Zhaokang Zhanga,
Lu Huangb,
Yiting Chenb,
Zhenli Qiub,
Xiangying Meng*c and
Yanxia Li*b
aCollege of Chemical Engineering, Fuzhou University, Fuzhou, 350108, China
bCollege of Materials and Chemical Engineering, Minjiang University, Fuzhou, 350108, China. E-mail: yxli09@163.com
cSchool of Medical Laboratory, Weifang Medical University, Weifang, 261053, China
First published on 2nd January 2024
In this work, a portable electrochemical glucose sensor was studied based on a laser-induced graphene (LIG) composite electrode. A flexible graphene electrode was prepared using LIG technology. Poly(3,4-ethylene dioxythiophene) (PEDOT) and gold nanoparticles (Au NPs) were deposited on the electrode surface by potentiostatic deposition to obtain a composite electrode with good conductivity and stability. Glucose oxidase (GOx) was then immobilized using glutaraldehyde (GA) to create an LIG/PEDOT/Au/GOx micro-sensing interface. The concentration of glucose solution is directly related to the current value by chronoamperometry. Results show that the sensor based on the LIG/PEDOT/Au/GOx flexible electrode can detect glucose solutions within a concentration range of 0.5 × 10−5 to 2.5 × 10−3 mol L−1. The modified LIG electrode provides the resulting glucose sensor with an excellent sensitivity of 341.67 μA mM−1 cm−2 and an ultra-low limit of detection (S/N = 3) of 0.2 × 10−5 mol L−1. The prepared sensor exhibits high sensitivity, stability, and selectivity, making it suitable for analyzing biological fluid samples. The composite electrode is user-friendly, and can be built into a portable biosensor device through smartphone detection. Thus, the developed sensor has the potential to be applied in point-of-care platforms such as environmental monitoring, public health, and food safety.
Graphene is a sp2 hybrid carbon nanosheet that offers high conductivity, specific surface area, and biocompatibility, making it a popular choice in biosensing devices. Methods used for preparation include mechanical stripping, chemical vapor deposition, and redox methods, which all have significant drawbacks, including high costs, long preparation times, and environmental pollution, which limit large-scale manufacturing.6,7 However, laser-induced graphene (LIG) technology is a promising alternative as it uses laser irradiation on a PI surface, causing rapid temperature increases that break C–O, CO, and N–C bonds and release CO, C2H2, and other gases. This process rearranges aromatic groups, transforming sp3 hybrid carbon atoms into sp2 hybrid carbon atoms and creating the basic structure of graphene, allowing for rapid preparation at a low cost.8,9 Composites that incorporate graphene, conductive polymers, and metal nanoparticles offer improved performance and expand the range of applications for these materials. Graphene/conductive polymer/metal nanoparticle composites can detect a variety of compounds and molecules such as glucose, urea, protein, and DNA,10–14 indicating promising prospects in the field of biosensing.
PEDOT is a conjugated polymer synthesized from 3,4-ethylene dioxythiophene (EDOT) monomer that exhibit excellent electrical conductivity and have been proven to enhance the electrochemical performance of biosensing electrode.15,16 Depositing of a PEDOT film on the LIG electrode surface via a potentiostatic method can enhance the mechanical properties, hydrophilicity and anti-pollution performance of LIG.17 Additionally, enzymatic oxidation-based amperometric glucose sensors utilizing electroactive noble metals (e.g., gold and platinum) have garnered increasing attention. Among them, gold nanoparticles exhibit exceptional performance in enzymatic electrochemical biosensing, owing to their stable chemical properties, good biocompatibility, good electrical conductivity, and high catalytic activity.18 Incorporating gold nanoparticles onto the LIG composite electrode surface can increase active sites, enhance the amount of immobilized enzyme, and improve the amperometric response range of the sensor.
As one of the important indicators of human health, glucose detection has always been a research hotspot. Investigators are committed to seeking cheaper, more accurate, minimal or non-invasive methods to quantify glucose levels in real-time. The detection of glucose concentration can be carried out through methods such as chromatography, spectroscopy, and electrochemical analysis. Electrochemical glucose sensors, which have the advantages of low detection cost, high reliability, and strong operability, have attracted much attention and have been widely studied and reported.
