A Field-effect transistor (FET) biosensor is a sensing device that detect changes in the interfacial potential associated with the adsorption of target molecules. FET biosensor is expected to be a simple measurement method by detecting target biomarkers without any labeling process. Recently, aptamers, single-stranded DNA molecules with specificity and stability, have attracted attention as receptors of FET biosensor [1,2]. Immobilization of aptamer onto metal oxide insulator surface of FET biosensors is often achieved by combining self-assembled monolayer (SAM) with cross-linking agents. Generally, an immobilization method using cross-linking agent, glutaraldehyde, between amino-modified aptamer and aminopropylsilane (APS) monolayer is often employed [3]. However, there is concern that aptamers may be immobilized by non-terminal amino group, because amino groups are also present within the bases constituting DNA structure. Consequently, decrease in affinity of the aptamer to the target molecule by involving collapse of the higher-order structure and inhibition of conformational change. In this study, we attempted immobilization method on the FET gate surface utilizing cross-linking thiol groups of aptamers to self-assembled monolayer instead of that between amino groups. By immobilizing thiol-modified aptamers onto APS monolayer via N-succinimidyl 3-(2-pyridyldithio) propionate (SPDP), we demonstrated specific immobilization at the aptamer termini for increasing the number of effective aptamers for binding of cortisol, which is stress-related hormone, was selected as a target.

The SiO₂ gate insulator of the FET was treated by piranha solution (H₂O₂ : H₂SO₄ = 1:4) for introduction of hydroxyl groups. Then, the FET chip was immersed in toluene solvent including 1%(v/w) 3-aminopropyltriethoxysilane in an argon atmosphere (60ºC for 7 min.). The SPDP was reacted with amino group of APS monolayer. After that, dithiothreitol (DTT) was added to the FET gate surface to form thiol group by cleavage of the disulfide bond within the SPDP. 100 nM 3’-thiol modified cortisol aptamer was immobilized to the FET gate surface by forming disulfide bond. Finally, 1 μM 2-mercaptoethanol was reacted with residual thiol sites to prevent non-specific adsorption of contaminating proteins. The FET characteristics were measured before and after the addition of cortisol for obtaining the threshold voltage shift (ΔV _g).

First, we compared the ability of reducing regent for cleavage of disulfide bond within SPDP to form the thiol group on the FET gate surface. A 5,5'-dithiobis(2-nitrobenzoic acid) (DTNB), which reacts with thiol group, was added to SPDP-modified surface treated by DTT, ascorbic acid or l-cysteine ethyl ester (LCEE). Following that, the absorbance derived from 5-mercapto-2-nitrobenzoic acid, which was generated from DTNB depending on the amount of thiol groups, was measured. As a result, thiol groups derived from SPDP were formed on the FET gate surface upon treatment with DTT. On the other hand, no absorbance of 5-mercapto-2-nitrobenzoic acid was observed when ascorbic acid or LCEE was used as reducing regents. Additionally, thiol groups were increased with SPDP concentrations in the range from 0.5 to 25 mM by DTNB absorbance measurements. Subsequently, electrical responses to cortisol using aptamer-immobilized FET were measured upon varying SPDP concentrations, with maximum response obtained at 1 mM. It should be noted that the FET response was acquired by increase in charge within the Debye length due to the conformational changes of the aptamer upon cortisol binding [4]. Finally, the sensitivity of aptamer-immobilized FET biosensor, which was fabricated using SPDP, were compared with that obtained using glutaraldehyde as the cross-linking agent. As a result, the gate voltage shift by addition of 1 μM cortisol when SPDP was utilized to fabricate the aptamer-immobilized FET biosensor was greater than the case of glutaraldehyde (Figure 1). These results indicated that cortisol was efficiently captured by the increase in the amount of aptamer, which was immobilized by 3’-terminus via disulfide bond. Therefore, the utilization of cross-linker for amine-to-thiol conjugation to nucleic acid immobilization could be useful for improving the sensitivity of FET biosensors.

[1] H. Chen et al., Analyst, 141, 2335-2346 (2016).

[2] N. Nakatsuka et al., Science, 362, 319-324 (2018).

[3] S. K. Vashist et al., Chem. Rev., 114, 11083−11130 (2014).

[4] H. Hayashi et al., Electrochemistry, 89, (2), 134-137 (2021).



