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Abstract
Surface plasmon resonance (SPR) has emerged as one of the most efficient and attractive techniques for optical sensors in biological applications. The traditional approach of an EC (electrochemical)-SPR biosensor to generate SPR is by adopting a prism underneath the sensing substrate, and an angular scan is performed to characterize the reflectivity of target analytes. In this paper, we designed and investigated a novel optical biosensor based on a hybrid plasmonic and electrochemical phenomenon. The SPR was generated from a thin layer of gold nanohole array on a glass substrate. Using C-Reactive Protein (CRP) as the target analyte, we tested our device for different concentrations and observed the optical response under various voltage bias conditions. We observed that SPR response is concentration-dependent and can be modulated by varying DC voltages or AC bias frequencies. For CRP concentrations ranging from 1 to 1000 µg/mL, at the applied voltage of -600 mV, we obtained a limit of detection for this device of 16.5 ng/mL at the resonance peak wavelength of 690 nm. The phenomenon is due to spatial re-distribution of electron concentration at the metal-solution interface. The results suggest that CRP concentration can be determined from the SPR peak wavelength shift by scanning the voltages. The proposed new sensor structure is permissible for various future optoelectronic integration for plasmonic and electrochemical sensing.
© 2022 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
Surface plasmons (SPs) are electromagnetic waves at the interface between the metal and dielectric. Although light cannot directly excite the SPs on a smooth metallic surface, surface plasmon resonance (SPR) can be achieved through prism or grating coupling [1]. SPs are particularly sensitive to the dielectric constant (index of refraction) at the interface and can be utilized for detecting surface binding [2]. This property has been used in many SPR-based biosensors [3,4]. SPR combined with electrochemical (EC) biasing, a technology called EC-SPR, exquisitely adopts optical characterization using SPR and electric potential modulation of EC impedance spectroscopy (EIS) [5,6] for biosensing applications. It is immune to the undesired signals from bulk refractive index changes, which improves detection selectivity and sensitivity [7].
A typical EC-SPR experimental setup utilizes the Kretschmann configuration [8], which comprises a thin metal layer on the front side of the substrate and a prism on the backside for SPR excitation. One key parameter of EC-SPR detection is the light incident angle, correlated with the resonant wavelength in the reflectivity spectrum. However, it is challenging to integrate microfluidics with SPR-based sensing at the device level. There are difficulties in adopting large numerical aperture optics for high spatial resolution and in scaling down the optical apparatus (such as the prism) required for reflection mode operation. Despite the broad applicability and excellent sensitivity of the typical SPR configuration, the experimental setup is complicated and time-consuming [9].
Periodic nanoholes in a thin metal sheet have been demonstrated to be a unique and viable method for light-SPR coupling [10,11]. In comparison to prism coupling, nanohole arrays have several unique advantages. First, operating at normal incidence simplifies the optical alignment, enabling future miniaturization and integration of supporting optics at the device level [1,12]. Second, the nanohole array has a small footprint, allowing for downsizing and integration into microfluidic systems and improved spatial resolution [13,14]. Third, in contrast to previous local surface plasmon techniques based on colloidal nanoparticles [15] or roughened surfaces, they can be fabricated with remarkable reproducibility. Last, nanohole arrays can be compatible with electrical detecting principles because the continuous metal layer is a conductive electrode [16]. Since metal plate nanohole arrays are fabricated using a semiconductor process, they can be easily miniaturized and integrated with other sensor devices. Extraordinary SPR based on chemo- and biosensors can be achieved by generating SPR from nanohole arrays [13,17].
The resonance peak in SPR is determined by the dielectric constants of the metal layer as well as the dielectric characteristics of media adjacent to the metal film. The application of an electric potential affects the dielectric characteristics near the metal-solution interface, leading to a variation in SPR signals [18]. It has been demonstrated that the formation of an electric double layer (EDL) [19] considerably impacts the SPR response. The change in SPR peak with electric potential is significant [20]. The capability to qualify the electric field locally using an SPR approach with high spatial and temporal resolution can be used in various bioscience applications [21].
