Non-faradaic Impedimetric Biosensing with Open Bipolar Electrode Platform

Abstract

We propose an impedimetric biosensor using the bipolar phenomenon with an open bipolar electrode (oBPE). The oBPE platform offers a structurally simple and highly flexible design. The unique electrochemical biosensor consists of a BPE, driving electrodes for detection, and an electrolytic cell as its components. We employed non-faradaic electrochemical impedance spectroscopy to detect biomolecular events occurring on the BPE surface. The biosensor showed a 45 % increase in impedance at 2 kHz after bovine serum albumin adsorption onto the BPE surface. Additionally, the significance of solution resistance, which was uniquely defined by the distance between driving electrodes and a BPE, was demonstrated through changes in impedance. We used the detection system as an immunosensor to detect the binding of C-reactive protein (CRP) to antibody-modified BPE, based on changes in impedance. The biosensor detected the binding of CRP to antibody without the need for redox-active reagents or complex signal amplification steps. The biosensor showed a linear response to the concentration of CRP in the range from 0.001 to 1 µg/mL, comparable to that of conventional three-electrode systems. These findings demonstrate the potential of non-faradaic impedimetric biosensors with an oBPE for sensitive detection of biomolecular events in biochemical analysis.

1. Introduction

The bipolar phenomenon induces electrochemical reactions on wirelessly connected conductors. This phenomenon enables the development of novel electrochemical systems that transcend the limitations of conventional electrochemistry, thereby facilitating the synthesis of advanced conductive materials and the fabrication of innovative electrochemical devices.^1,2 Bipolar electrochemistry systems are constructed with driving electrodes (DEs) and a bipolar electrode (BPE). The DEs form an electric field with a potential gradient across the BPE. The electric field drives electrochemical reactions at the BPE surface without requiring direct contact between the DEs and the BPE. We have previously demonstrated the fabrication of spatially defined conductive polymer films on BPEs through localized electrochemical polymerization.³ Furthermore, Inagi et al. have reported the synthesis of organic films with tailored gradient and pattern characteristics through precise control over the potential distribution along the BPE.^4,5

Bipolar electrochemistry-based biosensors, which use the bipolar phenomenon to detect analytes, are attracting attention as a novel sensing technique that leverages their unique characteristics.⁶ Common electrochemical biosensors consist of three electrodes: a working electrode, a reference electrode, and a counter electrode. Biosensors with a three-electrode electrochemical cell have been studied extensively due to their ability to detect trace amounts of analytes with high sensitivity and rapidity.^7,8 However, the biosensor with a three-electrode electrochemical cell is difficult to miniaturize and reduce costs because it has many restrictions on the configuration of electrochemical cells.^9–12 In contrast, biosensors using a bipolar phenomenon offer flexibility in electrochemical cell configurations due to their unique electrode arrangements and support designs, which are beneficial for miniaturization and cost reduction in achieving high-throughput analysis.

BPEs are classified into two types: closed-BPE (cBPE) and open-BPE (oBPE).¹³ The cBPE platform comprises separate electrochemical cells, enabling selective electrochemical reactions within each designated cell. This structure leads to higher current efficiency in cBPE platforms compared to oBPE platforms, which use a single electrochemical cell. cBPE-based electrochemiluminescence sensors which integrate cBPEs with electrochemiluminescence, are actively researched as promising tools that offer simple and highly sensitive detection methods.^14–17 Achieving precise control over electrochemical reactions is more challenging in oBPE platforms, as they operate within a single electrochemical cell. However, the simpler cell structure of oBPEs provides advantages in electrode design and miniaturization.

