3D-PMED (Three Dimensional Paper-Based Microfluidic Electrochemical Integrated Device) For Monitoring Of Sweat Metabolites
In the past few years, wearable electrochemical sensors have gained great importance, and this study aims to determine the use of 3D-PMED (three-dimensional paper-based microfluidic electrochemical integrated device) for observing sweat metabolites in real time. The construction of the 3D-PMED device was based on wax patterns screen printed on cellulose paper. This pre-patterned paper is transformed into a five-layered stack of paper by folding it four times. The first layer is the sweat collector, followed by the vertical channel layer; next comes the transverse layer.
The fourth is the electrode layer, and finally, the sweat evaporator. The sweat monitoring device took place through the integration of polyethylene terephthalate (PET) substrate with glucose screen printed sensors along with the newly invented 3D-PMED. Red ink was in use to model the process of sweat flow to demonstrate the 3D-PMED’s capability of collection, analysis, and evaporation of sweat owed to the filter papers’ capillary action and the wax’s hydrophobicity. With a high sensitivity of 35.7 µA mM-1 cm-2 and a low 5µM detection limit, the glucose sensor was designed considering the low concentration of glucose level in sweat.
To validate the sweat 3D glucose monitoring device, active on-body experimentation was carried out, and the same was advised to be readily expanded for continuous monitoring of alternate metabolites and electrolytes.
INTRODUCTION
The potential advantages and applications of wearable sensors in personalized medicines have garnered ample attention lately, with their ability to continuously monitor individual situations of health in real-time. (1-4). A vast range of wearable devices targeted to monitor blood oxygenation, heart rate, and electrocardiograms, besides several others, are already commercially available in the market and are increasingly being used by patients. 5 The same, however, is not applicable to chemical sensors, as they are not as popular or widely developed purely owed to the complicated and intricate principles of operations and structures of fabrication. (6,7).
This, however, does not mean that progress is limited in this area; instead, various research groups have recently developed different chemical wearable devices that can measure levels of biochemical analytes in sample sweat. (8,9). Some examples include epidermal-based sweat sensors for glucose analysis, 10 lactate 11, and pH 12 levels. This has been done by the J. Wang Group.
The same group has also successfully developed a sensing hybrid device that continuously monitors the wearer’s electrocardiogram and lactate. (13). Using screen printing technology, the sensor patches were designed on flexible polyester or paper used for tattoos. Gao et al. further proposed that by using a flexible PET (Polyethylene Terephthalate) substrate, a wearable incorporated sensor array can be designed by using deposition methods of electron beam and photolithography. 14 This incorporated patch, when worn, offers the ability to continuously monitor the electrolytes and the metabolites present in the sweat of a human. The pH and Ca2+ concentration measurements in body fluids can also be achieved by designing PET substrates using ion-selective electrodes. 15
However, one limitation of previous studies has been their prime consideration related to epidermal devices, focusing only on their stretchable and flexible properties. 16,17 There needs to be more effort placed on analyzing sweat fluid flow under the patch, as it is empirical for the simultaneous real-time sweat analysis and related comfortability factor in the long run for wearables. It should also be considered that the accumulation of sweat under sensors will lead to skin discomfort for the wearer. 18 To date, all paper-based monitoring devices targeted toward monitoring sweat metabolites have been proposed for their intrinsic characteristics of rollability, foldability, flexibility, ease of use, and low cost. 19-21 Furthermore, paper-based devices facilitate the flow of sweat owed to the capillary effect, thus contributing to sweat collection. 22
This study develops a 3D-PMED (three-dimensional paper-based microfluidic electrochemical integrated device) using a cellulose pre-patterned paper folding technique designed using modified wax-based screen printing. The 3D paper-based microfluidic electrochemical integrated device contains five layers. The first layer is the sweat collector. The second is the vertical channel, and the third is the transverse channel, the fourth is the electrode layer, and the sweat evaporator layer (Fig A)41. Additionally, the designed screen-printed electrodes were attached to the device’s electrode layer to measure the sweat analytes, thus allowing the patterned folded paper to form a flow channel (Fig B)41.
The collector absorbed the sweat from the skin and flowed to the next vertical channel layer through the capillary driving mechanism. The sweat evaporator ensured simultaneous sweat flow driven by the evaporation process, providing fresh sweat flow to the electrodes and eliminating the accumulation of sweat within the device. The 3D-PMED also had the ability to avoid direct sensor-to-skin mechanical contact by separating electrodes from skin contact. A glucose sensor was also designed in the 3D-PMED to monitor glucose levels in sweat samples. The glucose sensor is checked extensively for selectivity, repeatability, and sensitivity.
