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Non-Enzymatic Wearable Sensor for Electrochemical Analysis of Perspiration Glucose Xiaofei Zhu, Yinhui Ju, Jian Chen, Deye Liu, and Hong Liu ACS Sens., Just Accepted Manuscript • DOI: 10.1021/acssensors.8b00168 • Publication Date (Web): 16 May 2018 Downloaded from http://pubs.acs.org on May 17, 2018

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Non-Enzymatic Wearable Sensor for Electrochemical Analysis of Perspiration Glucose Xiaofei Zhu, Yinhui Ju, Jian Chen, Deye Liu, Hong Liu* State Key Laboratory of Bioelectronics, School of Biological Science and Medical Engineering, Southeast University, Nanjing 210096, China Supporting Information Placeholder ABSTRACT: We report a non-enzymatic wearable sensor for electrochemical analysis of perspiration glucose. Multi-potential 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, non-enzymatic, real-time monitoring

Wearable electronic devices equipped with physical sensors that can monitor physical activity and vital signs have been 1-5 commercially available for years. But these devices only provide limited information regarding human health and their sensitivity to human physiological states is usually too 6-8 low for practical application. So last several years have witnessed a trend to integrate sensors for chemical analysis onto wearable devices that can at molecular levels provide insight into human health state for applications such as early 9-13 diagnostics. For all kinds of wearable chemical sensing, 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 continuous14,15 ly monitored. For perspiration chemical sensing, electrochemical detection method is almost exclusively involved because it is simple, sensitive, intrinsically quantitative and consumes rela10,16,17 tively low electrical power. Sensors for perspiration glucose monitoring has attracted much attention and has been reported by several researchers owing to the growing world10,18-22 wide problem of diabetes. 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 18,19,23,24 early screening and even self-monitoring of diabetes. Almost all current wearable glucose sensor used an enzyme (e.g. glucose oxidase or dehydrogenase) to convert 25-27 glucose to detectable electrochemical signal. The enzyme is a dispensable reagent for disposable test strip in personal 28-30 blood glucose test. But for wearable sensing, the use of enzyme leads to several critical problems. First, as a biomolecule, the activity of enzyme can be affected by a range of factors such as temperature, pH and ionic strength which 31 can hardly be controlled under in-situ conditions. Second, enzyme can be degraded. Even for glucose oxidase that is

Figure 1. (a) Perspiration glucose analysis during physical activity using the wristband-based electrochemical sensor. (b) A 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 quasi-reference electrode (R). (c) A photograph of the smartphone with App for the perspiration analysis. The sensor was connected to the smartphone via Bluetooth. highly glycated and thus stable, its activity can gradually decrease with time which affects its shelf life and the ability 32-36 for long-time wearable monitoring. Third, the immobilization of enzyme on electrode involves processes such as covalent cross-linking, polymerization or sol-gel entrapment on the electrode surface that not only suppress the activity of the enzyme but the immobilized reagents on the electrode 37 may slow down the electron transfer for detection. Finally, commercially available enzyme is usually bio-sourced agent

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only for in-vitro analysis, it may not be safe for wearable 32,38,39 sensing. To solve these problems, it is highly desirable to develop a wearable sensor for non-enzymatic electrochemical detection of glucose. Actually, non-enzymatic electrochemical glucose detection has been investigated for decades. For these sensors, electrocatalytic materials of glucose including 40-42 43,44 45,46 bulk metal, metal oxide, alloy, metal nanomateri47,48 49-52 als and carbon nanocomposites were used. Despite decades of research, one of the most important issues to use them for wearable sensing is that the electrocatalytic oxidation of glucose usually requires an alkaline condition (pH > 11) 53-58 for acceptable selectivity, sensitivity and reproducibility. At physiological pH of body fluid (e.g. perspiration), a dramatic decrease or even complete loss of the electrochemical 59 response was observed. Besides, detection of glucose at non-alkaline physiological conditions suffer from serious electrode poisoning or competitive adsorption of chloride, 54,60-62 sulfides and phosphates. Here we report a non-enzymatic wearable sensor for electrochemical perspiration glucose analysis. Multi-potential 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 realtime 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. RESULTS AND DISCUSSION Au was used as the working electrode owing to its high electrocatalytic activity towards 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 measuring 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 pH. This is reasonable because the negative current step induces the reduction of proton on the surface of the electrode, low concentration of proton at high pH resulted in large overpotential for proton 63,64 reduction. 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.