In recent years, with the rapid development of intelligent wearable devices, new requirements and heights have been put forward for the development and application of biosensors.19 Flexible wearability, real-time monitoring, and high stability have become important directions for the development of glucose sensor applications. Flexible wearable devices are gradually improving human life through intelligence, personalization, and convenience, and the related market demand is also constantly increasing. Although the number of components that make up a flexible electrochemical sensor may vary depending on the specific application, it typically includes: substrate unit, sensing unit, decision-making unit, and energy supply unit. Flexible film is a common substrate form, such as polyethylene terephthalate (PET), polyimide (PI).20 PET and PI have almost no stretchability, making them suitable for fixation in areas where the skin is not easily deformed. Fabric has outstanding advantages in breathability and wear comfort, making it a more ideal substrate material for wearable devices.
The field of fabrication methods for functional circuits on 3D freeform surfaces and standalone stretchable sensing platforms is rapidly evolving. Researchers are continuously exploring new materials, fabrication techniques, and design strategies to overcome challenges related to mechanical durability, electrical performance, and integration of complex functionalities.21,22 For example, additive manufacturing involves using the techniques of 3D printing or inkjet printing to directly fabricate conductive traces and components on 3D surfaces. It allows for the creation of complex circuit geometries on irregular or curved objects.23–26 Utilizing stretchable materials as the base substrate for the circuit is a popular approach. These substrates, often made of polymers or elastomers, can deform and stretch without compromising the functionality of the circuit.27 Serpentine or meandering interconnect designs can accommodate the mechanical deformation of the substrate while maintaining electrical connectivity. Moreover, using methods of stamping or lamination transferred circuit components onto the 3D freeform surface.28
The sensing unit is the core component of flexible electrochemical sensors. According to the different detection mechanisms, electrochemical glucose sensors can be divided into non-enzymatic sensors and enzyme-based sensors. The non-enzymatic glucose sensor does not require direct use of enzymes during the detection process, so it is expected to achieve glucose detection under broader conditions. However, conventional non-enzymatic sensors usually cannot respond specifically to glucose molecules, making detection results more susceptible to interference. Relatively speaking, enzymatic glucose sensors typically have higher specificity. However, its detection performance is severely constrained by factors such as the effective load of enzymes and enzyme activity that led to poor reproducibility and stability during use. Therefore, exploring modified methods of electrodes with good performance electrode for obtaining stable binding and high GOx loading has extremely high practical significance. In addition to typical enzymes, the active catalytic materials of glucose sensors are mainly composed of polymers, metals, alloys, metal compounds, and other substances that can catalyze the oxidation of glucose.29 LIG electrodes have been widely developed due to their excellent biocompatibility, flexibility, pattern-ability, and ability to serve as a favorable environment for enzymes and promote electron transfer.30 The core of a decision-making unit is a microcontroller unit (MCU), which converts electrical signals into digital signals through components such as signal transduction, amplification units, and analog-to-digital converters. Currently, mainstream electrochemical sensors are applied relatively independently to a single individual, using selected features to perform predefined tasks. Smartphones can easily achieve portability and real-time detection.31 With the development of machine learning technology, flexible electrochemical sensors are expected to obtain a large amount of physiological information, and create a high-dimensional and multi-level database through distributed sensing unit arrays.32,33 Traditional energy supply units typically use commercial batteries directly, posing potential risks to the health of patients. Researchers have designed and developed various new wearable energy supply units to meet the demand for flexible and intelligent wearability, such as supercapacitor, triboelectric effect, biofuel cells, etc. Mercier et al. provide a self-powered model that utilizes glucose biofuel cells as energy sources to obtain energy from glucose in the intestine and measure glucose concentration.34
The aim of this study is to develop a portable glucose sensor based on patterned LIG composite electrode. To achieve this, PEDOT and Au NPs are deposited alternatively through a potentiostatic method, and GOx is then immobilized onto the surface of the composite electrode via chemical crosslinking. The resulting LIG/PEDOT/Au/GOx glucose biosensing interface demonstrates excellent specificity in catalyzing glucose molecules, which are converted along with oxygen in the substrate to gluconic acid and hydrogen peroxide, resulting in electron transfer on the electrode surface, and the electrochemical signals changes during this reaction process are recorded for glucose detection. What's more, the LIG/PEDOT/Au/GOx glucose biosensing interface can be combined with portable electrochemical workstation and smartphones to detect actual samples, expanding its potential applications.