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Nucleic acid amplification methods are widely used for sensitive detection of biomolecules utilizing signals derived from amplified DNA strands^[1]. Signal amplification by ternary initiation complexes (SATIC) has been attracted attention as high specific RNA detection method, because rolling circle amplification (RCA) by φ29 DNA polymerase requires the formation of ternary initiation complexes with DNA primers, circular DNA templates, and target RNAs^[2]. In SATIC system, target RNA was detected via fluorescence derived from the complex of guanine quadruplex (G4) structure, which generated within the amplified DNA strands, and thioflavin T (ThT)^[3]. To further simplify RNA detection using the SATIC system, it was expected to be effective to combine it with a field-effect transistor (FET) biosensor. A FET biosensor detects target molecules by altering the interfacial potential through the adsorption of charged molecules, thereby enabling simple label-free detection. Previously, we demonstrated the RNA detection via the negative charges of amplified DNA strands using the SATIC system, which originated from the DNA primer-immobilized FET sensor surface^[4]. This approach will be required the immobilization of different DNA primers for the detection of various targets. As the immobilization conditions of the DNA primers needed to be optimized for detection of each target RNA, there was concern that the versatility of the FET biosensor would be reduced. Here, the versatility of the FET biosensor could be reinforced by the capture of nucleic acid strands amplified in the bulk by the sensing interface formed under the same conditions. For the realizing the above approach, the interaction between the G4 structure and ThT was considered to be effective in capturing amplified DNA strands. In this study, we attempted to detect target RNA by capture of the amplified DNA strands containing the G4 structures using ThT-immobilized FET biosensor.

ThT derivatives (ThT-OE11^[5]) were immobilized to aminopropylsilane-modified FET gate surface via cross-linking by glutaraldehyde (1 h). After that, residual aldehyde group were capped by the ethanolamine (1 h). Subsequently, amplified DNA strands were generated by mixing target RNAs, DNA primers, circular DNA templates, the four deoxynucleotide triphosphates (dNTPs), and φ29 DNA polymerase at 37°C. Following that, the SATIC reaction solution was added to the ThT-immobilized FET (30 min.). Finally, the gate voltage (V _g)-drain current (I _d) characteristics were measured before and after the addition of the SATIC reaction solution for calculating gate voltage shift (ΔV _g).

The FET response to 100 nM target RNA was measured when the SATIC reaction solution was added ThT-immobilized FET biosensor. As shown in Figure 1(a), ΔV _g in the positive direction was obtained, indicating that the electron concentration in the channel was decreased by electrostatic interaction from negative charges of the amplified DNA strands. To optimize ThT immobilization condition, we evaluated the relationship between concentrations of ThT-OE11 solution and the sensor responses to target RNA. As a result, maximum response obtained at 500 mM ThT-OE11 solution. Subsequently, the sensor signal to 1 base-mismatched RNA as a negative control was almost no response (Figure. 1(b)). This result suggested that the nucleic acid strands were not generated by SATIC system, because the ternary initiation complexes did not form with 1 base-mismatched RNA. Thus, the ThT-immobilized FET with SATIC system specifically detected target RNAs. Finally, the measurement of ΔV _g at different concentrations of target RNAs were conducted using ThT-immobilized FET to examine the quantitative detectability. As a result, ThT-immobilized FET with SATIC system showed a concentration dependence in the range of 1 pM to 100 nM. In addition, we plan to obtain additional data on sensitivity by controlling the reaction time of FET biosensor with SATIC system. From these results, the capture of nucleic acid strands containing G4 structure by ThT-immobilized FET could be effective for RNA detection.

