**Plasmonic Sensors: A New Frontier in Nanotechnology**

Editors

**Samir Kumar Sungkyu Seo**

MDPI • Basel • Beijing • Wuhan • Barcelona • Belgrade • Manchester • Tokyo • Cluj • Tianjin

*Editors* Samir Kumar Korea University Sejong Republic of Korea

Sungkyu Seo Korea University Sejong Republic of Korea

*Editorial Office* MDPI St. Alban-Anlage 66 4052 Basel, Switzerland

This is a reprint of articles from the Special Issue published online in the open access journal *Biosensors* (ISSN 2079-6374) (available at: https://www.mdpi.com/journal/biosensors/special issues/plasm sensors).

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## **Contents**

## **About the Editors**

#### **Samir Kumar**

Samir Kumar, Ph.D., is a research professor (since 2021) at the Nano Biophotonics Lab, Korea University (Sejong, Korea). His previous academic positions were at Kyoto University (Japan, 2018–2021). He received his M.Sc. in physics from the University of Kalyani, India, in 2009, and his M.Tech. from the National Institute of Technology Hamirpur, India, in 2012. He received his Ph.D. in physics from the Indian Institute of Technology Delhi, India, in 2017. His research interests include designing, fabricating, and characterizing nano- and micro-electromechanical systems for sensor and actuator applications. He has authored more than 50 publications in scientific journals and proceedings.

#### **Sungkyu Seo**

Sungkyu Seo, Ph.D., is a professor (since 2009) in the Department of Electronics and Information Engineering at Korea University (Sejong, Korea). His previous academic positions were at UCLA (C.A., USA, 2008-2009). He is also an Adjunct Professor (since 2017) at the Department of Laboratory Medicine, Korea University (Guro Hospital, Seoul, Korea), and BK21 Plus Team Director (since 2014) at ICT Convergence Technology for Health & Safety, Korea University (Sejong, Korea). He received his B.S. degree in 1998 from Korea University (Seoul, Korea) and his Ph.D. degree in 2007 from Texas A&M University (TX, USA), all in electrical engineering. His previous academic positions were at UCLA (C.A., USA, 2008-2009). His research interests include bio photonics, nano optics, and MEMS/NEMS. He has published about 100 papers in scientific journals and proceedings.

## *Editorial* **Plasmonic Sensors: A New Frontier in Nanotechnology**

**Samir Kumar \* and Sungkyu Seo \***

Department of Electronics and Information Engineering, Korea University, Sejong 30019, Republic of Korea **\*** Correspondence<EMAIL_ADDRESS>(S.K<EMAIL_ADDRESS>(S.S.)

Plasmonics is the study of surface plasmons formed by the interaction of incident light with electrons to form a surface-bound electromagnetic wave [1]. Plasmons can configure light to nanoscale volumes, making them attractive for developing new technologies with a wide range of applications [2]. The intense fields created by plasmons can dramatically enhance various light–matter interactions, such as transmission [3], Raman scattering [4], and fluorescence [5], and can also lead to significant localized heating in metallic nanostructures, known as thermoplasmonics [6,7].

Surface plasmons are formed when the oscillation of free electrons on the surface of a metallic nanoparticle resonates with the incident light [8]. The resonance frequency of surface plasmons depends on the size, shape, and composition of the metallic nanoparticle, as well as on the dielectric properties of the surrounding medium [9]. Controlling the plasmon resonance frequency and tailoring the plasmon properties for specific applications is made possible by tuning these parameters. Applications of plasmonics can be found in nanotechnology, biophotonics, sensing, biochemistry, and medicine [10]. However, plasmonics has proven particularly useful in the development of biosensors [11]. The Scopus database contains 37,586 articles related to biosensors that have been published since 2018 (as of February 2023). Twenty-five percent of these studies contain plasmonicsrelated keywords.

