A method for preparing an all-carbon organic electrochemical transistor based on laser-induced graphene and its application

Abstract

The present invention discloses a method for preparing an organic electrochemical transistor based on laser-induced graphene, and designs and manufactures an all-carbon flexible high-performance OECT based on LIG electrode. Due to the porous structure of LIG, the normalized transconductance of the obtained OECT is as high as 31S cm ‑1 . The high performance achieved is superior to that obtained in Au-based OECTs. The high performance can be maintained after the device shape is deformed. In addition, LIG-based OECTs have much higher cycling stability than Au-based OECTs. In addition, sensitive detection of H2O2 and glucose is demonstrated. By integrating LIG-based OECTs in microfluidic chips, wearable sweat sensors can also be obtained, which are expected to monitor sweat glucose in real time and noninvasively.

Description

Technical Field

[0001] The present invention relates to a biosensor, and in particular to a preparation method of an organic electrochemical transistor based on laser-induced graphene and applications thereof. Background Art

[0002] Since its invention in 1984, organic electrochemical transistors (OECTs) have attracted increasing attention in the fields of chemical sensing and bioelectronics. OECTs consist of source, drain, and gate electrodes, with the source and drain connected by an organic semiconductor, also known as a channel. During the sensing process, the analyte undergoes an electrochemical, catalytic, or enzymatic reaction on the gate surface, generating an electrical signal that is converted into an amplified drain current. Importantly, OECTs are compatible with solutions, making them particularly advantageous in chemical and biological sensing. In addition, OECTs can operate stably at low voltages (e.g., <1V). These advantages make OECTs a perfect choice for many sensing scenarios.

[0003] Nowadays, flexible and portable sensors are being widely studied to meet the growing and complex needs of modern lifestyles. In order to expand the applications of OECTs, it is necessary and urgent to develop alternatives with high flexibility. Traditionally, OECTs are fabricated on rigid surfaces and rigid gold electrodes are usually used. In order to increase flexibility, OECTs need to be fabricated on flexible surfaces and flexible electrodes need to be used instead of rigid gold electrodes. Carbon nanotubes, silver nanowires, graphene, and MXene have been used as flexible electrodes for the manufacture of flexible OECTs. However, the manufacturing process is complex and it is challenging to precisely control the size and channel width of OECTs. In addition, some of these electrodes are chemically and / or mechanically unstable, limiting their practical applications. Therefore, there is an urgent need for flexible OECTs with high performance and low cost. Summary of the Invention

[0004] In order to address the deficiencies of the prior art, the present invention aims to provide a method for preparing and applying a low-cost flexible all-carbon organic electrochemical transistor.

[0005] To address the above technical problems, the present invention proposes a technical solution: a method for preparing an organic electrochemical transistor based on laser-induced graphene, using dodecylbenzenesulfonic acid (DBSA) as a surfactant to effectively combine poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) with a porous LIG electrode; the resulting all-carbon flexible organic electrochemical transistor (LIG-OECT) achieves sensitive sensing of H2O2 and glucose by modifying the porous LIG gate surface with platinum nanoparticles (PtNPs) and glucose oxidase (Gox).

[0006] Preferably, the channel size of the specific LIG-OECT is 4 mm (w) × 500 μm (l) × 500 nm (d).

[0007] Preferred, high flexibility (bending, torsion), high stability (8000 cycles), high normalized transconductance 31cm S -1

[0008] Beneficial effects:

[0009] (1) All-carbon OECTs were fabricated by using LIG as electrodes. The entire fabrication process is very simple, inexpensive, and controllable, making it promising for large-scale production. Due to the high porosity and conductivity of the LIG electrode, a 31S cm-1 CMOS performance was achieved in LIG-based OECTs. -1 The OECTs were characterized by a high normalized transconductance. In addition, the employed LIG and PEDOT:PSS have high chemical stability, and the LIG-based OECTs have high flexibility. These properties are superior to those observed in OECTs based on conventional rigid gold electrodes. The LIG-based OECTs are highly sensitive to the detection of hydrogen peroxide (H2O2) and glucose. The OECT device can also be attached to human skin after integration with a microfluidic chip for real-time monitoring of sweat glucose. The present invention develops a simple and cost-effective method for the fabrication of high-performance OECTs, which may find promising applications in chemical and biological sensing.

[0010] (2) For DBSA-doped PEDOT:PSS channels, PEDOT 0 The strength increase and PEDOT + The intensity drop is significantly easier than that of the pristine PEDOT:PPS channel. Therefore, due to the doping of DBSA, the high-quality PEDOT:PSS–LIG interface may facilitate the ion doping and dedoping of the PEDOT:PPS channel under bias, which well explains the high performance of LIG-OCT.