Polyimide (PI) film, with a thickness of 0.08 μm, was acquired from Tianjin Jiayin Nanotechnology Co., Ltd. Meanwhile, polyethylene terephthalate (PET) substrate with a thickness of 180 μm was obtained from Nanjing Xianfeng Nanotechnology Co., Ltd. The conductive silver paste and Ag/AgCl paste were purchased from Shenzhen Ausbon Co., Ltd. and Guangzhou Yinbiao Trading Co., Ltd., respectively. Glutaraldehyde (GA) (50%) was obtained from Fuzhou Xinyuhua Experimental Instrument Co., Ltd. Glucose oxidase (GOx) with a specific activity of 248.88 U mg−1 was purchased from Sigma-Aldrich (Shanghai) Trading Co., Ltd., while bovine serum albumin (BSA) with a purity of 98.0% was acquired from Beijing Dingguo Biotechnology Co., Ltd. Additionally, artificial sweat (ISO 105/E04 pH 7.0), artificial urine, and fetal bovine serum were obtained from Yuanye Biotechnology Co., Ltd. 11-Mercaptoundecanoic acid (MUA), 3,4-ethylene dioxythiophene (EDOT) with a purity of 99%, and phosphate buffer solution (PBS) with a concentration of 0.1 mol L−1 and pH of 7.0 were all purchased from Aladdin Reagent (Shanghai) Co., Ltd. Chloroauric acid tetrahydrate (HAuCl4·4H2O) with a concentration of 47.8% was acquired from Sinopharm Chemical Reagent Co., Ltd. Furthermore, potassium chloride, potassium ferricyanide, glucose, trisodium citrate, and other reagents were of analytical pure grade. Ultrapure water with a resistivity of 18.2 MΩ was obtained using a Millipore Autopure WR600A system (Millipore Ltd., USA) and used in all experiments.
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Fig. 1 Scanning electron microscope (SEM) images of bare LIG (A and B) and LIG/PEDOT (C and D), LIG/PEDOT/Au (E and F), LIG/PEDOT/Au/GOx (BSA) (G and H) composite electrodes. |
The energy dispersive spectrometer (EDS) was used to analyze changes in surface element content of the modified electrodes for further verification of the electrode assembly process. As shown in Fig. 2A, the surface of the PI film after laser printing was rich in C element, indicating a high degree of graphitization. The laser printing process was carried out in an empty atmosphere, resulting in a small amount of nitrogen and oxygen elements being doped during modification. This phenomenon may lead to a decrease in electron transfer efficiency on the electrode surface. However, it can improve the hydrophilicity of the electrode surface and increase the number of active sites to some extent. Fig. 2B indicates an increase in oxygen content, which proves that the EDOT monomer polymerized on the LIG electrode surface. The dense distribution of Au element in Fig. 2C indicates that a significant number of Au NPs were successfully assembled on the LIG/PEDOT electrode surface. Furthermore, the content of oxygen elements increased with GOx modification on the modified electrode, as shown in Fig. 2D, owing to the rich oxygen content of GOx. The content of N and O elements significantly increased in Fig. 3 due to the introduction of a considerable number of groups containing N and O elements after fixing GOx, BSA, and GA on the electrode surface.
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Fig. 2 EDS diagram of the element distribution on the electrode surface (A) LIG electrode, (B) LIG/PEDOT electrode, (C) LIG/PEDOT/Au electrode, (D) LIG/PEDOT/Au/GOx (BSA) electrode. |
The morphology and structure of LIG, LIG/PEDOT and LIG/PEDOT/Au composites were characterized by TEM (Tecnai F30 G2 STWIN 300 kV). As illustrated in Fig. 4A, the prepared material features a distinct flake graphene structure, indicating that laser printing technology has effectively transformed PI into graphene. The surface of graphene displays translucency in both the center and edge regions, indicating that the EDOT monomer has successfully polymerized on the electrode surface through electrodeposition (Fig. 4B). Fig. 4C depicts a uniform distribution of granular material with a size of around 20 nm adhering to the electrode surface, indicating that Au NPs have been deposited onto the surface.