Compared with standard SPR devices, the benefits of using plasmonic biosensors based on nanostructured metallic films are significant. Although studies on nanohole-based refractometric sensing have grown increasingly, there have been few reports in this field on implementing electrical measurements [22]. In this paper, we fabricated a nanohole array SPR sensor designed to be integrated with electrochemical techniques to characterize the sensing properties of C-Reactive Protein (CRP). We demonstrate the voltage and frequency dependencies of the SPR peaks. Understanding the effect of electrical potential on SPR phenomena paves the way for the combination of optical and electrochemical real-time monitoring.
2. Fabrication and measurement
2.1 Sensor fabrication
The sensor was fabricated on a glass substrate. The schematic steps to fabricate gold nanohole arrays are shown in Fig. 1. We first deposited Ti/Au with a thickness of 5/50 nm, followed by a 150 nm-thick SiNx dielectric layer coating by plasma enhanced chemical vapor deposition (PECVD). Nanoholes were next defined by electron beam lithography, and the subsequent RIE (reactive ion etching) to remove gold and SiNx in the holes. The diameter of the hole is 250 nm with a period of 750 nm. The sensor comprises a central electrode integrated with an gold nanohole array, which is surrounded by ring electrodes for electrochemistry. As shown in Fig. 2 (a), the nanohole plate is connected to the working electrode (in yellow). The concentric rings (red and green) encircling the working electrode function as the electrokinetic flow generator for gathering CRP molecules. Figure 2 (b) shows the nanohole pattern's scanning electron microscope (SEM) image.
Fig. 2. (a) Schematic diagram of the biosensor, which comprises ACEK electrodes integrated with a gold nanohole array. (b) SEM image of the nanohole array (scale bar: 500 nm). (c) Surface modification protocol on the gold nanohole array. (d) An experimental setup for SPR measurement.
2.2 Surface modification
Figure 2 (c) depicts the schematic diagram of the electrode modification procedure. First, we immobilized 5 µL of the anti-CRP aptamer (5'-SH AAA AAA GCC TGT AAG GTG GTC GGT GTG GCG AGT GTG TTA GGA GAG ATT GC-3’, purchased from Protech Technology Enterprise Co., Ltd.) with a concentration of 1 µM on the gold nanohole plate for 30 minutes. The surface was blocked with 10 µL of 95% ethanol solution (with 1 mM of mercaptohexanol (MCH), purchased from Sigma-Aldrich) for 15 minutes. Finally, ACEK electrodes were biased to behave as an electrokinetic flow generator for gathering the target analyte, CRP molecules, to bind with the anti-CRP aptamers.
2.3 Target analyte
The target analyte adopted in this study is CRP, one of the most well-known acute phase proteins composed of five subunits of the same polypeptides [23]. In physiological situations, acute-phase proteins exist in the blood, and their production increases or decreases fast in response to numerous diseases or threats [24], such as trauma, infection, acute myocardial infarction, and cancer. Blood CRP levels can rise to 1000 times after stimulation, such as acute inflammation, and have a half-life of 19 hours before reverting to normal [25]. CRP plasma concentrations are typically less than 1 µg/mL under normal conditions [26]. CRP was purchased from Sigma Aldrich. The concentration of CRP was adjusted from 1 to 1000 µg/mL using 10 mM phosphate-buffered saline (PBS). The PBS solution contains KCl (2 g/L), KH2PO4 anhydrous (2 g/L), NaCl (80 g/L), Na2HPO4 anhydrous (11.5 g/L) mixed with potassium ferricyanide (K3Fe(CN)6) and potassium ferrocyanide (K4Fe(CN)6) salt.