Furthermore, the processes of sweat flow were modeled to confirm simultaneous sweat flow by using red ink in the device. The final step was to monitor the glucose level through the sweat formed on the subjects’ skins after subjecting them to endure physical exercise for a sustained period (e.g., cycling). This was done to validate the 3D-PMED device’s practicability. The sweat glucose corresponding temporal profiles were then recorded using the amperometric method.
Fig A
Fig B[1]
[1] The 3D-PMED schematic diagram illustrated in Fig A defines the device’s layered structure. It includes five folds respectively 1) sweat collector, 2) vertical channel, 3) transverse channel, 4) electrode layer, and 5) sweat evaporator. The hydrophobic areas are highlighted in white and yellow formed using filter paper or wax screen printing respectively. Fig B illustrates the schematic diagram of the device on human skin. The tiered structure of the 3D flow through the device’s folds facilitates the fresh continual flow of sweat in the device from the skin under the electrodes.
Methods and Materials Used
Securing Of Chemicals
Aladdin Ltd was contacted for obtaining glucose oxidase (GOx) obtained from Aspergillus niger, ascorbic acid and L-lactate acid. Acetic acid, glucose, chitosan, and sodium chloride were secured from the Hushi Laboratorial Equipment Co, Ltd in Shanghai. 0.01 M PBS (phosphate buffered saline) was obtained from Sangon Biotech Co Ltd in Shanghai. Gwent Inc, situated in Torfaen, in the UK, supplied Graphite/Prussian blue ink, while Yingman Nanotech Ltd, located in Jiangsu province of China, helped obtain silver chloride/silver ink. NEXAR was acquired from XXX. No further purification was conducted on all reagents used. Jujo Chemical Co Ltd in Tokyo, Japan, was consulted for the purchase of the Insulator ink.
Fabrication Process Of the Device
The patterned cellulose paper, altered by wax screen printing, was used for the device. CorelDRAW was used to design the patterns for the 3D-PMED, further being outsourced for screen mesh fabrication. The pattern shapes were highlighted on the screen under white areas, depicting the penetration of wax to the hydrophobic regions through the mesh. Initially, the back of the 100 mesh nylon formed screen was covered by filter paper employing pasting to restrict the sliding of the piece relative to the mesh (Fig D). The screen mesh and the filter paper were rubbed over by a crayon over a hot plate. (Fig C). The heating made it easier for the wax to penetrate the paper to create well-defined hydrophobic barriers on the paper. The temperature was set at an appropriate 50ºC to 55ºC.
A temperature lower than those mentioned above restricted the wax from penetrating the filter paper, while a higher temperature was found to create enhanced lateral flow creating fuzzy patterns23.
In about 4 minutes, the filter paper had a distinct wax area formation. The filter paper with the wax pattern was then cut into a rectangular shape of 3cm by 15cm, considering the pattern edges, and folded into five equal 3cm into 3cm parts, forming the five-tier structure. As mentioned earlier, the screen mesh fabrication was outsourced, while the CorelDRAW-designed electrode patterns were used following previous works done in the area with slight changes.24 The process included first using silver/silver chloride ink before using Graphite/Prussian blue ink before the blue insulator layer. A curing period of 15 min at a temperature of 90º C was observed throughout each printing. Later pasting of electrodes was done over the electrode layer. The round working electrode has a diameter of 2.5 mm, whereas the counter electrodes and reference were circled in part with a respective area of 19.4mm2 and 2mm2.
Fig C
Fig D
Modification Of Electrode
Modification of glucose sensors was done as follows.14,25 2µL of Nexar: Ethanol = 1:10 was released on the surface of the working electrode and left to dry for 30 minutes at an ambient temperature. Further, 5% chitosan solution measuring 3µL was cast in 3% acetic acid and left to rest for 8 hours at 4ºC. Also, GOx (10mg mL-1) measuring 2µl was released on the surface of the working electrode and was left to dry for a period of two hours at an ambient temperature before releasing over it the 5% chitosan and 3% acetic acid on it. Lastly, 2µl of NEXAR: Ethanol 1:10 was released on the surface of the working electrode and left to dry out before storing at 4ºC as standby. (Fig E)
Fig E
Electrochemical Assays
All experiments performed were done within a chemical workstation. Amperometric sensors used contained 3 electrode systems, and all counter and working electrodes were printed using Graphite or Prussian ink with an Ag/AgCl electrode reference. All measurements of in vitro amperometric solutions in the experiments utilized 0.01 M PBS. The utmost effective catalyst for the electro-reduction of hydrogen peroxide, in terms of modified Prussian Blue electrodes, is amperometric detection26. The Prussian Blue modified molecules have a cubic geometry shape rendering their electrochemical selectivity as being practical. Instead, any molecule with a higher weight than the H2O2, e.g., uric acid or ascorbic acid, cannot penetrate the Prussian Blue lattice, thus offering a reaction resembling catalytic redox.