The structure of different electrodes was characterized using X-ray diffraction (XRD, Rigaku Smart Lab, Japan). As shown in Fig. 5A, the characteristic peak of graphene at 2θ = 25.92° was observed for all electrodes, indicating the successful graphitization of the PI film. While there is no obvious diffraction peak in the PEDOT polymer, thus it is not reflected in the XRD diagram. The reflection peaks of Au face-centered cubic crystal structure at 38.12°, 44.32°, 64.68°, 77.68° and 81.93° were observed for the LIG/PEDOT/Au and LIG/PEDOT/Au/GOx (BSA) electrodes, indicating that Au was successfully modified on the electrode surface.
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Fig. 5 (A) XRD spectra and (B) Raman spectra of LIG, LIG/PEDOT, LIG/PEDOT/Au, LIG/PEDOT/Au/GOx (BSA) composite materials. |
Raman spectroscopy is a valuable non-destructive tool for investigating the properties of carbon-based compounds. Herein, Raman spectroscopy further characterizes the graphitization effect of LIG electrode materials (Fig. 5B). Three typical graphene fingerprint vibration peaks are observed, which are G band at 1580 cm−1, D band at 1302 cm−1 and 2D band at 2594 cm−1. The intensity ratio of the G peak to the 2D peak is generally accepted to correlate with the number of graphene layers. Specifically, a higher G peak height than the 2D peak height indicates a multi-layer stacked structure in laser-induced graphene. Furthermore, the Raman scattering results of the composite electrode showed that the 2D peak disappeared, which potentially due to the change in the chemical environment and surface structure of the graphene after modification of the LIG electrode material, thereby affecting the scattering efficiency of graphene.
According to the calculation method in ref. 44, the effective area (Aeff) of electrode surface after modifications can be calculated according to Randles–Sevcik equation:
Ip = 2.69 × 105Aeffn2/3D01/2C0v1/2 | (1) |
After enzyme modification, the effective surface area of the electrode remained four times greater than its modified geometric area, which measures 0.07 cm2 (assuming π to be 3.14).
The amount of protein loaded onto the electrode surface could be calculated by Faraday formula:
![]() | (2) |
![]() | (3) |
Fig. 6B shows the cyclic voltammetry curves of LIG/PEDOT/Au/GOx (BSA) electrode in the presence of K3[Fe(CN)6] redox medium solution at different sweeping rates. The inset figure of Fig. 6B illustrates that both anodic and cathodic peak currents increased proportional with the increasing v1/2 (v is sweep rate) in the examined range of 5 to 200 mV s−1, which indicate that the surface reaction of LIG/PEDOT/Au/GOx (BSA) electrode is a typical diffusion-controlled reaction. The inset figure of Fig. 6B illustrates that both anodic and cathodic peak currents increased proportional with the increasing v1/2 (v is sweep rate) in the examined range of 5 to 200 mV s−1, which indicate that the surface reaction of LIG/PEDOT/Au/GOx (BSA) electrode is a typical diffusion-controlled reaction.
The catalytic activity of GOx towards glucose is significantly influenced by the presence of dissolved oxygen. The deoxidization experiment of LIG/PEDOT/Au/GOx (BSA) glucose sensor electrode was carried out to explore the influence of O2 on the performance of glucose sensor. The catalytic reaction process can be expressed by the following formula:1
GOx (FAD) + glucose → GOx (FADH2) + gluconolactone | (4) |
GOx (FADH2) + O2 → GOx (FAD) + H2O2 | (5) |
GOx (FADH2) ⇌ GOx (FAD) + 2e + 2H+ | (6) |
Flavin adenine dinucleotide (FAD) is an essential component of the GOx molecule that undergoes a redox reaction. Specifically, GOx (FAD) oxidizes glucose to gluconolactone acid, while the reduced GOx (FADH2) reduces O2 to H2O2 in the presence of O2. Direct electron transfer between GOx and the electrode surface results in an electrochemical response.39 Under nitrogen deoxidization conditions, GOx (FAD) is reduced to GOx (FADH2), but H2O2 production is inhibited, leading to only a limited electron transfer. As shown in Fig. 7B, the LIG/PEDOT/Au/GOx electrode generated a current signal that increased with the concentration of glucose solution in the absence of deoxidization. While under deoxygenation conditions, the current signal does not change significantly even with increasing glucose concentration. Therefore, efficient and stable glucose catalysis with GOx requires careful control of the dissolved oxygen level in the reaction environment.