2.4 Measurement setup
The measurement setup is shown in Fig. 2 (d). White light generated from a Xenon lamp (ASB-Xe-175) is coupled through an optical fiber and incident perpendicular to the sample. The reflected light is also collected by the fiber and then characterized by the spectrometer (Ocean Optics, HR 4000). The electrodes were connected to a potentiostat (SP-50) to provide both DC and AC electrical signals. This work applied AC signals with sweeping frequencies from 0.7 Hz to 100 kHz and DC voltages from -0.6 to 0.6 V.
3. Theory
This section provides a brief analysis of SPR peak wavelength shift at the metal-solution interface when the concentration of the solution is changed. The optical spectral variations when the system is exposed to an external voltage are also addressed.
3.1 SPR peak wavelength shifts with different target concentrations
First, we explain the shift in SPR peak through the wavelength interrogation technique, in which the resonance peak is obtained by tuning the wavelength while keeping the incident angle fixed. When the polychromatic light is shone onto the gold nanohole array, a longitudinal wave called surface plasmonic wave is generated. Since the plasmonic response of the material depends on the complex dielectric constant of metal, ${\varepsilon _m}$, described by Drude’s model [27] and the refractive index of the sensing media ${n_s}$, the phase matching could be explained by:
Compared with plasmonic sensors, a nanohole array structure crafted on the metal surface enhances the efficiency of refractive index-based sensors [28,29]. With nanohole arrays, the momentum matching condition in the sub-wavelength region can be achieved without using the prism coupling technique. The modified phase matching condition becomes,
3.2 Simulation model
We performed a simulation on the electric field distribution using the finite difference time domain (FDTD) method to demonstrate the performance enhancement of the proposed nanoholes. A comparative study was performed on two different structures. A 50 nm-thick gold film (see Fig. 3(a)) and a nanohole array of diameter 250 nm and period 750 nm on the same thickness of gold (Fig. 3(b)) were compared. The gold film is sandwiched between a BK7 glass substrate with the refractive index of 1.52 and the medium on top of the gold has a refractive index of 1.33. Under normal incident of light, SP modes are generated at the gold-glass interface. Compared with the gold layer-based sensor, the electric field distribution of the nanohole array sensor in Fig. 3 (b) clearly shows an evident improvement at resonance wavelength on top of the gold metal pads. Under normal incident of light, photons tunnel through the gold layer and excite the SPs at the gold -solution and gold -glass interfaces [30]. With nanohole arrays, a small imprint of SPs can be generated and detected by performing a collinear setup using ordinary light source (Xenon Lamp without using a highly focused laser) for a small volume (<1 µg/mL) of sensing solution. A similar comparative study was addressed by M. Couture et al. [31] and Pang et al. [32], showing nanohole array-based SPR sensor has high refractive index sensitivity. They also demonstrated that the sensitivity was correlated to the size of nanoholes (diameter and period).
Fig. 3. Simulated electrical field intensity distribution showing the generation of surface plasmons at (a) 50 nm gold thin film, and (b) 55 nm-thick gold nanohole arrays with a diameter of 250 nm and period of 750 nm (c) Sideview of the sensing area in (b). In the subsequent experiment, the CRP solution was dispensed on top of the sensor, and light was incident on the nanohole array. A bias voltage was applied during measurement. (Top figure of (c): a close-up view of the nanohole array in which the potential decay with the distance from the gold surface).
3.3 SPR peak wavelength shifts with the applied voltage
The metal-electrolyte interface can be the best candidate to explain the EDL phenomenon because the metal surface has excess positive charges, which can be balanced by a countercharge of the opposite sign by the electrolyte at the interface.
To explain the EDL phenomenon, we adopt the Gouy-Chapman-Stern model [33]. When the external voltage is applied to the metal-solution interface, the innermost layer of ions adsorbed on the metal surface is called the Stern layer. The remaining counter-ions are in the solution, making the system electroneutral. A potential profile is embedded in Fig. 3 (c), where ${V_0}$ and ${V_1}$ indicate the potentials at the gold surface, and Stern layer, respectively. The free electrons on the gold surface attract the counter-ions in the solution. A tightly bound counter-ions (Stern layer) are deployed on the metal surface, forming a capacitor. With the electrostatic interaction between gold's free electrons and ions (cations) in the solution, the potential in this region (from ${V_0}$ to ${V_1}$) decays linearly. The ions beyond the Stern layer further attract counter-ions (anions) but experience repulsive coulomb force from electrons in the gold layer. Hence another capacitor is formed with the value depending on the applied potential and the concentration of electrolyte [34]. The structure behaves like two capacitors connected in series.