Biosensors are devices that use a variety of physical, chemical, optical, electrochemical, and thermal processes to convert biological interactions into valuable information [12,13]. A biosensor has two components: a transducer and a bioreceptor. Bioreceptors are biomolecules that recognize the target analyte, whereas transducers convert this recognition into measurable signals [14]. Electrochemical and electrical transducers are at the forefront of revolutionizing biosensor technology. However, these systems are limited by their dependence on oxidizing/reducing agents [15]. Optical biosensors based on optical transducers play a prominent role in developing advanced biosensors due to their high detection accuracy and cable-free design [16]. Optical biosensors are capable of multiplex detection and remote sensing. Sensors based on absorption, surface plasmon resonance (SPR), and photoluminescence provide a next-generation optical biosensor platform [17–19]. There are several optical biosensing techniques which utilize plasmons, including surfaceenhanced Raman spectroscopy (SERS), surface-enhanced fluorescence spectroscopy (SEFS), and surface-enhanced infrared absorption spectroscopy (SEIRAS) [20–22].

This *Biosensors* Special Issue focuses on the theory and fabrication of plasmonic nanostructures; patterned surfaces; and devices for SPR, SEFS, and SERS-based biosensors. This Issue contains five research articles, five review articles, and one perspective article covering various topics related to fabrication and recent developments in plasmonic biosensors.

Gaur et al. investigated the interplay between SPR and lossy mode resonance (LMR) in this Special Issue. They compared these two types of sensing techniques in terms of sensitivity, detection accuracy, and figure of merit [23]. The SPR effect is a complicated physical phenomenon that occurs when light strikes a conductive layer of noble metal. This phenomenon occurs at the boundary between a medium with a low refractive index (buffer) and a medium with a high refractive index (sensor glass surface) [24]. In most

**Citation:** Kumar, S.; Seo, S. Plasmonic Sensors: A New Frontier in Nanotechnology. *Biosensors* **2023**, *13*, 385. https://doi.org/10.3390/ bios13030385

Received: 10 March 2023 Accepted: 13 March 2023 Published: 15 March 2023

**Copyright:** © 2023 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

plasmonic biosensors, the SPR effect detects the binding of biomolecules, such as proteins and DNA, to metallic nanoparticle surfaces. An interaction between a biomolecule and a nanoparticle changes the plasmon resonance frequency, which can be measured with a spectrophotometer. Unlike conventional biosensors, SPR-based biosensors offer several advantages, including label-free detection, high sensitivity, and real-time monitoring. In addition to SPR, LMR can also be used for optical sensing. LMR is a phenomenon that occurs when light is coupled into an optical waveguide and interacts with a lossy resonant mode of that waveguide [25]. The interaction between the light and the resonant mode results in a sharp dip in the transmission or reflection spectrum of the waveguide at the resonant wavelength. This effect is commonly observed in optical fibers, planar waveguides, and other structures supporting guided modes. LMR can be used in various applications, including sensing, filtering, and lasing. By detecting changes in the resonant wavelength, LMR sensors can measure changes in the refractive index, temperature, or other physical parameters in the waveguide environment [26]. The geometric configuration and the material supporting the SPR and LMR are key factors determining its performance. The authors investigated the properties of the bilayer (ITO + Ag structures) and trilayer (ITO + Ag + ITO structures) fiber-optic probes based on ITO, which enables simultaneous excitation of SPR and LMR. They found that LMR showed better sensitivity than SPR in the bilayer configuration. Moreover, as the thickness of the Ag layer increases, SPR depression becomes insensitive, and only LMR depression can be used for detection. Taking advantage of this property, a self-referencing sensor using the SPR depression as a reference point was proposed. In the three-layer configuration, they found that the first LMR depression in the visible region was less sensitive than the SPR depression when the thickness of the outermost ITO was varied. Conversely, when the thickness of Ag was varied, the resonance wavelength of the SPR depression shifted to the shorter wavelength side; however, the resonance wavelength of the LMR depression shifted to the longer wavelength side. The proposed sensor had a high sensitivity of 14 μm/RIU, good detection accuracy, and a Q-factor.