The inherent simplicity of oBPEs provides a versatile platform for the designing electrochemical sensors. However, the presence of multiple conductive pathways in oBPE platforms leads to reduced current efficiency compared to cBPE platforms. This limitation presents a significant challenge for their application in high-sensitivity biosensing. As a result, reported research efforts focused on biosensors with oBPE have been relatively limited. Most biosensors with oBPE use visual labels, such as electrochemiluminescence or colorimetric methods, and there are few reports on simple electrochemical biosensors that directly detect electrical changes on the oBPE surface resulting from biomolecular events.^18–20

Here, we describe a novel biosensor based on bipolar electrochemistry with an oBPE (Fig. 1). Our self-built biosensing system detects the adsorption and binding of biomolecules to the surface of oBPEs by measuring changes in impedance through non-faradaic electrochemical impedance spectroscopy (nEIS). In the oBPE setup, faradaic impedance spectroscopy, which employs redox sensitizers like ferricyanide and ferrocyanide, was found to be unsuitable due to competing electrochemical reactions occurring on the DEs. Thus, we opted to use nEIS in this study. nEIS does not provide the sensitization advantages associated with faradaic reactions, but it simplifies the reaction system.^21–23 The biosensor demonstrated adequate sensitivity for the electrochemical detection of C-reactive protein (CRP) through an antigen-antibody reaction.

Figure 1.

Schematic illustration of the impedance measurement based on bipolar phenomenon (A). Photographic image of the biosensor (B).

2. Experimental

2.1 Materials

Anti-human CRP monoclonal antibody (Ab) and human CRP were purchased from Oriental Fermentation Industry Co. (Tokyo, Japan). Bovine serum albumin (BSA) was purchased from Roche Diagnostics K.K. (Tokyo, Japan). 3-Mercaptopropionic acid (MPA) and N-hydroxysuccinimide (NHS) were purchased from Nacalai Tesque, Inc. (Kyoto, Japan). 1-(3-Dimethylaminopropyl)-3ethylcarbodiimide hydrochloride (EDC) was purchased from Tokyo Chemical Industry Co., Ltd. (Tokyo, Japan). Blocking peptide fragment (BP) was purchased from Toyobo Co., Ltd. (Tokyo, Japan). Other reagents used were special-grade or analytical reagents without further purification. Au-coated glass electrode (AuE, 100 mm × 100 mm, specific resistance 4 µΩ cm) was purchased from Geomatec Co., Ltd. (Kanagawa, Japan).

2.2 Preparation of biosensing system based on bipolar electrochemistry

The cross-section of an Au wire, with a diameter of 0.5 mm, was used as a DE. Two through holes, each with a diameter of 0.55 mm, were drilled 1.5 mm from the center of the bottom surface of a polytetrafluoroethylene (PTFE) rod with a diameter of 6.0 mm. The distance between the two holes was 3.0 mm. Au wire was threaded through these holes and trimmed to match the bottom surface of the PTFE rod. The assembly was then polished to eliminate any step between the bottom surface of the PTFE and the Au wire. The PTFE rod with the Au wire passing through it was connected to an acrylic resin pipe with an inner diameter of 6.0 mm and secured to the Z-axis stage via a dedicated jig. The Au wire was reinforced with a brass pipe with an inner diameter of 0.5 mm and connected to the electrochemical analyzer. The Z-axis stage was managed through specialized software, which controlled the distance between the DEs and a BPE. AuE was cut into 10 mm squares and used as a BPE. The BPE was sized to ensure it does not significantly affect the sensitivity of the biosensor. The AuE was subjected to vacuum plasma treatment for 30 s and washed with ethanol and pure water before use. A cylindrical glass container with a diameter of 70 mm and a height of 15 mm was used as the electrolytic cell. The cell was filled with the measurement solution during the measurement, and the BPE was placed on its inner bottom surface.

2.3 Electrode preparation

BSA-adsorbed BPE (BSA/AuE) was prepared by immersing AuE in 100 mM (M = mol L⁻¹) phosphate buffer (PB, pH 7.0) containing 1 mg/mL BSA solution for 1 h. After the adsorption process, the electrode was rinsed with pure water.