[0011] (3) When the laser power increases from 2.6 W to 3.0 W, the normalized transconductance of the OECT increases from 12 S cm -1 Increased to 31Scm -1 ( Figure 4 e). When the laser scanning speed is increased from 100 mm S -1 Increased to 150mm S -1 When the normalized transconductance changes from 31S cm -1 Down to 10S cm -1 ( Figure 4 f).

[0012] (4) The present invention also investigated the stability of the LIG-based OECT by applying an alternating current (AC) potential to the gate electrode. The drain current alternately responded to the applied potential, and the maximum current remained stable for 8,000 cycles, demonstrating the high stability of the LIG-based OECT.

[0013] (5) The OECT device can achieve sensitive detection of H2O2 and glucose, with the minimum detection limit of H2O2 reaching 1 nM and the detection limit of glucose reaching 10 nM. It can also be used for real-time monitoring of sweat glucose, with the minimum glucose concentration in the sample being 20 nM and the maximum being 1 mM.

[0014] (6) All-carbon flexible OECT based on LIG electrode. Due to the porous structure of LIG, the normalized transconductance of the obtained OECT is as high as 31 S cm -1 The high performance achieved is superior to that obtained in Au-based OECTs. The high performance can be maintained after the device shape is deformed. After twisting at a large angle (>90°), the transfer curve of the device is stable, and only a slight decrease in the drain current is observed ( Figure 5 c) These results indicate that LIG-based OECTs are promising for many applications in flexible electronics.

[0015] Furthermore, the LIG-based OECTs exhibited significantly higher cycling stability than Au-based OECTs. Furthermore, sensitive detection of H₂O₂ and glucose was demonstrated. By integrating the LIG-based OECTs into a microfluidic chip, a wearable sweat sensor could be developed, potentially enabling real-time and noninvasive monitoring of sweat glucose.

[0016] (7) LIG-based OECTs and their application in sweat sensing have been demonstrated. In the near future, this technology can also be extended to the fabrication and application of many other flexible and wearable bioelectronic devices. More importantly, the entire device fabrication process can be further optimized to improve the output performance of the obtained electronic devices. A very cheap (3000 yuan) laser engraving machine with poor spatial resolution was used in this invention. By reducing the channel length in the OECT using a high-resolution laser engraving machine, the output performance of the device can be greatly improved. BRIEF DESCRIPTION OF THE DRAWINGS

[0017] Figure 1 Design of an organic electrochemical transistor (OECT) based on laser-induced graphene (LIG). (a) LIG on a polyimide substrate; (b) The porous structure of LIG; (c) Schematic diagram of a LIG-based OECT; (d) The chemical structure of the semiconducting polymer PEDOT:PSS; (e) The high current modulation capability of the LIG-OCT; (f) The principle of glucose detection using an OECT.

[0018] Figure 2 Preparation and characterization of LIG electrodes. (a) Schematic diagram of laser-induced graphene electrodes on flexible polyimide films; (b) Raman spectra of LIG electrodes prepared at laser powers ranging from 2.6 to 3.0 W; (c) XPS C1s spectra of polyimide and LIG electrodes; (d) Surface porosity of LIG electrodes prepared at laser powers ranging from 2.6 to 3.0 W; (e) SEM images and corresponding surface porosity of LIG electrodes prepared at laser powers ranging from 2.6 to 3.0 W; (f) Laser scanning speeds of 100, 125, and 150 mm s -1 SEM cross-sectional images of the LIG electrode prepared at different laser scanning speeds; (g) film thickness and (h) resistance of the LIG electrode prepared at different laser scanning speeds.

[0019] Figure 3 The working mechanism of high-performance LIG-OCT. (a) SEM images of pristine and DBSA-doped PEDOT:PSS films on LIG; (b) Schematic diagram of the contribution of LIG electrodes in high-performance OECTs; (c) Schematic diagram of the UV-vis-NIR test setup; (d) UV-vis-NIR spectroelectrochemical spectra of pristine and (e) DBSA-doped PEDOT:PSS films at biases of -0.2 to 1 V; (f) UV-vis-NIR spectra of polyimide film; (g) 0 and (h)PEDOT + The intensity changes of the peaks at different bias voltages.

[0020] Figure 4 Performance test of LIG-based OECTs. (a) Preparation process of LIG-based OECTs; (b) output curve and (c) transfer curve of LIG-ECTs; (d) Maximum transconductance of OECTs with channel width / length (W / L) of 4000-1000 / 500 μm; (ef) Dependence of the normalized transconductance of LIG-ECTs on the surface porosity and thickness of the LIG electrode; (g) 0.7 V (V D =0.1V) pulsed gate transient response of the OECT.