[1] M. Falco et al., TrAC Trends in Anal. Chem., 2022, 148, 3, 116538

[2] H. Fujita et al., Anal. Chem., 2016, 88, 7137−7144

[3] J. Mohanty et al., J. Am. Chem. Soc., 2013, 135, 1, 367–376

[4] H. Hayashi et al., Talanta, 2023, 273, 125846

[5] M. Kuwahara et al., Molecules, 2020, 25, 4936



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Amperometric sensors with nucleic acid amplification are expected to be useful for sensitive RNA detection. Nucleic acid amplification reaction involves current changes associated with accelerated redox reactions. Among the nucleic acid amplification methods, we have attempted to apply the signal amplification by ternary initiation complexes (SATIC) system to amperometric sensors^1,2. Amperometric sensor with SATIC system detected target RNA by measuring the reduction of hydrogen peroxide by DNAzyme via ferrocene methyl alcohol (FMA). Although linear sweep voltammetry (LSV) was used in previous study, the current was included reactions other than reduction by DNAzyme, such as the charge of the electrical double layer and mass diffusion of redox species. Pulsed electrochemical measurement removes the other current by refreshing the diffusion layer of redox species with each pulse voltage. Therefore, in this study, we focused on pulsed electrochemical measurement methods to further increase the signal-to-noise ratio of amperometric sensors with SATIC system by removing the noise. Current values derived from nucleic acid amplification with the SATIC system were compared by LSV, differential pulse voltammetry (DPV), normal pulse voltammetry (NPV) and square wave voltammetry (SWV.)

DNA primers and 6-merchapto-1-hexanol (MCH) were immobilized on gold electrode via gold-thiol bond. Then, the SATIC reaction proceeded (37 ºC, 2 h) on the electrode by target RNA, a circular DNA template, and φ29 DNA polymerase. After that, the electrode was placed in phosphate buffer containing hemin which iron-containing porphyrin. DNAzyme was formed by the G-quadruplex (G4) in the amplified nucleic acid by binding to hemin. The current value changes resulting from the catalytic reaction of DNAzyme before and after the SATIC reaction were measured by LSV, DPV, NPV, and SWV. The ranges of voltage scan were -0.15〜-0.4 V vs. Hg/Hg₂SO₄ and -0.05〜-0.4 V vs. Hg/Hg₂SO₄ in LSV and other pulsed electrochemical measurement, respectively. The step potential was 5 mV, pulse wide was 50 ms, and pulse periode was 100 ms. Additionally, the pulse amplitudes of DPV and SWV were 50 mV and 25 mV, respectively. The measuring points each were acquired after 40 s from applying the pulse signals.

First, each electrochemical measurement methods in a simple redox reaction were compared. The current density corresponding to the reduction reaction of FMA was measured using a polished bare gold electrode. The results showed an increase in current density depending on FMA concentration in all electrochemical measurement methods. Furthermore, the slope of the current value change relative to the FMA concentration was found to be greater in the order LSV &lt; SWV ≦ DPV ≦ NPV. As the pulse eliminated the contribution of the charging current due to the electrical double layer, the quantitative detectability of DPV, NPV and SWV was higher than LSV. Based on the above results, the signal-to-noise ratio of amperometric sensors with SATIC system was compared using each electrochemical measurement method. The amount of current density changes with the progression of the SATIC reaction to 1 µM Target RNA was measured. As a result, current density changes of LSV: 1.98 μA/cm², DPV: 12.2 μA/cm², NPV: 3.41 μA/cm² and SWV: 9.04 μA/cm² were observed (Figure 1). In NPV measurement, the variation in current density change was more varied than that using other electrochemical measurements. As the pulse signal were impressed from the open circuit potential in NPV, the current included not only electron transfer but also mass diffusion of FMA mediator, resulting in the variation of repeatability in each measurement. Moreover, in DPV and SWV measurements, the changing in current density were greater than that of LSV measurement. This result could be attributed to the fact that the contribution of the current due to the charging electrical double layer and mass diffusion of FMA mediator was removed by the impressed pulse signal. However, the signals obtained from measurements using DPV and SWV represent the slope of the current value change relative to the potential scan. Therefore, it is not advisable to compare each electrochemical measurement at a single RNA concentration point. In the presentation, we will further discuss the signal-to-noise ratio of measurement methods of amperometric sensors using the SATIC system by comparing the concentration dependence of the target RNA concentration obtained from each electrochemical measurement.

Reference:

[1] H.Fujita et al., Anal. Chem. 88, 7137−7144 (2016).

[2] H.Saze et al., 2023 ECSJ Fall Metting, Abstr, S12_2_04, [in Japanese]



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Currently, aqueous zinc batteries (AZBs) with zinc anode and aqueous electrolyte are focused on as one of the post-lithium-ion batteries because they can be a safe and inexpensive energy storage device. To date, a number of compounds, such as manganese oxide, vanadium oxide, Prussian blue analogues, and so on, have been investigated as cathode materials for AZBs. Although these materials are promising, suitable materials for cathode are still being studied for practical use. A variety of charge-discharge mechanisms are found in cathode of AZBs, which is influenced by many factors such as the composition, crystal structure, and morphology of the compounds used. Therefore, it is essential to understand the reaction mechanism to study new AZB systems.