The preparation of the Ab-immobilized electrode and the binding of CRP to the electrode were conducted as follows: First, bare AuE was immersed in an aqueous solution of 10 mM MPA for 1.0 h to form a self-assembled monolayer of MPA on the AuE surface (MPA/AuE). After washing with pure water, MPA/AuE was immersed in an aqueous solution containing 0.2 M EDC and 0.2 M NHS for 30 min to activate the terminal carboxy groups on the MPA. After rinsing unbound molecules with pure water, the electrode was immersed in 100 mM PB (pH 7.0) containing 10 µg/mL Ab for 1 h to immobilize Ab covalently (Ab/MPA/AuE). Unbound Ab was removed by rinsing it with pure water. Blocking treatment to avoid non-specific adsorption of biomolecules was performed by immersing Ab/MPA/AuE in 100 mM PB (pH 7.0) containing 0.5 mg/mL BP for 30 min to achieve blocking (BP/Ab/MPA/AuE). Finally, CRP was bound to Ab on BPE by immersing BP/Ab/MPA/AuE in 100 mM PB (pH 7.0) containing 0–1.0 µg/mL CRP for 1 h at 37 °C (CRP/BP/Ab/MPA/AuE). BPEs were washed with pure water and stored in 100 mM PB at 4 °C until impedance measurements were taken.

2.4 Impedance measurement

nEIS was conducted using the electrochemical analyzer (IM6, ZAHNER-elektrik GmbH & Co. KG). During the measurement process, the electrolytic cell was filled with 30 mL of 100 mM PB (pH 7.0), and the BPE was placed on the inner bottom surface of the cell. The DEs were positioned directly above the BPE. Unless otherwise specified, the distance between DE and BPE (d_DE–BPE) was 0.3 mm. The DEs were linked to the electrochemical analyzer. The electrochemical analyzer provided an alternating voltage with an amplitude of 0.1 V, operating in the frequency range from 1 MHz to 1 Hz for impedance measurements.

3. Results and Discussion

3.1 Characterization of the biosensor with oBPE

To evaluate the fundamental characteristics of the biosensor, nEIS was performed both with and without AuE as the BPE. When the BPE was placed in the electric field formed by DEs, the impedance decreased by 17–42 % across the overall frequency range (Fig. 2A). The impedance consists solely of contributions from the DEs and the electrolyte between them in the absence of BPE, as shown in the equivalent circuit model in Fig. 2B. In contrast, inserting the BPE into the electric field establishes a new conductive path via the BPE in parallel to the existing conductive path between the DEs, which is formed solely through the electrolyte (Fig. 2C). The decrease in impedance thus shows that the presence of the BPE, which is not directly coupled to the DE, was reflected in the measured impedance value of this sensing system.

Figure 2.

Bode plots obtained in the absence and presence of BPE (A). Schematic illustration of measurement conditions and equivalent circuit models (B and C).

To assess the detection capability of the biosensor for biomolecular adsorption, impedance measurements of BPE with physically adsorbed BSA, a model protein, were conducted without an electrochemical probe (Fig. 3A). The impedance measured with BSA/AuE showed an increase across the frequency range of 1 Hz to 60 kHz compared to the impedance measured with the unmodified AuE (Fig. 3B). The maximum increase in impedance was approximately 45 % at 2 kHz (Fig. 3C). This change in impedance is attributed to the increased resistance of BPE (R_BPE) that resulted from the inhibition of current flow along the DE-BPE path, caused by the adsorption of BSA onto the surface of the BPE. Consequently, it has been demonstrated that the biosensor with an oBPE effectively detects the adsorption of biomolecules on the BPE surface through changes in non-faradaic impedance.

Figure 3.

Schematic illustration of AuE and BSA/AuE (A). Bode plots comparing AuE and BSA/AuE (B). The rate of change in impedance to BSA adsorption on AuE (C).

3.2 Effect of distance between DE and BPE for detection of biomolecules

In this biosensing system, the distance between DEs and a BPE (d_DE–BPE) influences the resistance of the conductive pathway through the BPE. The distance, which is a unique value of the biosensing system, defines the resistance derived from the electrolyte between the DEs and the BPE (R_{Sol(DE–BPE)}). Therefore, we investigated the relationship between d_DE–BPE and the detection of BSA, assuming that d_DE–BPE affects the sensitivity of the biosensor. As a result, a smaller d_DE–BPE yielded larger impedance changes when the distance ranged from 0.3 mm to 1.5 mm (Fig. 4). A 30 % impedance change was observed at a d_DE–BPE of 0.3 mm at 1 kHz. The resistance of the electrolyte solution between DE and BPE R_{Sol(DE–BPE)} decreases as d_DE–BPE becomes smaller. Therefore, it is considered that as d_DE–BPE decreases, the current flowing through the DE-BPE path increases, which in turn increases the sensitivity of the BPE to impedance changes.