[0021] Figure 5 Flexibility and stability test of LIG-based OECTs. (a) Photograph of flexible LIG-ECT; (b) Transfer curve of LIG-ECT in a bent state; (c) Transfer curve of LIG-ECT before and after twisting; (de) LIG-OCT cycle test after 8000 cycles. G =0-0.7V, V D =0.1V.

[0022] Figure 6Contribution of the LIG electrode. (a) Transfer curve and normalized transconductance of an OECT based on an Au electrode; (b) Transfer curve of an OECT based on LIG and Au source and drain electrodes; (cd) CV curves of a PEDOT:PSS film coated on LIG and Au electrodes during 200 cycles; (e) Normalized transconductance of an OECT based on Ag / AgCl, LIG, Pt, carbon paper, and glassy carbon gates (LIG as source and drain); (f) Transfer curve and normalized transconductance of an OECT based on an Ag / AgCl gate (LIG as source and drain).

[0023] Figure 7 Sensitive detection of H2O2 and glucose based on LIG-ECT. (a) Schematic diagram of LIG-ECT as a glucose sensor; (b) Working principle of LIG-OCT for glucose detection; (c) Sensitive detection of H2O2; (d) ΔI in LIG-ECT D Response to the logarithm of H2O2 concentration; (e) Sensitive detection of glucose; (f) ΔI in LIG-ECT D Response to the logarithm of glucose concentration.

[0024] Figure 8 A wearable sweat glucose sensor based on LIG-ECT. (a) Schematic diagrams of the integration of a microfluidic chip and a LIG-OCT wearable sweat sensor; (c) Photograph of the wearable sweat glucose sensor on a volunteer's arm; (d) Glucose concentrations in 10 samples. DETAILED DESCRIPTION

[0025] The present invention will be described in detail below with reference to the accompanying drawings and specific embodiments.

[0026] Example 1

[0027] (1) Preparation of laser-induced graphene (LIG) electrodes: Graphene electrodes were in situ laser engraved on a 100 μm thick polyimide film using a CO2 laser (wavelength 10.6 μm) with a step size of 1 pixel. The laser operating mode was grating mode, the power was 3 W, the scanning speed was 100 mm / s, and the focal length was 5 cm. The electrodes were three coplanar electrodes (source, drain, and gate). The source and drain were parallel and the distance between them was 500 μm, which was defined as the channel width of the transistor. The gate and source were 2 mm apart, and all three electrodes were rectangular (3×12 mm). 2 (2) Organic semiconductor formulation: 0.5 μL of DBSA (dodecylbenzenesulfonic acid) and 500 μL of PEDOT:PSS (1.0-1.3 wt% aqueous solution) were mixed and stirred vigorously for 20 minutes. Then, 2.5 μL of GOPS (silane coupling agent) solution was added and stirred for another 2 hours.

[0028] (3) Preparation of organic electrochemical transistor (OECT) device: 15 μL of the mixture solution prepared in step 2 was drop-casted on the LIG source-drain electrode pattern on the polyimide film, which served as the channel layer in the OECT. A prefabricated PET frame (prepared by a laser engraving machine) was used to control the drop-casting area to 5×6 mm. 2 The drop-cast channel was annealed at 120°C for 1 hour. The organic semiconductor channel dimensions were 4 mm (w) × 500 μm (l) × 500 nm (d). 100 μL of phosphate-buffered saline (PBS) was then added to the semiconductor as the electrolyte layer for electrochemical transistor testing.

[0029] Organic electrochemical transistor (OECT) performance test: Transfer curve: drain voltage is -0.1V, gate voltage is swept from 0V to 2V. Output curve: drain voltage is from -1V to 0V, gate voltage is from 0V to 2V, with intervals of 0.2V.

[0030] H2O2 detection: Use the chronopotentiometry mode with a drain voltage of 0.2 V and a gate voltage of 1 V. Glucose detection: Use the chronopotentiometry mode with a drain voltage of 0.2 V and a gate voltage of 0.5 V.

[0031] Example 2

[0032] Surface modification of the gate of the organic electrochemical transistor: First, the gate is close to the source end (3×4mm 2 A layer of platinum nanoparticles (2.5 mA, 40 s) was sputtered with ions in the area above the platinum nanoparticles (Fig. 2) to specifically recognize H2O2. Subsequently, 10 μL of glucose oxidase solution (20 mg / mL) and 5 μL of chitosan solution (1 wt%) were added dropwise above the platinum nanoparticle area and refrigerated at 4°C for 12 hours.

[0033] Figure 1 Design of organic electrochemical transistors (OECTs) based on laser-induced graphene (LIG).

