Electrochemical Biosensors - Sensor Principles and Architectures

Abstract

Quantification of biological or biochemical processes are of utmost importance for medical, biological and biotechnological applications. However, converting the biological information to an easily processed electronic signal is challenging due to the complexity of connecting an electronic device directly to a biological environment. Electrochemical biosensors provide an attractive means to analyze the content of a biological sample due to the direct conversion of a biological event to an electronic signal. Over the past decades several sensing concepts and related devices have been developed. In this review, the most common traditional techniques, such as cyclic voltammetry, chronoamperometry, chronopotentiometry, impedance spectroscopy, and various field-effect transistor based methods are presented along with selected promising novel approaches, such as nanowire or magnetic nanoparticle-based biosensing. Additional measurement techniques, which have been shown useful in combination with electrochemical detection, are also summarized, such as the electrochemical versions of surface plasmon resonance, optical waveguide lightmode spectroscopy, ellipsometry, quartz crystal microbalance, and scanning probe microscopy. The signal transduction and the general performance of electrochemical sensors are often determined by the surface architectures that connect the sensing element to the biological sample at the nanometer scale. The most common surface modification techniques, the various electrochemical transduction mechanisms, and the choice of the recognition receptor molecules all influence the ultimate sensitivity of the sensor. New nanotechnology-based approaches, such as the use of engineered ion-channels in lipid bilayers, the encapsulation of enzymes into vesicles, polymersomes, or polyelectrolyte capsules provide additional possibilities for signal amplification. In particular, this review highlights the importance of the precise control over the delicate interplay between surface nano-architectures, surface functionalization and the chosen sensor transducer principle, as well as the usefulness of complementary characterization tools to interpret and to optimize the sensor response.

Keywords: bioelectronics.; biosensors; electrochemistry; review.

Figure 1.

Elements and selected components of a typical biosensor [1, 2, 3].

Figure 2.

A sample amperometric measurement: According to Kueng et al. [28] this is a typical hydrodynamic response of their biosensor to glucose followed by several injections of ATP measured in phosphate buffer at 650 mV in reference to Ag/AgCl. The change in current response is proportional to the ATP concentration as glucose is consumed at the glucose oxidase (GOD) and hexokinase (HEX) modified electrode surface.

Figure 3.

The schematic and performance results of a potentiometric sensor for the detection of hybridization between complementary single-stranded (ss) oligonucleotides on the phase boundary PVC membrane/solution. The arrangement of the probe is shown in two configurations where the ss oligonucleotide is attached along (I) or parallel (II) to the PVC membrane surface. The potentiometric detection of the hybridization with complimentary strands is shown: (A) with complementary and unspecific interactions with non-complementary oligonucleotides; (B) with complementary oligonucleotides; (C) demonstration of the reproducibility after regeneration in a buffer of 0.01 M NaOH for 15 minutes (upward arrow) and subsequent injections of the corresponding oligonucleotides (downward arrow) [38].

Figure 4.

An overview of selected cyclic voltammograms with reference to the work of Li et al.: i) a single generic linear voltage sweep; ii) a cyclic voltammogram of the Prussian Blue-modified (PB) glassy carbon electrode was prepared in 0.1 M KCl + 0.1 M HCl at a scan rate 50 mV/s; iii) cyclic voltammograms of the PB sol-gel modified glassy carbon electrode and iv) the cholesterol/PB sol-gel modified glassy carbon electrode in a phosphate buffer (pH 6.8) at varying concentrations of the analytic solution: a) blank solution; b-f) blank solution + cholesterol 1×10⁻⁵, 2×10⁻⁵, 3×10⁻⁵, 4×10⁻⁵, 5×10⁻⁵ mol/L. [47]. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.

Figure 5.

Chronoamperometric curves for detecting cholesterol in a phosphate buffer (pH 6.8), using Prussian Blue (PB) modified electrodes and cholesterol oxidase (ChOx)/PB sol-gel electrodes. The working electrode was polarized at -0.05 V (vs. Ag/AgCl) in a stirred solution. A 3σ detection limit of 5.4 × 10⁻⁷ M [47]. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.