Figure 4.

Influence of d_DE–BPE on the rate of change in impedance at 1 kHz before and after adsorption of BSA onto BPE.

In contrast, the change in impedance almost plateaued when d_DE–BPE became larger. The plateau of the change indicated that the contribution of impedance of BPE (Z_BPE) to the total impedance decreases as the R_{Sol(DE–BPE)} increases. The distance between DEs (d_DE–DE), serving as a conductive path between DEs, is maintained at a constant 3 mm in this study. Considering the distance, the results in Fig. 4 demonstrate the importance of the solution resistance value, as determined by d_DE–BPE and d_DE–DE (R_{Sol(DE–BPE)} and R_Sol(DE–DE)), in this measurement system. Additionally, when the d_DE–BPE value decreased, significant errors were observed in the impedance measurements. Therefore, we adopted a d_DE–BPE of 0.3 mm for the other experiments in this study, as this value produced a substantial impedance change and reasonable errors due to protein adsorption. The larger error at small d_DE–BPE is likely due to the significant difference in the state of the electrolyte solution from the bulk solution, as well as the device's positional accuracy limitations. The electrolyte solution was trapped in a narrow space between a BPE and the PTFE rod with DEs. Furthermore, the smaller the d_DE–BPE, the more strongly positional deviations affect the data. Therefore, we aim to enhance sensitivity further by investigating the arrangement and shape of DEs and a BPE in future studies.

3.3 Changes in impedance at each modification step of BPE for CRP detection

We used the binding reaction between CRP and Ab as a model immunoreaction to evaluate the performance of the biosensing system with an oBPE for immunosensing. CRP, used as an analyte, is an acute-phase protein that increases with inflammation, infection, and tissue injury. CRP is an important biomarker for diagnosing and monitoring inflammatory diseases, cardiovascular risks, and certain cancers.^24,25 Thus, a simple, cost-effective CRP detection method is needed for early diagnosis and high reliability.^26,27 To detect the binding of CRP to Ab immobilized BPE, we first monitored the change in impedance at each step of surface modification of the BPE. The surface modification steps were as follows (Fig. 5A): Firstly, an MPA monolayer was formed on bare AuE (MPA/AuE). Next, Ab was immobilized on the carboxy groups of MPA (Ab/MPA/AuE). A blocking treatment was conducted using BP (BP/Ab/MPA/AuE), and subsequently, CRP was bound to the Ab (CRP/BP/Ab/MPA/AuE). The Bode plots for each modification step showed an apparent increase in impedance in the frequency range of 10 to 100 Hz (Fig. 5B). This result indicates that the measurement system in this study is capable of detecting various modifications on the BPE and the binding of CRP to Ab-immobilized on BPE as a change in impedance. Figure 5C shows the difference in impedance values at 30 Hz measured at each modification step compared to that of unmodified AuE. The impedance at 30 Hz increased with each modification step, indicating that the deposition of modifier molecules enhanced the electrical insulation of BPE. CRP binding, induced by the antigen-antibody reaction, resulted in the most significant impedance increase (6 kΩ) in each step.

Figure 5.

Preparation scheme of BPE surface modification (A). Bode plots of BPEs at each modification step: AuE, MPA/AuE, Ab/MPA/AuE, BP/Ab/MPA/AuE, and CRP/BP/Ab/MPA/AuE (B). The inserted figure in (B) indicates Bode plots magnitude in the range from 10 Hz to 100 Hz. Impedance changes at 30 Hz of BPEs at each modification step for the impedance of AuE (C).