The charging of the EDL capacitor leads to the free electron profile re-distribution of the gold surface. It further changes the free electron concentration of the metallic surface layer. In such a case, the free electron variations result in the corresponding change of the relative dielectric constant of the layer [35]. Hence the dielectric constant of the gold layer can be expressed as the summation of the unperturbed dielectric constant of the metal and the induced dielectric due to the formation of a double layer. Since the dielectric constant of metal is directly related to the plasmonic vibration, the applied voltage alters the plasma frequency of metal. The change in electron charge density linearly affects the plasma frequency of metal, which is responsible for the further shift in the resonance peak.
3.4 SPR peak wavelength shifts with AC frequency
SPR measurement's sensitivity can be enhanced by applying an electric potential [18]. This sensitivity enhancement was also discovered to be frequency dependent, wherein the shift in the wavelength of the SPR response is more prominent at lower frequencies. The frequency dependence is due to the adsorption of electrolyte ions on the surface of the electrode. When an AC signal is applied, the electrodes behave as an electric capacitor, where one plate is the gold film and the other is the adsorbed ions. This capacitance can be described as:
wherein A is the area, and $\Lambda $ is the distance between the two plates. Based on this interpretation, the variables in the numerator of Eq. (3) are all constant, which signifies that the value of C changes with the distance ($\Lambda $). The frequency dependence of the double layer capacitor was reported to be caused by its capacitance behavior [36], which can be described as: wherein ${C_{f = 1\; Hz}}$ is the capacitance at a low frequency, f is the applied frequency, and $\alpha $ is a constant. It can be observed in Eq. (4) that the capacitance varies indirectly with the frequency, where an exponential decay can be obtained upon the increase in the frequency. Equation (3) and 4 state that the increased applied frequency leads to an increased distance between the two plates of the capacitor (electrode and electrolyte), thus an exponential increase in the capacitance. At higher frequencies, there will not be enough induced electric charges on the electrode surface, while at lower frequencies, excess induced charges will be attracted to the surface [37].4. Results and discussion
4.1 Characterizations of SPR spectra with various CRP concentrations
Prior to the characterization of the SPR spectra with various CRP concentrations, the reflection spectra of bare sample, with CRP aptamer, and with MCH, all dispensed with PBS solution, were recorded and shown in Fig. 4 (a). Two spectral dips were observed, undergoing a red shift with the inclusion of anti-CRP aptamer and a further shift with MCH for surface blocking. As the CRP solutions with different concentrations were dispensed to the nanohole arrays, the reflection spectra are shown in Fig. 4 (b). Again, both spectral dips experience a red shift when the concentration, and thus the refractive index, increases. In Fig. 4 (a) and (b), the SPR peak around 611 nm is associated with plasmon generation at the solution-metal interface, while the one at around 690 nm is attributed to the metal-glass interface [38]. Since light is normally incident on the sample from the air, the plasmonic peak wavelength at the solution-metal interface is calculated from the following equation [39],
Fig. 4. (a) Reflection spectra of the bare device, after applying CRP aptamer, and after applying MCH (all dispensed with PBS solution). (b) Reflection spectra of the sensor with the CRP concentration ranging from 1 to 1000 µg/mL. (Inset : a close-up view of the spectra near 690 nm)
On the other hand, the plasmonic peak at the metal-glass interface is also affected by the solution's refractive index since light transmits through the solution. Hence, the plasmonic resonant peak shift in the metal-glass interface is larger than that at the solution-metal interface, in which the resonant peak can be expressed by the following equation [38], with ${n_g}$ being the refractive index of glass,