Figure 6.

a) i) Immunosensor for the in-plane impedance measurements between conductive electrodes, ii) interdigitated electrode for the in-plane impedance immunosensing [60]. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission; b) illustration of the principle behind the change in network impedance upon hybridization of ssDNA probe oligomers to its complement interdigitated microsensor electrode [16]. With kind permission of Springer Science and Business Media.

Figure 7.

Structure and principal function of a penicillin-sensitive EnFET. The enzyme penicillinase is immobilized on a pH-sensitive Ta₂O₅ surface [11]. The reference electrode is labelled as RE. Reproduced by permission of The Royal Society of Chemistry.

Figure 8.

Schematic diagram of a silicon nanowire FET [84]. The field effect occurs as charged particles or molecules bind to the surface of the nanowire and change its conductive properties, analogous to the gate voltage and channel conduction of a conventional FET.

Figure 9.

i) A schematic representation of the design to use silicon nanowire FET as a biosensor for the label-free detection of DNA hybridization; ii) the change of resistance of the silicon nanowire array as a function of hybridization time ((1) 1.0 nM non-complementary DNA as control, (2) 25 fM, (3) 100 fM and (4) 1.0 nM target DNA in buffer); iii) the change of resistance of the silicon nanowire array as a function of hybridization time in 25 fM of (1) fully complementary, (2) one-base mismatched, and (3) two-base mismatched target DNA in TE buffer. The TE buffer solution (pH 8.5 10 mM Tris-HCl-1.0 mM EDTA-30 mM NaCl buffer) was used as the hybridization, washing, and working buffer. Reprinted with permission from [87]. Copyright 2007 American Chemical Society.

Figure 10.

An illustration to demonstrate the concept of the influence of surface interactions on nanowire conduction. In larger wires, for example with micrometer dimensions, the surface-to-volume ratio is relatively small. Even if surface interactions, such as binding with charge particles or biological species, influence conduction near the surface of the wire, there is still a large portion of the wire's interior available for uninterrupted conduction. To help visualize this, the illustration uses a mental construct of ‘conduction channels’ within the wire. As the dimensions of a wire decrease to the nanometer regime, the surface-to-volume ratio drastically increases. Therefore, the same external influences increasingly influence the conduction both on the wire surface and in the wire interior, thereby decreasing the possibility for uninterrupted conduction.

Figure 11.

An overview of the experiment by Foley et al. to measure the step-wise decrease in conductance as a gold nanowire is thinned by pulling. The discrete values of the conductance in units of 2e²/h has been corrected for the impedance of the measurement equipment. Reprinted with permission from [95]. Copyright 1999, American Institute of Physics.

Figure 12.

The image on the left illustrates the sensing principle of nanowire-based biosensing with an FET configuration. The image on the right demonstrates the concept of multi-analyte biosensing with nanowire FET configurations [99]. Reproduced by permission of the MRS Bulletin.

Figure 13.

Array of nanowires made from DNA-tagged Au nanoparticles using line patterns made with EUV-IL [24].

Figure 14.

A modified schematic illustration of a combined surface-plasmon optical and electrochemical setup from Badia et al. [110].

Figure 15.

Measured results of the biosensor of Kang et al. for the detection of the enzymatic reaction between horseradish peroxidase and a film of polyaniline (PAn) on a gold electrode in the presence of H₂O₂: i) CV curve (a) and reflectance change (ΔR) (b) as a function of a potential at a fixed incident angle of 62.2°; ii) SPR measurements of reflectivity versus incidence angle at varying electrochemical conditions 1) bare gold electrode 2) reduced PAn film after maintaining a potential of 0 V for 5 minutes 3) oxidized PAn film after maintaining a potential of 0.9 V for 5 min iii) SPR kinetic curves at a fixed angle of 62.8° resulting from the HRP/PAn reaction at various concentrations of H₂O₂; iv) the relationship between the slope of the ΔR-t curves from iii) and the H₂O₂ concentration [116].