3.4 CRP quantification

To assess the quantitative performance of the biosensor, BP/Ab/MPA/AuE was treated with various concentrations of CRP solutions (0–1.0 µg/mL). The response to CRP binding was displayed as a rate of impedance change relative to untreated BPE (BP/Ab/MPA/AuE). The impedance response measured by using the BPEs treated with CRP increased with the increasing concentration of CRP (Fig. 6A). The impedance response at 30 Hz, which showed the most significant rate of impedance change, exhibited a linear relationship to the CRP concentration in the measured range (Fig. 6B). The linear response observed across a range of CRP concentration from 0.001 to 1.0 µg/mL is described by the following equation: ΔZ = 2.62 log C + 10.5 (R² = 0.997), where ΔZ represents the impedance response [%], and C is the CRP concentration [µg/mL]. Experimentally obtained limit of detection (LOD) was 0.002 µg/mL. The LOD was estimated using the equation: LOD = 3σ/b, where σ represents the standard deviation of the blank measurement, and b represents the slope of the regression line. R.N. Dalialla et al. performed the detection of CRP using faradaic impedance measurement with a three-electrode cell and reported a sensing performance of a linear range 1 fg/mL–1 µg/mL and LOD 0.01 fg/mL, which are significantly higher than the results reported in this study.²⁸ On the other hand, A.S. Tanak et al. performed the CRP detection using nEIS with ZnO electrode. They reported a result of linear range 0.01 µg/mL–20 µg/mL and LOD 0.10 µg/mL in human serum, and linear range 0.01 µg/mL–10 µg/mL and LOD 0.10 µg/mL in whole blood.²⁹ Therefore, although the measurement conditions are different, the performance of the proposed biosensor is estimated to be similar to that of nEIS measurements using a three-electrode system. In addition, the biosensor did not exhibit a clear impedance response to the increasing concentration of BSA used as a control sample. Therefore, these results demonstrate that the developed biosensor with an oBPE has the capability to detect antigen-antibody reactions and quantify CRP in a label-free, single-step reaction format, while ensuring adequate specificity and sensitivity.

Figure 6.

CRP concentration dependence of impedance spectrum. The vertical axis shows the rate of change in impedance from CRP concentration of 0 µg/mL to each CRP concentration (A). Dependence of impedance response of the biosensor at 30 Hz on the concentration of CRP as an analyte and BSA as a control sample (B).

4. Conclusions

In this study, we developed a simple, label-free electrochemical biosensor based on bipolar electrochemistry with an oBPE. The fabricated biosensor exhibited a clear impedance response to biomolecular events on the oBPE surface. It achieved sensitive detection of CRP over a concentration range (0.001 to 1.0 µg/mL) with linearity and reproducibility. This biosensing system does not require redox-active reagents and signal amplification steps, relying solely on a single-step antigen-antibody interaction for detection. The compact electrode configuration and simple measurement protocol suggest strong potential for miniaturization and low-cost fabrication, making the system suitable for point-of-care testing and other resource-limited diagnostic applications. Consequently, this work demonstrates the feasibility and utility of oBPE-based non-faradaic impedimetric biosensors as a platform for rapid, label-free, and cost-effective detection of biomolecules.

CRediT Authorship Contribution Statement

Arisa Tochigi: Funding acquisition (Lead), Investigation (Lead), Methodology (Lead), Visualization (Lead), Writing – original draft (Lead)

Mizuki Kondo: Methodology (Supporting), Supervision (Equal), Writing – review & editing (Supporting)

Takashi Kuwahara: Conceptualization (Lead), Funding acquisition (Lead), Supervision (Lead), Writing – review & editing (Lead)

Conflict of Interest

The authors declare no competing financial interests.