These advantages above have successfully been used for acquiring interference-free and sensitive probes for the detection of H2O2. 27 Since H2O2 has the overpotential to accelerate electrocatalytic oxidation of interferents in sweat, such as ascorbic acid and uric acid, in conventionally printed electrodes, in the fabrication process, Prussian Blue modified printed biosensor was used to decrease the potential of consequent electrochemical interferences. 28 To achieve on-body measurements and calibration curves, -0.1V (against an Ag/AgCl) applied potential was used in the amperometric detection. 25, 26, 28, 29
[1] Fig R1 – The current amperometric response at -0.1V (reference electrode versus Ag/AgCl) with a 0.1 mM glucose successive addition with concentration ranging between 0mM – 1.6mM in M PBS solution 0.01; The calibration plot (inset).
Fig R2[3]
Figure R 3[4]
Fig R-4[5]
Figure R5[6]
Experiment Results
Glucose Sensor Characteristics
The concentration of glucose present in sweat is much lower compared to blood; however, it is still an effective glucose level indicator in humans. 30-33 Having said the low concentration between 0.25 – 1.5 mM range also makes developing practical sweat-based glucose sensors difficult.
In this study, many major analytical parameters had to be evaluated beforehand to successfully achieve a sweat-based glucose sensor that is high performing in terms of repeatability, linearity, and sensitivity. 34 It was also imperative that the glucose sensor offered a reproducible and selective response, high sensitivity, and a low detection limit. The core of the glucose sensor was the Glucose Oxidase (GOx), an enzyme vastly used to enhance glucose sensors. Glucose Oxidase simultaneously catalyzes the process of oxidation in glucose, transforming it into gluconolactone while generating H2O2, whereas Prussian Blue directly reduces H2O2 electrochemically. Prussian Blue is effective against selectively detecting H2O2 in oxygen presence and other interferents acting as a redox mediator, owed to its geometrical shape and extremely low -0.1 V (versus Ag/AgCl reference) 35 – 37 catalytic potential.
In Fig R1, the glucose sensor amperometric response is illustrated after the addition of 10 µl 1 M drops of glucose with concentrations of 0 – 1.9 mM, using a beaker containing 10 mL 0.01 M PBS. The measurement process and system were controlled at an ambient temperature. The fig represents a good linearity of 0-1.9 mM glucose concentrations depicting a correlated 0.99 coefficient covering the human sweat physiological glucose concentration range (0.25 – 1.5 mM). 38 The glucose sensor sensitivity touched the 35.7µA mM-1 cm-2 (the electrode area consisted of 4.906 mm2) offering quick response. The glucose sensor’s detection limit was analyzed with current amperometric responses after successive addition of 5, 10, 20, and 40 µM.; the lowest detection limit was computed AT 5 µm.
The human perspiration existing interferents (e.g., 0.1 mM NaCl, 2 µM UA, 5 mM lactate, and 2 µM AA) influence was examined using the sensor response. No significant signals of interference were found compared to the drops of glucose. The A fluctuating curve was achieved with the high volume lactate addition, which was acceptable in this study. The reproducibility of the inter-electrode in the glucose sensor was also evaluated, with the five sample-based standard deviations not being more than 10%. This was taken as satisfactory reproducibility considering the low-cost route of electrode fabrication. This thus concluded the good performance of the adapted glucose sensor.
3D PMED Fluid Flow Modelling
The sweat fluid flow modeling using the 3D PMED was done by releasing red ink on the PMED and monitoring the flow process in real time. Taking two plastic plates and placing the 3D PMED in between with holes on both plates to ensure the entering and exporting of red ink, the first and last layer, sweat collector and sweat evaporator, were aligned with the holes created in the two plates. Securing them using two clips on both sides. Gradual successive red ink addition on the sweat collector surface was done following the 50, 100, 150, 200, and 250 µL measurements, respectively—the filter paper’s capillary effect limited ink flow within a second. The later unfolding of the paper revealed a complete flow of red ink through the layered channels from the first to the fifth layer.
Clear boundary creation was witnessed between the hydrophilic channels and hydrophobic barriers. The first drop of 50µL remained on the first layer of the sweat collector, only covering a vast area. This layer was intended to collect the maximum possible perspiration. As the drops of ink increased, it gradually penetrated all channels reaching the next layers until all chambers were filled. This revealed the evidence that the 3D PMED has the capacity and ability to catch perspiration from human skin. The 3D PMED can thus prove to be a promising sweat accumulation-based sensor.