Label-free detection of biomolecules by localized surface plasmon resonance (LSPR) has excellent potential for point-of-care (POC) testing. SPR occurs at the metal–dielectric interface and is sensitive to changes in the refractive index of the dielectric layer. In contrast, LSPR occurs at the surface of metallic nanoparticles and is sensitive to changes in the local environment around the nanoparticle. LSPR results from the confinement of surface plasmons in nanoparticles (with a size comparable to the wavelength of the light used to excite the plasmons) [27]. In developing LSPR-based POC devices, a key challenge is to produce large-scale LSPR substrates that are reproducible and have high throughput. In a study included in this Special Issue, Kim et al. fabricated wafer-scale LSPR substrates using reproducible high-throughput techniques such as nanoimprint lithography, wet etching, and the glancing angle deposition (GLAD) technique [28]. First, they fabricated hard masks of SiO2 nanodots on a transparent sapphire wafer using nanoimprint lithography, which was anisotropically etched with a solution of H2SO4 and H3PO4, resulting in a patterned sapphire substrate (PSS). Finally, they fabricated an LSPR substrate with Au on the PSS by GLAD and used it to detect biomolecule binding processes without labeling. The GLAD technique allowed for the formation of Au nanostructures on the PSS due to the PSS's three-dimensional structure (triangular–pyramidal). The Au nanostructures formed by the GLAD technique showed a red-shifted LSPR peak compared to those formed by vertical deposition. This suggests that the LSPR properties can be controlled by the shape of the PSS and deposition conditions, such as the PSS's angle and rotation.

GLAD is a physical vapor deposition technique used to fabricate complex nanostructures and thin films with controlled porosity and morphology [29,30]. In this technique, the material is deposited obliquely onto a substrate at a low angle of incidence, typically between 70◦ and 87◦, resulting in highly anisotropic structures with directional growth [31]. Nanostructures fabricated by the GLAD technique exhibit high sensitivity, enhanced optical and catalytic properties, periodicity, and controlled morphology, which makes them

attractive for sensory applications [32,33]. In another study in this Special Issue, Yadav et al. provided a detailed overview of recent advances in various nanostructures fabricated via GLAD and their applications in the biomedical field [34]. In addition to discussing the various ways to fabricate nanostructures using the GLAD technique, they also considered the assembly configuration, the effects of different growth parameters, and the advantages of GLAD-based nanostructures over conventional nanoparticles and substrates. In addition, the authors highlighted several advantages of GLAD compared to other chemical deposition processes. For example, GLAD does not require precursor materials, which provides safety from toxic precursors and by-products. The chemical composition and thickness in this deposition can be controlled at the atomic level. Higher temperatures are not required, so this technique can produce heat-sensitive substrates.

Of all the metals reported for plasmonic sensors, Ag and Au are considered the most useful plasmonic materials. They show the naturally lowest ohmic losses at optical frequencies due to their strong plasmon resonances in the visible range, their biocompatibility, and their stability [35,36]. In this Special Issue, Gahlaut et al. presented highlights of recent developments in the field of nanostructured Ag substrates for plasmonic sensing, with a focus on SPR and SERS, over the past decade [37]. They mainly focused on the chemical methods of solution phase synthesis and physical methods such as vapor phase deposition, GLAD, and lithographic techniques for synthesizing Ag nanostructures ranging from nanoparticles to nanocubes, nanotriangles, nanorods, and nanowires. In addition, the authors also discuss the latest spectroscopic techniques, focusing on plasmonic enhancement of biosensing methods such as SPR/LSPR, SERS, SEF, and SEIRAS using Ag as plasmonic material.