[0034] Our all-carbon OECTs are fabricated on a flexible polyimide film using LIG as electrodes and poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) as channel materials. First, a CO2 laser was used to prepare LIG electrodes on the surface of the polyimide film ( Figure 1 a), whose movement is controlled by a computer. The electrodes can be easily and precisely constructed without a shadow mask, and the source-drain distance can be precisely controlled. The surface morphology of the LIG electrode can be easily controlled by changing the fabrication parameters ( Figure 1 b). Then, an organic semiconductor layer (e.g., PEDOT:PSS) can be prepared as a channel layer using a drop-casting method ( Figure 1 Finally, a flexible LIG-based OECT was obtained ( Figure 1 c). Note that all three electrodes in our OECT are made of LIG, a cost-effective, chemically stable, and flexible carbon material, and the channel layer is made of a carbon-based polymer. Therefore, our device is called an all-carbon OECT. Our all-carbon OECT can be easily bent and twisted. In addition, due to the porous structure of the LIG electrodes, LIG-based OECTs may have high output performance ( Figure 1 e), making it promising in sensing applications ( Figure 1 f).

[0035] Figure 2 Preparation and characterization of LIG electrodes.

[0036] LIG electrodes were patterned on polyimide films using a CO2 laser engraver operating in raster mode ( Figure 2 a) Electrode patterns can be easily designed using a computer and fabricated with ease and precision. Laser engraving carbonizes the polyimide due to the ultra-high local temperature, forming a graphene pattern that, due to its high conductivity, can serve as an electrode. Clearly, the laser power and speed of the laser scanning significantly influence the performance of the resulting LIG electrode.

[0037] First, we found that the laser power has a significant effect on the performance of the obtained LIG electrode. Raman spectroscopy was used to characterize the properties of the obtained LIG electrode. The D, G, and 2D bands of graphene were clearly observed in the Raman spectrum, proving the successful preparation of graphene ( Figure 2 b). At a low laser power of 2.6 W, the D band is strong and the 2D band is weak, indicating that the defect density of the obtained LIG is high. Increasing the laser power leads to a decrease in the D band intensity and an increase in the 2D band intensity. When the laser power is increased from 2.6 W to 3.0 W, the 2D / D ratio increases from 0.2 to 2, indicating a significant decrease in the defect density. The chemical properties of a typical LIG electrode were then characterized using X-ray photoelectron spectroscopy (XPS). Compared with polyimide, the O1s and N1s XPS peak intensities of LIG decrease significantly, which is due to the carbonization of polyimide, during which N and O are released in the form of gas. The content of O and N in LIG decreases with increasing laser power. In particular, almost no N is observed in LIG prepared at 3.0 W, indicating a high degree of carbonization. By deconvolution, the C1s XPS spectrum of LIG can be seen as three peaks at 283.98, 284.78, and 288.48 eV ( Figure 2c). These three peaks represent the vibrations of C–C, C–O, and O–C=O, respectively. As the laser power for LIG preparation increases from 2.6, 2.8, to 3.0 W, the intensities of the C–O and C=O peaks decrease significantly, indicating better carbonization at higher laser powers. These XPS results are consistent with the Raman results. More interestingly, the LIG electrode has a high density of microscopic pores, which were characterized by SEM ( Figure 2 e). In all LIG electrodes prepared at laser powers between 2.6 and 3.0 W, the size of these pores in the LIG was approximately 1–10 μm. These pores are formed due to carbonization and gas evaporation during the laser engraving process. The content of these pores on the LIG surface is highly correlated with the applied laser power. As the laser power increases from 2.5 W to 3.0 W, the pore content of the LIG increases from 11.4% to 86.1% ( Figure 2 de). The increase in pore content will lead to an increase in specific surface area and interfacial capacitance, both of which may affect the output performance of LIG-based OECTs.

[0038] In addition, the laser scanning speed also affects the performance of the obtained LIG electrode. -1 The top-view SEM images of the LIG electrodes prepared at 175 mm s are similar, which may be due to the similar local temperature at a fixed laser power. In contrast, a higher laser scanning speed leads to a narrower thickness of the LIG electrode. -1 At high laser scanning speeds of 150 mm s, the prepared LIG electrodes are thin and have uneven thickness. -1 At the laser scanning speed of 100 mm / s, the thickness of the prepared LIG electrode is 26 μm. When the laser scanning speed is reduced to 125 and 100 mm / s -1 When the LIG thickness increases to 38 and 56 μm ( Figure 2 fg). At 75mm s -1 At low laser scanning speeds of 100 mm / s, the LIG electrode becomes too thick and easily falls off from the underlying substrate. In addition, the laser scanning speed also affects the resistance of the obtained LIG electrode. As the laser scanning speed increases from 100 mm / s -1 Increased to 150 mm s -1 , the resistance increases from 59Ω to 110Ω( Figure 2 h). Therefore, the electrical properties of the obtained LIG electrode can be optimized by changing the speed of the laser scanning.