From Eq. (5) and 6, the SPR peak wavelengths depend on the periodicity of the nanohole lattice and the change of the dielectric constant in the surrounding medium. SPs can travel along the Au-dielectric interface and be confined to the gold nanoholes’ edges, enhancing plasmon resonant peak sensitivity [40]. The results in Fig. 4 (b) were further analyzed by plotting the resonant peak wavelength shifts of CRP concentrations from that of MCH. As depicted in Fig. 5 (a), the resonant peak around 611 nm was red-shifted linearly with the logarithmic scale of the CRP concentration within the range of 1 and 1000 µg/mL. A fitting coefficient (R2) value of 0.9973 was obtained. In Fig. 5 (b), the resonant peak around 690 nm was also red-shifted linearly with the log scale of CRP concentration with an R2 of 0.9777. A well-fitting curve can detect concentrations as low as 1 µg/mL, which is lower than the typical detection ranges of CRP using point-of-care testing (POCT) [41]. Since the plasmonic resonant peak shift at the wavelength around 690 nm is larger than that around 611 nm, we only characterize SPR around 690 nm under various DC and AC bias conditions in the subsequent experiments.
Fig. 5. Wavelength shift of the resonant peak around (a) 611 nm and (b) 690 nm, extracted from Fig. 4(b).
4.2 Characterizations of SPR spectra with various DC applied voltages
The plasmonic properties under various applied voltages were explored next. The SPR spectrum at a CRP concentration of 1 µg/mL changes with the applied DC voltage, as represented in Fig. 6 (a). The corresponding SPR peak wavelengths around 690 nm were plotted against various DC voltages (see Fig. 6 (b)). The linear spectral response with R2 of 0.9945 is attributed to the formation of the EDL, as illustrated in Fig. 3 (c). When the gold-electrolyte interface is exposed to an external electric field, electrochemical processes at the interface result in the EDL, creating a parallel plate capacitor. Ions at the interface are polarized, changing the concentration of ions, which further leads to changes in the dielectric constant of metal. Since the plasmonic resonance is directly related to the charge carrier concentration at the surface, a shift in the SPR peak was observed with the applied voltage.
Fig. 6. (a) SPR spectra of CRP concentration of 1 µg/mL at various applied voltages. (Inset : a close-up view of the spectra near 690 nm). (b) The corresponding wavelength shift at around 690 nm at various applied voltages.
In Fig. 7 (a), voltage-dependent resonant peak wavelengths were plotted against different concentrations of CRP. First, as shown in Fig. 7 (b) the amount of wavelength shift in detecting CRP is bias-dependent. It was observed that the wavelength shift is the highest at the applied voltage of –600 mV and gradually decreases when the voltage swings toward positive. The phenomenon is attributed to the charging of the voltage-dependent EDL capacitor. Under a negative voltage bias, a larger number of electrons transfer towards the interface, while positive ions are expelled away from the surface. In the extreme case, a blue shift, instead of red shift, of the wavelength is observed at a CRP concentration of 1 µg/mL at -600 mV. The phenomenon is mainly attributed to the attraction of positive ions in the solution to the sensor surface due to a large (compared with other applied voltages in the experiment) negative voltage on the electrode. The blue shift indicates the generation of opposite polarity of charges in the solution near the sensor metal-solution interface. Second, though it is unsurprising to observe the wavelength shift with the CRP concentration under applied bias, the slope of the curve is also concentration-dependent from Fig. 7 (c). The phenomenon can be explained by the surface binding when the CRP concentrations are high. Since more CRP molecules were bound to the surface of gold, electrons transferred to the gold plate were blocked by the immobilized multi-molecular layer, which reduces the influence of the EDL. Electron transfer to the gold plate becomes more efficient when the concentration becomes lower, thus enhancing the wavelength shift. The above results suggest that our proposed techniques are more efficient in distinguishing CRP at low concentrations with the help of an external electric field. Moreover, the lowest analyte concentration that can be detected is referred as LOD. Based on signal-to-noise, the LOD is expressed as 3 SD/s [42], in which SD is the standard deviation of three samples tested in our experiment and the s is the slope of the curve in Fig. 7(b). At -600 mV, the calculated LOD value for our SPR detection device is 16.5 ng/mL.