Figure 16.

EC-OWLS flow cell. The laser is coupled into the grating at a certain incoupling angle. The ITO-coated waveguide acts as the working electrode, the reference electrode is placed on the topside of the flowcell and the counter electrode at the outlet.

Figure 17.

A) EC-OWLS curve showing adsorption as a function of time. Additionally, the signal also reacts to an applied potential. Prior to the adsorption steps the system was electrically cycled to obtain stable conditions. PLL-g-PEG/biotin, neutravidin and biotin conjugated GOx were adsorbed in three steps with buffer rinsing in between. After completing the adsorption, a potential of 1 V was applied several times in order to measure the current generated by the enzymes at different substrate concentrations. B) Current after 10 min as a function of the glucose concentration. The applied potential was 1 V as stated on the left side.

Figure 18.

An overview of an SPM technique combined with electrochemical techniques; a) SEM image and schematic top view of an AFM-tip-integrated frame electrode; b) schematic cross section of an AFM-tip integrated biosensor. Enzymes are shown to be immobilized at the surface of the scanning-probe tip-integrated electrode; c) schematic cross section of the experimental setup for imaging glucose transport through a porous membrane during simultaneous AFM mapping [157]. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.

Figure 19.

Topographic images obtained using constant-current mode of SECM imaging as presented by Kurulugama et al. of (A) an undifferentiated PC12 cell, (B) early neurite development and (C) a differentiated PC12 cell. In their work they also provide images using other feedback modes, namely constant-impedance and constant-distance. Reprinted with permission from [166]. Copyright 2005 American Chemical Society.

Figure 20.

a) Simplified scheme for magnetically labelled biomolecule detection. The magnetic label, functionalized with biomolecule A, interacts with the complementary biomolecule B at the magnetoresistive sensing surface. The resulting magnetic fringe field changes the resistance of the magnetoresistive sensor, which is measured as a change in voltage (ΔV) at a fixed sensing current (I). b) Simplified scheme for the use of magnetically labelled streptavidin to detect the location of pre-hybridized biotinylated DNA. First, the immobilized DNA on the magnetoresistive sensor surface is hybridized with biotinylated target DNA. Second, the magnetically labelled streptavidin is introduced to detect the hybridized DNA by binding to the biotinylated hybridized target DNA. The magnetic fringe field of the labels is detected by the magnetoresistive sensors as in a) [185].

Figure 21.

Scheme of direct and indirect electron transduction. a) Direct transduction: the electrons only generate a measurable current if the reaction takes place close to the surface. b) Indirect transduction: with the help of a mediator, that shuttles the electrons between the reaction site and the surface, much larger distances can be overcome. c) Enzyme bound via conducting CNT. d) ELISA-like sandwich setup, where an enzyme labelled antibody binds to the detected antigen.

Figure 22.

Schematics of supported lipid bilayer architectures used for sensor applications including electrochemical biosensors: a) hybrid bilayer, b) solid-supported lipid bilayer (SLB), c) tethered lipid bilayer (t-SLB), d) polymer-supported bilayers (pSLB), e) pore-spanning lipid bilayers (nano-BLM) and f) a nano-BLM with an incorporated protein for single channel recordings.

Figure 23.

Upon becoming triggered to the open state, the pore filters and selects only those specific ions or charged molecules that are designed to pass through the channel [284]. From Doyle, D. A. et al., SCIENCE 280:69-77 (April 3, 1998). Reprinted with permission from AAAS.

Figure 24.

Schematic of an ion-channel switch (ICS) biosensor as created by Cornell et al.: a) the addition of the target analyte inhibits channel formation by anchoring the pore, which lowers the electrochemical conductivity of the membrane, b) an alternative mechanism, where the introduced analyte replaces an existing bond and releases the pore. This then promotes the formation of channels and increases the membrane conductance [286].