The effect of potential value on chronoamperometry is to regulate the rate and direction of electrochemical reactions. Adjusting the potential value can control reaction rate and selectivity, ultimately influencing the changes in current. In order to enhance the electrochemical performance of the prepared electrode, we compared the i–t curves of glucose solutions with varying concentrations measured at different potentials. Analysis of the i–t curves in Fig. 7C reveals noticeable differences in current response under different constant potentials. The current response increases with an increase in potential until it reaches optimal response at a constant potential of 0.9 V. However, further increase of potential to 1.0 V results in a decrease of current response. It can be concluded that although high potential is beneficial for generating high current response, excessive potential may activate interfering substances and produce many intermediates, leading to irreversible reactions and inhibition of glucose oxidation. Therefore, the LIG/PEDOT/Au/GOx (BSA) electrode produces the best catalytic effect on glucose at a potential of 0.9 V.
To investigate the impact of glucose sensing in various pH and temperature environments, we conducted chronoamperometry experiments on glucose catalysis using LIG/PEDOT/Au/GOx electrodes in different detection environments. Fig. 7E shows that at pH 7.4, the LIG/PEDOT/Au/GOx (BSA) electrode exhibited the highest response to the current generated by electrochemical enzyme catalysis of glucose solution. It is well known that GOx enzyme activity is reduced in alkaline environments and enhanced in acidic environments. This is due to the conformational changes in the enzyme molecules caused by the acidic or alkaline environment, which affect GOx's ability to bind to glucose. Therefore, to maintain the optimal enzyme activity and stability of glucose oxidase, it is usually necessary to perform reactions or preservation in a neutral or slightly acidic environment. In addition, the activity of enzymes and the electrical properties of liquids are influenced by temperature. To ensure the reliability of measurement data, the influence of temperature on sensor performance should be tested (Fig. 7F). The results indicate that the enzyme activity of the electrochemical sensor may be inhibited at an ambient temperature of 23 °C, but the sensor can still maintain a high response current over a wide temperature range. Therefore, pH and temperature have an impact on the performance and practical application of enzyme electrodes. In practical applications, it is necessary to regulate the pH value with neutral buffer solutions and control the ambient temperature within the range of 27 to 37 °C. When a sensor is simultaneously affected by the pH and temperature, it can be addressed by employing a multimodal sensing decoupling mechanism. Each sensing modality can be calibrated and optimized independently to eliminate the cross-interference between pH and temperature.45,46 Calibration can be performed to establish the correlation between the output of each sensing modality of the pH and temperature, enabling the creation of suitable calibration curves or algorithms. By integrating the data from multiple sensing modalities, accurate measurements of the pH and temperature of the solution can be obtained. Through decoupling the sensing mechanisms, each modality can work independently without being influenced by other modalities, enhancing the accuracy and reliability of the sensor. In the current study, working curves can be created in different pH solution by fixing the temperature and in different temperature by fixing pH to perform calibration.