[0039] Figure 3 The working mechanism of high-performance LIG-OCT.

[0040] Since the LIG surface is highly porous, the quality of the coated PEDOT:PSS channels may be greatly affected. We added dodecylbenzenesulfonic acid (DBSA) surfactant to PEDOT:PSS to make it fully contact with the porous carbon electrode surface (reducing the interfacial energy). When not doped with DBSA, PEDOT:PSS cannot be evenly coated on the porous LIG ( Figure 3 a), so the output curve of the corresponding OECT becomes highly unstable. PEDOT:PSS cannot fully contact the porous surface of the LIG electrode, resulting in parasitic capacitance. In contrast, despite the porous structure, DBSA-doped PEDOT:PSS is uniformly coated on the LIG surface ( Figure 3 b) The introduction of DBSA reduces the interfacial energy, thereby ensuring uniform coating of PEDOT:PSS on LIG. This improved surface coating contributes to the high performance achieved in LIG OECTs. When a gate voltage is applied, the interface between the porous gate and the electrolyte can enrich the surface charge to form a strong local electric field, allowing PEDOT:PSS to be more efficiently ion-doped, which well explains the high performance achieved in LIG-OCT. In addition, as a surfactant, DBSA can make the PEDOT:PSS film more hydrophilic, which may facilitate ion transport at the electrolyte / PEDOT:PSS interface.

[0041] The porous structure of PEDOT:PSS@LIG may greatly affect the ion doping / dedoping in the PEDOT:PSS channel under bias. To verify this, the ultraviolet-visible-near-infrared (UV-vis-NIR) spectra of the PEDOT:PSS channel connected by the LIG electrode were studied at different bias voltages. Figure 3 The strong light absorption at <500 nm is attributed to the absorption of the polyimide film ( Figure 3 f). It is reported that the absorption peak at about 650nm is attributed to the neutral state of PEDOT. + and bipolar PEDOT 2+ Broad absorption peaks are formed at ∼1000 nm and >1200 nm, respectively. When 0.1 vol% DBSA is doped in the PEDOT:PSS channel, a much faster increase in the intensity of the ∼650 nm peak is observed when the bias voltage increases from 0 to 0.7 V ( Figure 3 e), indicating rapid dedoping of ionized PEDOT in the channel. This rapid dedoping is likely due to the high-quality PEDOT:PSS–LIG interface and the large surface area of ​​PEDOT / PSS coated on LIG. Consequently, the conductivity of the channel can be effectively reduced at low bias voltages. In contrast, in the pristine PEDOT:PSS channel connected to the LIG electrode, the ionized PEDOT shows almost no neutralization process when the bias is increased from -0.2 to 1 V ( Figure 3 d), which may be due to the poor coating of the original PEDOT / PSS on LIG. Then, the PEDOT 0 and PEDOT + The intensity of the peak changes to facilitate comparison ( Figure 3 gh). For DBSA-doped PEDOT:PSS channels, PEDOT 0 The strength increase and PEDOT + The intensity drop is significantly easier than that of the original PEDOT:PPS channel. Therefore, the high-quality PEDOT:PSS–LIG interface due to DBSA doping may be beneficial for the ion doping and dedoping of the PEDOT:PPS channel under bias, which well explains the high performance of LIG-OCT ( Figure 3 e).

[0042] Figure 4 Performance test of LIG-based OECTs.

[0043] Flexible OECTs were fabricated on polyimide substrates using LIG as electrodes. 15 μL of the organic semiconductor mixed solution described in Example 1 was added to the substrate and then thermally annealed at 120°C to prepare the channel layer ( Figure 4 a). Typically, the channel width and length are controlled to 4 mm and 500 μm, respectively. The channel thickness is controlled to 500 nm. Note that the materials used are cost-effective, and the fabrication process is very simple and convenient. Furthermore, device fabrication can be easily scaled for large-scale production. Therefore, our strategy can be used for low-cost and large-scale fabrication of OECTs.