Previously, we investigated the applicability of Li(Ni_0.33Mn_0.33Co_0.33)O₂ (NMC₁₁₁) as a cathode for AZBs¹. In the report, it is found for the first time that NMC₁₁₁ can be used as an AZB cathode by delithiation in 1 M KOH aq. (delithiation treatment). In order to implement the NMC cathode in AZB in the future, it is important to understand its delithiation treatment and charge-discharge behavior. During the delithiation process, it was found that water molecules inserted into the interlayer spaces and expanded the interlayer spacing, thereby facilitating the ion diffusion and redox activity². The purpose of this study was to reveal the charge-discharge behavior of NMC₁₁₁ cathode by characterization and electrochemical measurement, leading to the improvement of the performance of NMC₁₁₁ cathode in AZB.

The cathode was prepared by mixing NMC₁₁₁, acetylene black, Ketjen black, polyvinylidene difluoride (PVdF) (45 : 4 : 1 : 4 in weight ratio) diluted with an amount of N-methyl-2-pyrrolidinone (NMP) and coated it on a Ni porous substrate. A beaker cell was used for delithiation treatment with an NMC₁₁₁ cathode as the working electrode, a Cu foil as the counter electrode, a Hg/HgO electrode as the reference electrode, and 1 M KOH aq. as the electrolyte. Then a constant current of 170 mA g^-1 was applied to an NMC₁₁₁ cathode for 1 hour to deintercalate Li⁺ from NMC₁₁₁. Charge-discharge tests were performed using a beaker cell with a delithiated NMC₁₁₁ cathode, a Zn plate anode, a Hg/HgO reference electrode, and 1 M KOH (sat. ZnO) electrolyte, with a C rate of 0.1 C (= 17 mA g^-1) and cut-off potentials of -0.4－0.7 V vs. Hg/HgO. Then characterizations, X-ray Diffraction analysis (XRD) and Inductively Coupled Plasma Optical Emission Spectroscopy (ICP-OES), were performed on the NMC₁₁₁ cathode after charge-discharge tests. In addition, the WE used for Electrochemical Quartz Crystal Microbalance (EQCM) measurements was prepared by using a 6 MHz Au-coated quartz crystal (1.33 cm²) as a substrate, mixing NMC₁₁₁ and PVdF (99 : 1), diluting it with an appropriate amount of NMP, and loading about 10－50 μg onto the gold. Then, delithiation treatment and electrochemical measurement were performed in EQCM cell.

First, the crystal structures of NMC₁₁₁ cathodes after charge-discharge test were confirmed by ex-situ XRD. The results showed that there was a single peak attributed to the interlayer spacing of the cathode after charge, while the peak after discharges was split into two peaks (Figure 1). The interlayer spacing was calculated to be 0.702 nm after charging, it changed to 0.694 and 0.714 nm after discharging. These results suggests that the charge-discharge reaction of NMC₁₁₁ occurs by intercalation of ions. In addition, both an increase and a decrease in the interlayer spacing were observed during the discharge process, which indicated that ions could be inserted into the interlayer in stages. Next, the cathode composition was measured by ICP-OES after charge and discharge to identify the intercalants. The concentration changes of potassium, lithium, and zinc in the NMC₁₁₁ cathode after charge and discharge were small, less than 5 mAh g^-1 converted to specific capacity. Considering that the discharge capacity is about 120 mAh g^-1, the dominant intercalants are not Zn²⁺, K⁺, or Li⁺. Thus, insertion of OH^-, H⁺, and H₃O⁺ may have occurred since this battery system doesn't contain other ions. Finally, EQCM measurements could detect the weight of intercalants inserting into the cathode interlayer. The result of intercalant identification by EQCM measurements will be also discussed.

[1] H. Hayashi, et al., 63rd Battery Symposium in Japan, Abstr., 1H22 (2022). [in Japanese]

[2] H. Hayashi, et al., 91st ECSJ Annual Meeting, Abstr., S8-3_3_12 (2024). [in Japanese]

Figure 1 (a) Charge-Disharge curves of 1st cycle (b) XRD profiles at each point



Figure 1

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