R.; Hsiao, M. W.; Yeager, E. B. Electrochemical oxidation of glucose on single crystal gold surfaces. J. Electroanal. Chem. Interfacial Electrochem. 1989, 260, 475−485. (63) Bard, A. J.; Faulkner, L. R. Electrochemical Methods: Fundamentals and Applications, 2nd ed.; John Wiley & Sons: New York, 2001. (64) Bard, A. J.; Parsons, R.; Jordan, J., Eds. Standard Potentials in Aqueous Solution; Marcel Dekker: New York, 1985. (65) Larew, L. A.; Johnson, D. C. Transient generation of diffusion layer alkalinity for the pulsed amperometric detection of glucose in low capacity buffers having neutral and acidic pH values. J. Electroanal. Chem. Interfacial Electrochem. 1989, 264, 131−147. (66) Jensen, M. B.; Johnson, D. C. Fast wave forms for pulsed electrochemical detection of glucose by incorporation of reductive desorption of oxidation products. Anal. Chem. 1997, 69, 1776−1781. (67) Johnson, D. C.; LaCourse, W. R. Liquid chromatography with pulsed electrochemical detection at gold and platinum electrodes. Anal. Chem. 1990, 62, 589−597. (68) Schmid-Wendtner, M. H.; Korting, H. C. The pH of the skin surface and its impact on the barrier function. Skin Pharmacol. Physiol. 2006, 19, 296−302. (69) Hughes, S.; Johnson, D. C. Amperometric detection of simple carbohydrates at platinum electrodes in alkaline solutions by application of a triple-pulse potential waveform. Anal. Chim. Acta 1981, 132, 11−22. (70) Garcia, C. D.; Ortiz, P. I. BHA and TBHQ quantification in cosmetic samples. Electroanalysis 2000, 12, 1074−1076. (71) LaCourse, W. R. Pulsed Electrochemical Detection in HighPerformance Liquid Chromatography; Wiley: New York, 1997. (72) Garcia, C. D.; Henry, C. S. Direct determination of carbohydrates, amino acids, and antibiotics by microchip electrophoresis with pulsed amperometric detection. Anal. Chem. 2003, 75, 4778−4783.

(73) Bindra, D. S.; Wilson, G. S. Pulsed amperometric detection of glucose in biological fluids at a surface-modified gold electrode. Anal. Chem. 1989, 61, 2566−2570. (74) Surareungchai, W.; Deepunya, W.; Tasakorn, P. Quadruplepulsed amperometric detection for simultaneous flow injection determination of glucose and fructose. Anal. Chim. Acta 2001, 448, 215−220. (75) Hoyer, B.; Loftager, M. Suppression of the chloride interference effect on solid-state cupric ion selective electrodes by polymer coating. Anal. Chem. 1988, 60, 1235−1237. (76) Park, S.; Park, S.; Jeong, R.-A.; Boo, H.; Park, J.; Kim, H. C.; Chung, T. D. Nonenzymatic continuous glucose monitoring in human whole blood using electrified nanoporous Pt. Biosens. Bioelectron. 2012, 31, 284−291. (77) Hendricks, J. O. Industrial fluorochemicals. Ind. Eng. Chem. 1953, 45, 99−105. (78) Ghilane, J.; Hapiot, P.; Bard, A. J. Metal/polypyrrole quasireference electrode for voltammetry in nonaqueous and aqueous solutions. Anal. Chem. 2006, 78, 6868−6872. (79) Wu, H.; Wang, J.; Kang, X.; Wang, C.; Wang, D.; Liu, J.; Aksay, I. A.; Lin, Y. Glucose biosensor based on immobilization of glucose oxidase in platinum nanoparticles/graphene/chitosan nanocomposite film. Talanta 2009, 80, 403−406.

G

DOI: 10.1021/acssensors.8b00168 ACS Sens. XXXX, XXX, XXX−XXX