3D – PMED Based Glucose Sensing Experiments
As mentioned earlier, the 3D-PMED integration with glucose-based sensors was measured in real-time by absorbing sweat from human skin. The experiment was conducted by incorporating the device between two plates made of plastic, with the glucose sensor integrated within. The sensor was further connected to the workstation, which was analyzed electrochemically. Fig R4 depicted the current amperometric responses at -0.1 V (versus Ag/AgCl reference electrode) the current amperometric responses after adding glucose successively in the following quantities: 0.5, 1.0, 1.5, 2.0 mM, and PBS. The fluid’s evaporation was relatively slow. Thus, another filter block paper was placed on the last layer of the sweat evaporator so that increased fluid flow could be made possible. With a break of 500s between adding new fluid, the calibration plot was depicted as linear with a 17.49µA mM-1 cm-2 sensitivity. Thus good concentration linearity and repeatability were achieved, ranging between 0-1.9 mM.
The glucose sensor sensitivity in the device was, however, decreased owing to the difference in the 3D PMED device and the beaker. The Standard deviation, however, lay between 0.02 – 0.12, which was acceptable as the performance of the device’s sensor. It was considered suitable for glucose monitoring from sweat samples. The 3D-PMED device and the glucose sensor were then analyzed by integrating it as a monitoring device that can be practically used to measure glucose levels on human skin. This was done by designing a ventilated band that prevented delamination and protected the device from folding into a 60mm into 80mm device attached to the forearm of the selected subject.
Consent was achieved before the experiment from all issues involved, and they were informed that the findings would be used for research work only. Using a cycle ergometer, practical body measurement of glucose through sweat was conducted using the electrochemical workstation. After a 10-minute warm-up, a 25 minutes cycling session at a constant speed and a 10 minutes interval were set as a protocol for all subjects.
All involved subjects were also asked to drink honey water before exercising with a honey-to-water volume ratio of 1:19 so that blood glucose could be raised. The measurement started after the 10-minute warm-up exercise, aiming for subjects to acquire enough sweat during the measurement time. The cycle ergometer calculated all subjects’ calorie and heart rates in real-time. The heart rate naturally rose after warm-up, with the calorie also rising linearly owed to constant exercising intensity; the consumption of blood glucose during exercise gradually lessened till the end of the practice ranging between 1.5mM to 0.04mM.
The heart rate recovered during the break between cycling, while the level of glucose also increased a bit, leveling off the blood glucose regulation. Since the level of metabolism and glucose differ among subjects, there was a variance in the heart rate of all issues. The results revealed an active change in the glucose concentration during the exercise. Besides the differences occurring in the three subjects, the glucose change concentration in sweat showcased the exercise throughout the process. This also proved that the 3D-PMED showed decent performance and can be actively used as a form of on-body-based measurement.
Conclusions
The 3D-PMED device was initially proposed to be constructed and used as a wearable device to monitor glucose levels through sweat. The device is designed by channeling a 3D flow using the five-folding technique on a pre-patterned paper filter. The first layer absorbs the sweat on the skin, the sweat collector, further flowing into the second and third layers of the vertical channel and transverse channel, respectively, using the capillary driving technique, flowing further into the electrode layer and the final sweat evaporator layer. The device is precisely designed for the collection and circulation of sweat and the sensor isolation to monitor sweat. The proposed 3D-PMED device effectively controlled sweat flow and successfully accumulated sweat.
Further on, the screen-printed-based glucose sensor helps detect glucose from sweat. The 3D-PMED successfully proved to be a practical and usable device with essential characteristics: low cost and easy and straightforward use. A few challenges appeared during the experiment and remained unsolved, such as the sensor stability and the device’s size of channeling. Future experiments will see a downsizing of the device to enhance wearing comfortability and further addition of pH electrodes and sodium and lactate sensors. The current need, however, is to develop a printed flexible circuit board to facilitate wireless monitoring through the device.
Interest Conflicts
No Conflicts of Interests Are to Be Declared.
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[1] The 3D-PMED schematic diagram illustrated in Fig A defines the device’s layered structure. It includes five-folds, respectively 1) sweat collector, 2) vertical channel, 3) transverse channel, 4) electrode layer, and 5) sweat evaporator. The white and yellow hydrophobic areas are using filter paper or wax screen printing. Fig B illustrates the schematic diagram of the device on human skin. The tiered structure of the 3D flow through the device’s folds facilitates the fresh continual flow of sweat in the device from the skin under the electrodes.
2) Figure R1 – The current amperometric response at -0.1V (reference electrode versus Ag/AgCl) with a 0.1 mM glucose successive addition with concentration ranging between 0mM – 1.6mM in M PBS solution 0.01; The calibration plot (inset).
[3] Fig R2 – Glucose sensor inter-electrode reproducibility table
[4] Fig R3 – amperometric current responses based selectivity experiment of the biosensor with successive glucose addition, furthered by interfering electroactive species of NaCl and last glucose in 0.01 M PBS at -0.01 V applied potential.
[5] Fig R4 – Calibration plot with 0mM to 2.0mM error bar range
[6] Figure R5 – Measurement of on-body glucose in sweat using 3D-PMED