[0044] The obtained all-carbon OECTs showed promising output performance. In the output curves, the drain current increased with increasing source-drain voltage, and the total current decreased with increasing gate voltage ( Figure 4 b). These results indicate the depletion mode of the PEDOT:PSS-based OECT, which is consistent with the results of the transfer curves ( Figure 4 c). Notably, the transconductance of the all-carbon OECT normalized to the channel geometry was determined to be 31 S cm -1 Furthermore, the transconductance of our OECTs is linearly related to the W / L ratio of the device ( Figure 4 d), which is consistent with the results in the literature. Compared with OECTs in the literature, especially OECTs based on non-metallic electrodes, the performance of our device in this work is relatively high.

[0045] The characteristics of LIG greatly affect the output performance of the obtained OECTs, in which the transconductance is systematically compared. As shown in the figure, the normalized transconductance of the OECT increases from 12S cm to 12S cm when the laser power increases from 2.6W to 3.0W. -1 Increased to 31Scm -1 ( Figure 4 e). When the laser scanning speed is increased from 100 mm S -1 Increased to 150mm S -1 When the normalized transconductance increases from 31Scm -1 Down to 10S cm -1 ( Figure 4 f). Interestingly, the increase in transconductance is closely related to the increase in pore density and LIG thickness. These results indicate that the OECT transconductance is proportional to the pore content and thickness of the LIG electrode. In addition, an ultrashort response time of 7 milliseconds was observed in LIG-based OECTs ( Figure 4 g). The high stability and short response time make LIG-based OECTs promising for many bioelectric applications.

[0046] Figure 5 Testing the flexibility and stability of LIG-based OECTs.

[0047] Our LIG-based OECTs are highly flexible. The LIG-based OECTs on polyimide substrates can be easily bent using two fingers ( Figure 5 a). In addition, the output performance of the LIG-based OECT remains stable under high shape deformation. When the device is bent in the positive or negative direction at different angles, the transfer curves are similar to those in the unbent OECT ( Figure 5 b), indicating that the LIG-based OECT has high flexibility and stability. In addition, after twisting at a large angle (>90°), the transfer curve of the device is stable, and only a slight decrease in the drain current is observed ( Figure 5 c) These results indicate that LIG-based OECTs are promising for many applications in flexible electronics.

[0048] We also investigated the stability of the LIG-based OECTs by applying an alternating current (AC) potential on the gate electrode. The drain current alternately responded to the applied potential, and the maximum current remained stable for up to 8,000 cycles ( Figure 5 de), indicating the high stability of LIG-based OECTs.

[0049] Figure 6 The contribution of LIG electrode.

[0050] We compared the performance of LIG-based OECTs with conventional OECTs based on gold electrodes. The gold electrodes were prepared using ion sputtering (glass substrate, sputtering current 2.5 mA, time 270 s). The electrode pattern size was the same as that of the LIG electrode in Example 1, with a coplanar three-electrode structure. The Au-based OECT of the same size showed a scattering of only 23 S cm -1 The normalized transconductance ( Figure 6 a), which is significantly smaller than that of LIG-based OECTs (31S cm -1 ). An OECT can be viewed as a series circuit consisting of a source electrode, a semiconductor, and a drain electrode. The source and drain electrodes are considered resistors with constant resistance values. Higher electrode resistance should reduce the ability of the gate to regulate the drain current. In our case, the resistance of LIG is 60Ω, which is an order of magnitude higher than that of the gold electrode ( Figure 6 b). However, higher transconductance was observed in LIG-based OECTs. Therefore, the porosity of LIG should contribute to the enhanced transconductance in LIG-based OECTs. In the subsequent cyclic voltammetry electrochemical stability comparison ( Figure 6 cd) and found that the capacitance of PEDOT:PSS on the LIG electrode was 10 mF and that on the Au electrode was 5 mF. Due to the high capacitance, the PEDOT:PPS channel can be more efficiently doped on LIG, which explains the high transconductance obtained in LIG-OCT.

[0051] In addition to transconductance, the stability of LIG-based OECTs is also higher than that of Au-based OECTs. After 200 cyclic voltammetry (CV) cycles in phosphate-buffered saline (PBS), the capacitance of PEDOT:PSS-coated LIG retained 102% of its initial value (from 9.6 mF to 9.8 mF) ( Figure 6 c). In addition, no morphological changes were observed after 200 CV cycles. In contrast, the capacitance of the PEDOT:PSS-coated Au electrode could only maintain 27% of its initial value (from 5.2 mF to 1.4 mF) ( Figure 6 d), showing worse stability. Optical images show that many protrusions and wrinkles appeared after 200 CV cycles, indicating poor morphological stability, which may be attributed to the mismatched swelling between PEDOT:PSS and the Au electrode in solution. The porous structure of LIG increases the interfacial area between PEDOT:PSS and LIG, reducing the deformation caused by membrane expansion in solution. Therefore, the LIG-based OECT has high stability during chemical detection in solution. On the other hand, the capacitance of PEDOT:PSS on the LIG electrode is about twice that on the Au electrode at the beginning and about 7 times after 200 operation cycles. The larger capacitance means that ions in the electrolyte can dope PEDOT:PSS more effectively, resulting in high OECT performance.

