Innovations in biomedical nanoengineering: nanowell array biosensor
© The Author(s) 2018
Received: 8 February 2018
Accepted: 26 March 2018
Published: 11 April 2018
Nanostructured biosensors have pioneered biomedical engineering by providing highly sensitive analyses of biomolecules. The nanowell array (NWA)-based biosensing platform is particularly innovative, where the small size of NWs within the array permits extremely profound sensing of a small quantity of biomolecules. Undoubtedly, the NWA geometry of a gently-sloped vertical wall is critical for selective docking of specific proteins without capillary resistances, and nanoprocessing has contributed to the fabrication of NWA electrodes on gold substrate such as molding process, e-beam lithography, and krypton-fluoride (KrF) stepper semiconductor method. The Lee group at the Mara Nanotech has established this NW-based biosensing technology during the past two decades by engineering highly sensitive electrochemical sensors and providing a broad range of detection methods from large molecules (e.g., cells or proteins) to small molecules (e.g., DNA and RNA). Nanosized gold dots in the NWA enhance the detection of electrochemical biosensing to the range of zeptomoles in precision against the complementary target DNA molecules. In this review, we discuss recent innovations in biomedical nanoengineering with a specific focus on novel NWA-based biosensors. We also describe our continuous efforts in achieving a label-free detection without non-specific binding while maintaining the activity and stability of immobilized biomolecules. This research can lay the foundation of a new platform for biomedical nanoengineering systems.
Nanoarrays, which consist of probe biomolecules immobilized on a chemically modified surface, have attracted much attention for the development of nanobiosensors. However, there is no simple nanometric system that can commonly be used for high-selectivity nanobiosensor applications. This NWA structures with functional materials at the nanometer scale have become more important as technology advances and the paradigm shifts from single material and 2D planar structures to complex compounds and 3D architectures. It allows for the fabrication of unique and integrated devices such as gas sensors , optical photonic crystals , fluidic devices , stretchable electronics , ring shapes with ferromagnetic materials , and so on. Especially, nanoimprint lithography (NIL) is a promising technique that can produce low-cost, large-area fabrication, high throughput [34, 35]. However, pattern resolution is primarily determined by the mold and E-beam lithography (EBL) or Focused ion-beam lithography (FIB), and NIL cannot overcome the resolution limits of the patterning tools and is often limited by the type of materials [36–38]. NIL was used to fabricate protein nanoarrays with an inert poly ethylene glycol (PEG) polymer as the resist material. The most widely used inert materials are uncharged PEG-based polymers [39, 40] and self-assembled monolayers (SAMs) [41–43], but the latter is incompatible with the NIL process owing to difficulties in direct imprinting. We developed a very effective and widely applicable method for fabricating nanopatterns of a PEG hydrogel for protein NWA.
Construction of such NWAs on Au electrode with controlled dimension and density could allow for quantitative analysis on the level of single lipid vesicle with more reaction sites and low signal-to-noise (S/N) ratio. Non-specific binding has been a hurdle to most label-free detection methods since many proteins present in a biological sample adhered to the surface non-specifically, giving false positives. Strategies to overcome this problem involve specific chemistries, which limit non-specific binding, but they do not completely eliminate non-specific adsorption [29, 30, 44–46]. Thus, it would be of great benefit to develop a NWA electrode with controlled geometry and density that can capture individual liposomes without non-specific binding. To block non-specific binding, UV-curable PVA hydrogel, which acted as an inert barrier against non-specific adsorption of liposome was used so liposome only can bind to Au exposed area.
Electrochemical impedance spectroscopy (EIS) is commonly used for label-free detection of analytes such as proteins, DNA and peptides. EIS based immunosensors have high sensitivity and are relatively simple to operate in comparison to other immunological methods that involve optical or piezoelectric instrumentation. EIS-based immune sensors typically utilize antigen–antibody immune layers that are very thin and have low electric permittivity. Molecular binding is detected by the interruption of the faradaic current at the electrode, which generates an impedance signal [31, 32, 47]. In the impedance system, we measure the change in impedance at the electrode interface at the double layer and in solution. EIS can transduce these electrode boundary phenomena to electric signals.