Draw a working curve between the chronocurrent values obtained at different concentrations and the logarithm of glucose concentration. It can be observed that the current response of LIG/PEDOT/Au/GOx (BSA) electrode increases exponentially with the logarithm of glucose solution concentration from 0.5 × 10−5 mol L−1 to 2.5 × 10−3 mol L−1, with an excellent sensitivity of 341.67 μA mM−1 cm−2 and an ultra-low limit of detection (S/N = 3) of 0.2 × 10−5 mol L−1. While the current response of the electrode without GOx modification remains unchanged with the increase of glucose solution concentration (Fig. 8B). Compared with commercial conventional glucose meter (Counter Plus 7600P), the present sensor has a higher sensitivity and wider detection range (Table 1). In order to further evaluate the glucose detection performance of the electrode, the comparison of electrochemical sensors for glucose detection based on enzymatic and non-enzymatic sensors was listed in Table 2. From Table 2, it can be seen that non-enzymatic sensors exhibit higher sensitivity due to the excellent catalytic performance and conductivity of nanoenzymes, and often displayed weak competitiveness in complex matrices due to less specificity of nanoenzymes than biological enzymes. Electrochemical glucose sensors based on LIG combined with nanoenzymes, exhibit ultra-high sensitivity.24,39 The non-enzymatic sensors and the present enzyme sensor both use patternable flexible LIG electrodes, promoting the development of LIG technology from a single detection component to an integrated intelligent detection system. The continuous semiconductor laser with 450 nm wavelength in this study has a lower cost and is environmentally friendly. Moreover, the mesoporous structure of LIG electrode improves the electrochemical response, and is conducive to electrode modification. The modification of Au and PEDOT on LIG electrode effectively enhances the sensitivity of glucose oxidase sensor. Furthermore, PEDOT can protect LIG mesoporous structure and have excellent anti fouling ability, making the modified electrode more stable and capable of analyzing actual samples. Research shows that it does not require a special solution environment and can analyze samples in pH and temperature environments of human sweat, urine, and blood samples, avoiding unnecessary measurement steps. In addition, the sensor shows excellent portable for real-time detection combined with smartphone.
Methods | Dynamic ranges/(mmol L−1) | Working curve equations | Correlation coefficients | LOQ/(mmol L−1) |
---|---|---|---|---|
Blood glucose meter | 0.6–33.3 | y = 5.42x + 0.65 | 0.9453 | 0.6 |
The present method | 0.005–2.5 | y = 4875.48 × exp(x/0.48) + 0.48 | 0.9932 | 0.002 |
Electrode | Recognition elements | Detection methods | Sensitivity (μA mM−1 cm−2) | LOD (μM) | Linear range | Measurement environment | Operating voltage (V) | Characteristics | Ref. |
---|---|---|---|---|---|---|---|---|---|
a RGO (reduced graphene oxide); GCE (glassy carbon electrode); FCA (ferrocenecarboxylic acid); PEG (polyethylene glycol); SWCNT (single-walled carbon nanotubes); LSGE (laser-scribed graphene electrode); RFG (reduced functionalized graphene); CS (chitosan); PB (Prussian blue); NC (nanocomposite); SPCE (screen-printed electrode); GOD (glucose oxidase). | |||||||||
Cu2O/PEDOT/RGO/GCE | Cu2O | Chronoamperometry | 4700.00 | 0.14 | 0.004–3.2 mM | 0.1 M NaOH | 0.55 | The response time was only 0.9 s, which had good stability and selectivity. No actual sample analysis was conducted | 47 |
LSGE/AuNPs | AuNPs | Linear sweep voltammetry | 0.044 ± 0.005 mA/log(M) | 6.30 | 10 μM–10 mM | 0.5 M KOH | No | The electrode preparation method is simple, low-cost, and has good selectivity. No stability and actual sample analysis was done | 48 |
Ni/Au/LIG | Ni | Chronoamperometry | 3500.00 | 1.50 | 0–1 mM | 0.05 M NaOH | 0.5 | The requirement for solution environment is relatively low, and the porous encapsulation reaction chamber can better collect sweat. Has good selectivity, but has not undergone stability analysis | 49 |