Fig. 7. (a) Peak wavelength shift around 690 nm vs. DC voltage at various CRP concentrations. (b) Wavelength shift vs. CRP concentrations at different DC bias voltages (The curves are extracted from (a)). (c) The slope of the curve fitting in (a) of each concentration.
4.2 Characterizations of SPR spectra with various AC frequencies
We explored how the resonant peak wavelength changed with externally applied frequencies. With the AC signals sweeping between - 12 and -2 mV at frequencies from 0.7 to 100 kHz, the reflection spectra of the CRP concentration of 1 µg/mL were plotted in Fig. 8 (a). The wavelength shift was extracted and shown in Fig. 8 (b). The SPR wavelength shift decreased from 1.57 to 0.24 nm when the frequency increases from 0.7 Hz to 100 kHz. It was previously reported that the thickness of the double layer decays exponentially with the increase of applied frequency [34], which can be verified from Fig. 8 (b) that the wavelength shift was observed to decay exponentially with increased frequency. This frequency dependence can be attributed to the movement of adsorbed ions on the electrode surface. Fewer ions are attracted toward the electrode surface at higher frequencies. In contrast, excess ions are attracted toward the surface at low frequencies, producing more wavelength shift because the amount of delocalized ions increases [35]. With the range of DC bias voltages (see Fig. 6 (b)) and AC sweeping frequencies (Fig. 8 (b)), the amount of SPR wavelength shift with the DC bias voltages is more significant than that with AC frequency. It implies that characterizing CRP concentrations under a selective DC bias is preferred when the SPR spectral response is considered.
Fig. 8. (a) SPR spectra of CRP concentration of 1 µg/mL at various AC frequencies. (Inset : a close-up view of the spectra near 690 nm) (b) The corresponding wavelength shift around 690 nm at various AC frequencies.
4.3 Discussions on the voltage-controlled SPR biosensing
Characterizations of SPR properties using the proposed device structure indicate a new method of voltage-controlled SPR sensor. In contrast to conventional EC-SPR biosensors in which an angular scan is performed to characterize the reflectivity of target analytes, our sensors can correlate the SPR peak wavelengths to CRP concentrations by providing different DC voltages or AC bias frequencies. It implies that CRP concentration can be determined from the SPR peak wavelength shift by sweeping the voltages. As a result, a real-time measurement can be achieved with the proposed sensor because modern electronic equipment can easily provide a fast voltage sweep. In addition, voltage-resonant peak wavelengths measurement suggests that the corresponding slope is concentration-dependent (see Fig. 7(b)), which provides an additional parameter to benchmark target analytes.
5. Conclusion
We have experimentally investigated the sensing response of SPR sensors with external bias voltages for biological research applications. Unlike the typical prism configuration, the SPs were generated from nanohole arrays, which avoid characterizing plasmonic properties at various incident angles. We tested our device for different CRP concentrations and observed the optical response. The 16.5 ng/mL detection limit of CRP is achieved from the resonant peaks at the wavelength of around 690 nm at the bias voltage of -600 mV. We further investigated the behavior of the SPR in response to different applied voltages. The SPR peak showed a linear shift with voltage. The voltage-dependent SPR peak shift can be explained from the EDL. When the Au-electrolyte system was subjected to an external voltage, electrochemical reactions at the interface occurred, resulting in the formation of a parallel plate capacitor
Funding
National Science and Technology Council, Taiwan (111-2218-E-002 -025 -, 111-2221-E-002 -188 -MY3).
Disclosures
The authors declare no conflicts of interest regarding this article.
Data availability
Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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