In addition, we developed a highly sensitive immunosensor for quantitative detection of various proteins using wafer-scale nanowell array (NWA) electrodes. Wafer scale fabrication methods have low throughput and are limited to small areas [33–36]. Therefore, in this study, we used a krypton-fluoride (KrF) stepper semiconductor process with a wavelength of 248 nm for the fabrication of NWA electrodes on 6-inch wafer. This wafer-scale fabrication method of nano-patterns has rapid, high-throughput and is highly reliable for the fabrication of NWA biosensor.
2 Nanoengineering of NWA structure
2.1 3D Au nanobox arrays
2.2 Double oxide deposition and etching (DODE) lithography
High aspect-ratio nanostructure fabrication techniques have been developed by Morton et al. [54–57]. Involved nanomold imprinting to transfer patterns on a substrate, sacrificial layers deposition for bulk silicon substrate etching and polymer replication processes, which cause the nanostructure manufacturing process being difficult, time-consuming, and expensive. However, the DODE technique can simply fabricate wafer-scale high aspect-ratio nanostructures with nanoscale patterns by isotropic oxide deposition and anisotropic oxide etching processes without e-beam lithography or nanoimprint lithography processes. In addition, the DODE method does not use a polymer resin as a masking layer in the dry etching process so the nanostructures fabricated using the DODE process can be used as a polymer-free nanotemplate for various nanoelectronic applications.
2.3 Hydrogel NWA electrodes on Au substrate
2.4 Wafer-scale fabrication of NWA
It was confirmed that each NWA structure had a diameter of 500 nm with interspacing of 200 nm and a height of 200 nm. The NWA electrode, the counter electrode and reference electrode were combined into the electrode holder. The liquid reagents were treated to the electrode and electrochemical analysis was performed.
3 Biomedical nanoengineering platform
3.1 Enhancement of integrated electric nanobiosensor utilizing NWA geometry
To demonstrate its usefulness, we fabricated a NW array-based gene assay for EC sensing, which has attracted a great deal of attention due to its high sensitivity, low cost, and high portability [62–70]. Figure 1b displays an illustration of our experimental electrode, composed of eight Au pads. The thickness of the resist layer was varied between 150 and 200 nm. And, each NW was 50 nm in diameter and separated by 500 nm. The atomic force microscope (AFM) picture, in Fig. 1c, illustrates that arrayed 50 nm NWs were successfully fabricated. Inside the NWs, probe ssDNA molecules were immobilized. The diameter of each streptavidin molecule used in this study was approximately 10 nm . The diameter of each streptavidin molecule used in this study was approximately 10 nm . Therefore, the maximum allowed number of the streptavidin–biotinylated DNA within a 50 nm NW is less than 3. The actual number should be much less than 3, since the actual diameter of the exposed gold dot is smaller than 50 nm due to the cutting angle of the fabricated resist layer. To directly observe whether the ssDNA molecules could be immobilized within the NWs, we performed AFM measurements. Initially, we attempted to obtain good AFM pictures on 50 nm NW arrays, but failed (due to the NW size, which might be too small and deep for an AFM cantilever to enter the holes). Instead, we performed AFM measurements on a 200 nm NWAs. Figure 1d shows an AFM picture of a NWA electrode subsequent to the ssDNA immobilization process. There are several bright spots inside each NWA, which serves as indirect confirmation that several probe DNA/streptavidin complexes could be attached inside the NWA. There are also bright spots on the resist layer, but these DNA/streptavidin complexes would not contribute to the EC signals. Although this AFM picture was taken on the 200 nm NW array, it demonstrates that the NW geometry concept, proposed in Fig. 1c, could work for our 50 nm NW array.
3.2 Electrochemical nanobiosensor: DNA
3.3 Electrochemical nanobiosensor: functional lipid vesicle
3.4 Electrochemical nanobiosensor: Immunosensor
We determined the charge transfer resistance (Rct) of the anti-STIP-1 antibody layer to be 36.96 kΩ. After antigen treatment, the Rct was determined to be 48.39, 57.26, 62.98, 68.26, 75.13 and 80.70 kΩ, for 10 pg/mL, 100 pg/mL, 1 ng/mL, 10 ng/mL, 100 ng/mL, and 1000 ng/mL STIP-1, respectively, (Fig. 15a). To compare the sensitivity of the NWA biosensor with that of other electrode-based biosensors, the same treatments were applied to a bare electrode without NWA. For the bare electrode (Fig. 15b), the base Rct was 52.52 k Ω, which represents an increase of 42% over the base Rct of the NWA electrode. The Rct values for 10 pg/mL, 100 pg/mL, 1 ng/mL, 10 ng/mL, 100 ng/mL, and 1000 ng/mL STIP-1 were 54.22, 55.34, 59.94, 61.64, 65.43 and, 65.83 kΩ, respectively.