AgNRs/AuNPs–GO–CNT/LIG | Ag/AuNPs | Chronoamperometry | 1317.69 | 0.08 | 0.1–5 mM | 0.1 M NaOH | 0.7 | The porous graphene sensor modified with nanocomposites has good selectivity, can also detect pH value, and exhibits extraordinary stability, maintaining sensitivity of over 91% under environmental conditions for 21 days | 32 |
GOx/FCA/PEG/SWCNT | GOx | Chronoamperometry | 5.50 | 28.00 | 0–10 mM | PBS (pH 7.4) | 0.3 | A simple electrochemical sensing platform based on single walled carbon nanotube (SWCNT) electrodes has good selectivity and stability. No actual sample analysis was conducted | 1 |
GOx/LIG | GOx | Chronoamperometry | 43.15 | 431.00 | 0.431–8 mM | PBS (pH 7.0) | 0.8 | Simple preparation, low cost, and reliable stability. Has good selectivity and stability, but has not undergone actual sample analysis. The detection limit is high, and sensing needs further optimization | 50 |
GOx/Ti3C2Tx MXene | GOx | Cyclic voltammetry | 93.75 | 12.10 | 0.1–10 mM | PBS (pH 7.0) | No | A glucose data collection and sharing system based on cloud platform is developed, but the detection limit is high. Has good selectivity and stability, but has not undergone actual sample analysis | 51 |
GCE–RFG–RGO–GOD–CS | GOD | Chronoamperometry | 46.71 | 79.65 | 0.08–3 mM | PBS (pH 7.0) | −0.5 | It had good sensitivity with a relative standard deviation (RSD) of 4.49%. But has not undergone selectivity and actual sample analysis | 52 |
GOx/Au/PEDOT/LIG | GOx | Chronoamperometry | 341.67 | 2.00 | 0.005–2.5 mM | PBS (pH 7.4) | 0.9 | The continuous semiconductor laser with 450 nm wavelength has a lower cost and is environmentally friendly. The 3D mesoporous LIG electrode materials, combined with Au and PEDOT modification endows the sensor with excellent sensitivity, stability, and anti-fouling ability. The flexibility and simple method, combined with smart phones are expected to achieve portable real-time detection. It has good application prospect in analysis of actual samples, such as blood, sweat, and urine | This work |
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Fig. 9 i–t curves of LIG/PEDOT/Au/GOx (BSA) electrode for selectivity (A), reusability (B), stability (C); i–t curves of flexible LIG/PEDOT/Au/GOx (BSA) electrode under different bending states (D). |
Sample | Additive concentration (mmol L−1) | Current (μA) | Recovery rate (%) |
---|---|---|---|
Urine | 0 | 1.30 | — |
0.005 | 2.49 | 61.34 | |
0.25 | 5.24 | 111.93 | |
1.0 | 15.91 | 90.52 | |
FBS | 0 | 1.49 | — |
0.005 | 3.64 | 112.37 | |
0.25 | 7.08 | 159.66 | |
1.0 | 14.65 | 81.72 | |
Sweat | 0 | 1.49 | — |
0.005 | 2.90 | 72.68 | |
0.25 | 5.89 | 125.00 | |
1.0 | 14.96 | 83.46 | |
PBS | 0 | 1.46 | — |
0.005 | 3.40 | — | |
0.25 | 4.98 | — | |
1.0 | 17.60 | — |
In this work, the displayed sensor has good anti-interference performance, enabling it to accurately measure glucose concentration in complex samples. In order to further verify the reliability of the sensor in actual sample detection, commercial blood glucose meters and LIG/PEDOT/Au/GOx (BSA) electrodes were used to detect actual samples containing 1 mM glucose. As shown in Fig. 10B, the results showed that for actual samples containing interfering components, commercial blood glucose meters showed significant errors, while glucose sensing electrodes showed more stable and reliable measurement values. In all, due to the low cost and portability, blood glucose meters, as commercial products, have developed maturely and are widely accepted by consumers. Compared with commercial conventional glucose meter, apart from high sensitivity and wide detection range, the present sensor has higher precision and reusability. Both of them can accurately detect complex matrix samples. But blood glucose meters are not friendly to sweat detection, which may be because the complex electrolytes in sweat interfere with glucose detection. These highlights of the present sensor, as a reliable tool for glucose monitoring, has good application prospect in challenging environments. More efforts should be made in device preparation and intelligent detection.
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Fig. 11 SenSit Smart U disk detection device connected to Android smartphone detects (A) FBS, (B) sweat and (C) urine containing different concentrations of glucose. |
This journal is © The Royal Society of Chemistry 2024 |