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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

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. 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 towards the bulk solution. The pH 68 range of human perspiration is from 4.0 to 7.0. 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 have 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 near the electrode surface which was much smaller in volume than the bulk solution. So it should not cause irritation to 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

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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).

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. 65-67

As previously reported, the anodic wave at −0.4 V was owing to oxidation of the carbonyl group of glucose, and the anodic wave at 0.2 V was because of 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 po54 tential, when the passivating oxide layer was removed. 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 69,70 passivate the electrode. For continuous wearable sensing,

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 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. 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 71-74 clean the electrode surface. So for wearable sensing of perspiration glucose, multi-potential steps were now applied to the Au working electrode (Figure S1 in the Supporting Information). The multi-potential 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 multi-potential 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 75 electrode surface. The Kel-F membrane is a type of fluorocarbon materials, which can further repel charged molecules


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(e.g. amino acids and ascorbic acid). For wearable sensing, the Kel-F has unusual and outstanding properties. It is chemically stable, which can resist fuming nitric acid, aqua regia o and continued exposure to air at 200 C. It is insoluble in most solvents. It has high softening points and extremely low 77 brittle points. The chemical stability and durability of the Kel-F layer ensured the robustness of the sensor. The effect

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ed 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 quasireference 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-time wearable sensing process, no irritation to skin was observed by any of the volunteers. Results indicated that during the two weeks period of time, the RSD for perspiration analysis was about 4.6 %, which demonstrated the reproducibility and practical applicability of the sensor. Conclusions

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. of interferents in real perspiration such as ascorbic acid, uric acid, lactic acid and glutamic 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 amount of glucose was 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 -1 30 to 1100 μM. The sensitivity of the sensor was 114 μA mM -2 cm .The limit-of-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. 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 includ-

We have reported the non-enzymatic wearable sensor for electrochemical analysis of perspiration glucose. Multipotential steps were involved for electrocatalytically 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 activities, 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 HPLCMS. 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. 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), 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 threedimensional printer (M-Jewelry) was used to print the wrist-


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band. CHI 8520d workstation from Chenhua (Shanghai, 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 oC. Citrate buffered solution with different pHs, including 2.2, 3.0, 4.0, 5.0, 6.0, 7.0, 8.0, were prepared. Threeelectrode 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 seconds. A current of -10 µA was applied for 200 seconds. For the pretreatment that change the pH of the solution, a negative potential of -2.0 V applied for 20 seconds before the chronopotentiometric pH measurement as mentioned above. Glucose Detection. Glucose detection in 0.10 M PBS (pH 5.0) was conducted by applying multi-potential steps to the working electrode. Three potential steps were involved, including a high negative potential of -2.0 V for 20 seconds, a moderate potential of 0.2 V for 5 seconds and a positive potential of 1.0 V for 2 seconds. Preparation of Pseudo-Reference Electrode. For wearable sensing, a polypyrrole pseudo-reference 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 three-electrode cell with a Ag/AgCl reference electrode, 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 report78 ed (Figure S3 in the Supporting Information). 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 ultrasonically cleaned 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, 5% Nafion and 8% Kel-F oil mixture were mixed together by 1 : 2 rato, and was dropcast on the electrode. 76 The electrode was dried in an oven for 5 h at 60 ◦C. 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 con-

stant potential of −0.25 V in a 5.0 mL solution containing 1.0 79 mM H2PtCl4 and 0.50 M H2SO4. For wearable sensing, 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 three-dimensional 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 (3.7V, 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

Supporting Information Information about the fabrication and experiments. This material is available free of charge on the ACS Publications website. AUTHOR INFORMATION

Corresponding Author [email protected]

Notes The authors declare no competing financial interests. ACKNOWLEDGMENT 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)


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