The differences in the impedance spectra for each electrode before and after STIP-1 addition are shown in Fig. 15c. For the bare electrodes, the impedance at low frequencies interfered with the measurement of the limit of detection (LOD). This low signal-to-noise ratio (S/N ratio) was caused by non-specific binding of STIP-1. It is reasonable to estimate an LOD of 1 ng/mL for the bare electrodes. The LOD was estimated to be 10 pg/mL or less for the NWA electrodes, which suggests a 100-fold improvement in the LOD when using EIS with NWA electrode. The sensitivity of the NWA impedemetric immunosensor was better for each analyte concentration tested when compared the sensitivity of the bare electrode sensor. Therefore, bare electrode has a larger area (4 mm × 2 mm) than NWA electrode (1.75 mm2). Basically, electrodes with a larger area provide the sensor with a better S/N ratio because number of binding site is increased. However, in this study NWA’s S/N ratio was higher than the bare electrode although NWA had a smaller area than the bare electrode. The signal for the negative control (without antibody) was not significant, which further supports our finding of high specificity for the NWA electrodes.
Based on these results, the electrochemical impedemetric immune-sensors using NWA electrodes can be applied for label-free detection, with low levels of non-specific binding. Because NWA electrodes are optimal for the selective docking of single molecules, they reduce non-specific binding and enhance electrochemical responses. Therefore, NWA has high sensitivity and selectivity as well as very low LOD.
As previously mentioned, we applied modified randles circuit to fit our experimental data. In the Bode plot, there was a change of the magnitude of the impedance (|Z|) at the low frequency (< 100 Hz). Moreover, the signal in this region was dominated by the dielectric behavior of the electrode. For the high frequency region (> 104 Hz), the resistive region appears and resistance in the solution determines the signal [78–81]. In low frequency region, the magnitude of impedance increases with increasing STIP-1 concentrations, but the impedance change is negligible in high frequency region. The Rct increased by 66.77% for 1000 ng/mL STIP-1, over the value obtained with 10 pg/mL STIP-1. In contrast, the Rs increased by only 8.78%, which indicates that Rs does not reflect the STIP-1 concentration. This means that higher concentrations of STIP-1 cause increased binding of STIP-1 to the anti-STIP-1 antibody layer, which creates a denser and thicker electrical isolation layer [82–87]. These properties were observed in the low frequency region by EIS. Changes of the double layer properties at lower frequency would be more dramatic than the changes in the solution resistance at the higher frequency region.
NWA effectiveness compared with conventional ELISA method
Direct redox reaction
Limit of detection
0.2 × 10−2 U/L (μmol/L)
0.13 × 10−4 U/L (μmol/L)
200 μL per test
5 μL per test
Narrow (10 ng–1 μg/mL)
Wide(10 pg–10 μg/mL)
In addition, a highly sensitive electrochemical impedimetric immunosensor based on wafer-scaled NWA was successfully applied for quantitative detection of various proteins. At low frequencies, binding of protein to the immuno-affinity layer on the NWA electrodes resulted in large changes in impedance, and this effect was not observed at high frequencies. Thus, double-layer properties are more useful than solution resistance for qualifying protein. From the NWA electrodes, we calculated the change in Rct and estimated LOD for specific protein as stress induced phosphoprotein-1 (STIP-1) to be 10 pg/mL which represents a ≥ 100-fold improvement over the bare electrodes which are an electrode without NWA. Electrochemical impedimetric NWA biosensor can also be applied for label-free detection without non-specific binding by means of the selective docking of immobilized antibodies. These properties provide high sensitivity and selectivity for wafer scaled NWA sensor. The results also show the advantages of using NWA over large areas to improve performance and decrease costs. We anticipate that the in vitro nanobiomedical device will lead the way for the realization of digitized nanomedicine at the molecular level.
SJ and JL helped writing the manuscript. HC and JK and HL guided manuscript preparation. All authors read and approved the final manuscript.
The authors declare that they have no competing interests.
Ethics approval and consent to participate
This research was supported by the Basic Science Program of the National Research Foundation of Korea (NRF) funded by the Ministry of Education, Science and Technology (2014–052607).
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