[0052] On the other hand, the output performance of OECTs is often closely related to the choice of gate electrode, including non-polar Ag / AgCl electrodes and polar Pt, LIG, and other carbon electrodes. We compared the transfer curves of OECTs based on Ag / AgCl, LIG, Pt, glassy carbon, and carbon paper gates ( Figure 6 e) The normalized transconductance of the LIG-gate-based OECT is 31 S cm -1 , which are similar to those based on polarizable gates and lower than those based on non-polarizable Ag / AgCl gates (49 S cm -1 ). The on-off ratio is another important parameter of OECTs. OECTs with non-polarizable Ag / AgCl gates show the highest on-off ratio. However, Ag / AgCl is chemically unstable and incompatible with enzymes. Compared with polarizable gates of Pt, carbon paper, and glassy carbon, the LIG gate leads to a relatively high on-off ratio, especially at high scan speeds, which may be due to the presence of a porous structure on the LIG surface. In addition, due to the larger specific surface area, more enzymes can be attached to the LIG gate, which is beneficial for chemical and biological detection.

[0053] Figure 7 Sensitive detection of H2O2 and glucose based on LIG-ECT.

[0054] Our LIG-based OECTs can be used for chemical and biological detection. Hydrogen peroxide (H2O2) is an intermediate product for detecting various metabolites (such as glucose, lactose, and lactic acid). As described in Example 2, the LIG gate electrode ( Figure 7 a). Electrochemical oxidation of H2O2 on the gate surface under positive bias catalyzed by PtNPs (Reaction 1).

[0055]

[0056] This oxidation reaction generates electrons which are then transferred to the gate electrode, resulting in an effective gate voltage (V G-Eff )Increase( Figure 7 b). As a result, the drain current changes, achieving sensitive detection of H2O2. When H2O2 is detected, the drain current of the OECT drops sharply. The drain current decreases (ΔI D ) is closely related to the concentration of the detected H2O2 solution. Quantitatively, ΔI D An exponential fit was obtained in the range of 1 nM to 1 mM, with a coefficient R 2 >0.99( Figure 7 d) Even though the response is not linear, the lowest detection limit can reach 1 nM.

[0057] Furthermore, our OECTs can quantitatively detect glucose. Glucose levels in sweat and blood are important indicators of health and diabetes. For glucose detection, as described in Example 2, the gate electrode was modified with chitosan mixed with glucose oxidase (GOx).

[0058] During detection, glucose is electrochemically oxidized under the catalysis of GOx on the gate surface (Reaction 2).

[0059]

[0060] In this reaction, H2O2 is produced, and the generated H2O2 can be detected by our OECT to achieve quantitative detection of glucose. As shown in the figure, glucose can be quantitatively detected in the range of 10nM-1mM, ΔI D An exponential fit was obtained over the 10 nM to 1 mM range, with an R-squared coefficient > 0.99. More importantly, a linear response was achieved over the concentration range of 10 nM to 10 μM. Note that the drain current was 3 times larger than the gate current. This signal enhancement is a unique advantage of OECTs. The drain current here was 17 times larger than that of Au-based OECTs, demonstrating the superiority of LIG-OECTs. These results demonstrate that our flexible and cost-effective LIG-OCT can be used for quantitative glucose detection over a wide concentration range, with a detection limit of 10 nM.

[0061] Figure 8 Wearable sweat glucose sensor based on LIG-ECT.

[0062] The glucose level in sweat is closely related to the glucose concentration in blood. Therefore, non-invasive detection can be achieved by detecting glucose in sweat, and wearable sweat sensors can be used to achieve continuous monitoring of health status ( Figure 8 a). Here, LIG-OCT is integrated with a microfluidic chip for efficient collection and detection of sweat. The microfluidic chip consists of six layers ( Figure 8 b). The bottom layer is the skin adhesive layer, which has 5.3cm 2 The sweat collection area is covered by the bottom layer, OECT detection layer, transfer layer, sweat retention layer, and top sweat outlet layer. The microfluidic channels are made using a laser engraving machine, which helps to create precise patterns in a short time. The cross-sectional area of ​​the microfluidic channels is only 0.04mm 2. Due to capillary forces, the sweat produced can be collected by the sweat collection area and then flow into the OECT detection chamber for quantitative glucose sensing. The excess sweat then flows into the sweat retention layer (sweat reservoir) and finally flows out through the outlet of the top layer. Red ink was first used to simulate the flow of sweat within the microfluidic chip. The red ink filled the microfluidic channels, OECT detection chambers and sweat collection areas. This process can be completed in just 10 seconds, and only 20μL of ink is enough to fill the entire microfluidic chip. Therefore, our sweat sensor can be used to effectively collect and sense sweat on human skin.