Funding

JSPS: JP21K04075

JSPS: JP23KJ1019

Uchida Energy Science Promotion Foundation

TAKEUCHI Foundation

NAGAI promotion foundation for science of perception

Footnotes

A. Tochigi: ECSJ Student Member

T. Kuwahara: ECSJ Active Member

References

• 1)  N. Shida, Y. Zhou, and S. Inagi, Acc. Chem. Res., 52, 2598 (2019).
• 2)  S. Masturah binti Fakhruddin, K. Ino, K. Y. Inoue, Y. Nashimoto, and H. Shiku, Electroanalysis, 34, 212 (2022).
• 3)  T. Kuwahara, K. Sato, M. Kondo, and M. Shimomura, Synth. Met., 198, 274 (2014).
• 4)  N. Shida, F. Kitamura, T. Fuchigami, I. Tomita, and S. Inagi, ChemElectroChem, 3, 465 (2016).
• 5)  M. Ohira, Y. Koizumi, H. Nishiyama, I. Tomita, and S. Inagi, Polym. J., 49, 163 (2017).
• 6)  K. Rahn and R. Anand, Anal. Chem., 93, 103 (2021).
• 7)  M. Sumitha and T. Xavier, Hybrid Adv., 2, 100023 (2023).
• 8)  E. Cesewski and B. N. Johnson, Biosens. Bioelectron., 159, 112214 (2020).
• 9)  M. S. Wilson and W. Nie, Anal. Chem., 78, 2507 (2006).
• 10)  N. Ruecha, J. Lee, H. Chae, H. Cheong, V. Soum, P. Preechakasedkit, O. Chailapakul, G. Tanev, J. Madsen, N. Rodthongkum, O.-S. Kwon, and K. Shin, Adv. Mater. Technol., 2, 1600267 (2017).
• 11)  Ö. A. Yokuş, F. Kardaş, O. Akyıldırım, T. Eren, N. Atar, and M. L. Yola, Sens. Actuators, B, 233, 47 (2016).
• 12)  W. Xu, K. Fu, and P. W. Bohn, ACS Sens., 2, 1020 (2017).
• 13)  E. Villani and S. Inagi, Anal. Chem., 97, 5837 (2025).
• 14)  H. Motaghi, S. Ziyaee, M. A. Mehrgardi, A. A. Kajani, and A.-K. Bordbar, Biosens. Bioelectron., 118, 217 (2018).
• 15)  X. Zhang, R. A. Lazenby, Y. Wu, and R. J. White, Anal. Chem., 91, 11467 (2019).
• 16)  E. Rafatmah and B. Hemmateenejad, Microchim. Acta, 186, 684 (2019).
• 17)  W. Zhao, Y. Ma, and J. Ye, J. Electroanal. Chem., 888, 115215 (2021).
• 18)  R. Liu, C. Zhang, and M. Liu, Sens. Actuators, B, 216, 255 (2015).
• 19)  M. Bhaiyya, P. Rewatkar, M. Salve, P. K. Pattnaik, and S. Goel, IEEE Trans. Nanobioscience, 20, 79 (2021).
• 20)  Z. Wang, Y. Z. Shen, M. Xu, J. Zhu, C. Ma, X.-Y. Hu, and Q. Xu, Sens. Actuators, B, 414, 135924 (2024).
• 21)  H. Adam, S. C. B. Gopinath, H. Krishnan, T. Adam, M. A. Fakhri, E. T. Salim, A. Shamsher, S. Subramaniam, and Y. Chen, Bioelectrochemistry, 161, 108800 (2025).
• 22)  C. Ibau, M. K. M. Arshad, S. C. B. Gopinath, M. Nuzaihan M.N., M. F. M. Fathil, and S. A. Shamsuddin, Int. J. Biol. Macromol., 162, 1924 (2020).
• 23)  G. N. Abdelrasoul, A. Afreen, S. MacKay, M. Tamura, M. A. Shah, D. P. Khasa, R. R. Montgomery, and A. I. Ko, Anal. Chim. Acta, 1107, 135 (2020).
• 24)  S. Black, I. Kushner, and D. Samols, J. Biol. Chem., 279, 48487 (2004).
• 25)  T. W. Du Clos, Ann. Med., 32, 274 (2000).
• 26)  M. Cui, Z. Che, Y. Gong, T. Li, W. Hu, and S. Wang, Chem. Eng. J., 431, 133455 (2022).
• 27)  T. Bryan, X. Luo, P. R. Bueno, and J. J. Davis, Biosens. Bioelectron., 39, 94 (2013).
• 28)  N. Dalila R, M. K. M. Arshad, S. C. B. Gopinath, C. Ibau, M. Nuzaihan M.N., M. F. M. Fathil, U. Z. Mohd Azmi, and P. Anbu, J. Taiwan Inst. Chem. Eng., 138, 104450 (2022).
• 29)  A. S. Tanak, B. Jagannath, Y. Tamrakar, S. Muthukumar, and S. Prasad, Anal. Chim. Acta X, 3, 100029 (2019).