[0063] The wearable LIG-ECT sweat sensor can be easily attached to different body parts such as the arm ( Figure 8 c) for collecting and detecting sweat glucose. The integrated sweat sensor demonstrated excellent linearity when used to detect artificial sweat. Next, the integrated sweat sensor was used to detect sweat samples. Nine sweat samples were successfully tested. The glucose concentrations in samples 2 and 5 corresponded to normal human sweat glucose concentrations, sample 9 to that of a person experiencing hypoglycemia, and the remaining samples to those experiencing hyperglycemia. The lowest glucose concentration in the sample ranged from 20 nM to 1 mM, with an accuracy rate of 94.5%.

Claims

1. A method for preparing an organic electrochemical transistor based on laser-induced graphene, characterized in that : Flexible and easy-to-pattern laser-induced graphene LIG is used as an electrode, including a source electrode, a drain electrode and a gate electrode; a CO2 laser engraver working in a grating mode patterns the LIG electrode on a polyimide film, and the source electrode and the drain electrode are connected by an organic semiconductor, and the organic semiconductor uses dodecylbenzenesulfonic acid as a surfactant to effectively combine poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) with the porous LIG electrode; the fabricated all-carbon flexible organic electrochemical transistor realizes sensitive sensing of H2O2 by modifying platinum nanoparticles on the surface of the porous LIG gate, and sensitive sensing of glucose by modifying platinum nanoparticles, glucose oxidase and chitosan on the surface of the gate; The engraving parameters of laser-induced graphene were CO2 laser, wavelength 10.6 μm, focal length 5 cm, power 3.0 W, scanning speed 100 mm / s, step size 1 pixel; The formula of organic semiconductor is 500µL poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) with a concentration of 1.0~1.3wt% aqueous solution, 0.5µL dodecylbenzenesulfonic acid, 2.5µL silane coupling agent, 15µL of the mixed solution is drop-casted on the electrode area as the channel, and annealed at 120℃ for 1 hour; The dimensions of the organic semiconductor channel w×l×d are 4 mm×500 μm×500 nm; Preparation of laser-induced graphene LIG electrodes: In-situ laser engraving of graphene electrodes on a 100 μm thick polyimide film was performed using a CO2 laser. The laser operating mode was grating mode, the power was 3 W, the scanning speed was 100 mm / s, and the focal length was 5 cm. The electrodes were three coplanar electrodes, the source and drain were parallel, and the distance between them was 500 μm, which was defined as the channel width of the transistor. The distance between the gate and the source is 2 mm, and the size of the three electrodes is 3×12 mm. 2 Rectangle; Organic semiconductor ratio: 0.5µL dodecylbenzenesulfonic acid and 500µL 1.0~1.3wt% poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) aqueous solution were mixed and stirred vigorously for 20 minutes; Then, 2.5 μL of silane coupling agent solution was added and stirred for another 2 h; Preparation of organic electrochemical transistor devices: 15 μL of the mixture solution was drop-coated on the LIG source-drain electrode pattern on the polyimide film, which was used as the channel layer in the organic electrochemical transistor; the drop-casting area was controlled to 5×6 mm using a prefabricated PET frame 2 ; The drip channel was annealed at 120°C for 1 hour; 100 μL of phosphate buffered saline was then dripped onto the semiconductor as the electrolyte layer for the electrochemical transistor test; Surface modification of the gate of an organic electrochemical transistor: First, a 3×4 mm 2 A layer of platinum nanoparticles was ion sputtered on the area with a sputtering current of 2.5 mA and a time of 40 s for specific identification of H2O2; then 10 µL of a 20 mg / mL glucose oxidase solution and 5 µL of a 1 wt% chitosan solution were dripped on the platinum nanoparticle area.

2. The method for preparing an organic electrochemical transistor based on laser-induced graphene according to claim 1, characterized in that: The prepared organic electrochemical transistor has high flexibility, high stability, and a normalized transconductance of up to 31 cm S -1 .

3. Application of the organic electrochemical transistor based on laser-induced graphene prepared by the method according to claim 1, characterized in that: PtNPs and GOx were modified on the LIG gate to achieve highly sensitive H2O2 and glucose detection.

4. The use of an organic electrochemical transistor based on laser-induced graphene prepared by the method according to claim 1, characterized in that: Used to prepare wearable sweat sensors.

Images