Metallic nanoparticles, especially Au, have been widely used in plasmonics due to their excellent plasmonic properties. However, the discovery of two-dimensional (2D) nanomaterials, such as MXenes, has revealed new possibilities for plasmonic applications [38,39]. MXenes are a class of 2D transition metal carbides, nitrides, and carbonitrides that have recently shown promise as materials for plasmonics [40]. They have excellent electrical conductivity, mechanical strength, and chemical stability, which makes them attractive for various applications. One of the advantages of MXenes over metallic nanoparticles is their tunability. The plasmonic properties of MXenes can be tuned in the mid-IR to THz regions of the spectrum, and their optical constants (such as the dielectric constant and refractive index) can be tuned by controlling their chemical composition, surface functionalization, number of layers, and morphology [41]. For example, the wavelength of the plasmon resonance of MXenes can be tuned by changing the thickness or the type of functional group on the surface [42]. MXenes have shown great potential in various plasmonic applications, including biosensing, imaging, and photovoltaics. They can also be integrated with other materials, such as polymers or nanoparticles, to create hybrid plasmonic systems with enhanced optical properties. Some researchers have coined the term "smart MXene" to describe MXene-based hybrid materials that exhibit unique properties related to their applications [43,44]. Smart MXene quantum dots (SMQDs), a new and rapidly emerging class of nanomaterials, are tantalizing candidates for SPR biosensor development given their intriguing optical properties, which include light absorption, photoluminescence, and electrochemiluminescence [43]. In their study, contained in this Special Issue, Mousavi et al. discussed recent advances in ultra-sensitive SPR nano biosensors based on SMQDs [45]. They began with an introduction to SPR and SPR biosensors. Then, they explained and discussed the different types of SMQDs and SPR biosensors based on SMQDs. Finally, they presented the useful characteristics of SMQDs for developing SPR biosensors and their biomedical applications, and discussed the current limitations of these biosensors. Overall, MXene plasmonics is a rapidly growing field with promising prospects for various applications, and it is expected to impact future plasmonic sensors' development significantly.

In this Special Issue, Park et al. discussed innovative advances in biosensors led by DNA nanotechnology [46]. Nanotechnology has made considerable progress in the last decade, allowing us to overcome the limitations of using DNA as the sole form of genetic material and thus develop a new method for constructing biosensors [47]. DNA aptamers can bind to a specific target with a strong affinity. DNA aptamers and enzymes can be produced via the systematic evolution of ligands by exponential enrichment (SELEX), and can bind to specific targets [48]. Aptamers are short, single-stranded DNA or RNA molecules that can bind to specific targets with high affinity and selectivity, similar to antibodies. Various research groups have presented plasmonic biosensing strategies based on aptamers [49]. In aptamer-based plasmonic sensors, the aptamer is immobilized on a metal surface, such as gold or silver. The binding event between the aptamer and the target molecule causes a change in the optical properties of the metal surface, which can be detected by various optical methods. The optical signal can be measured by various methods, e.g., fluorescence, absorbance, and SPR [49]. This review highlighted recent advances in the development of nucleic acids for the purpose of constructing plasmonic biosensors through studies published over the past five years. The authors discussed applying simple aptamers and origami-shaped structures in SPR/LSPR, SERS, and SEF sensing. They discussed several aspects that should be considered to improve the sensing process, including the nonspecific adsorption of unwanted molecules and the fabrication of homogeneous plasmon surfaces. The combination of aptamers and plasmonics has great potential to improve the sensitivity and selectivity of biosensors and enable the development of new diagnostic tools for various applications.

Plasmonics has made significant advances in the field of SERS, a highly effective analytical technique for identifying and detecting molecules [33,50]. SERS significantly amplifies Raman scattering signals when molecules are near plasmonic surfaces [4]. This enhancement is due to the excitation of LSP in metallic nanostructures, which generate strong electromagnetic fields that interact with molecules. SERS is a surface-sensitive technique that can detect analytes even at the level of single molecules [51]. In this Special Issue, Beeram et al. provide an excellent overview of recent trends in SERS-based plasmonic sensors for disease diagnostics, biomolecule detection, and machine learning techniques [52]. The authors review the work of the past decade and use simplified language to address the needs of an interdisciplinary audience. In the first section, the authors discuss the need for plasmonic sensors in biology and the advantages of SERS over existing technologies. Next, the authors discuss the use of SERS-based biosensors for disease diagnosis, focusing on cancer and respiratory disease detection, such as the recent detection of severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2). They then discuss advances in the detection of microorganisms, such as bacteria, with particular emphasis on plasmonic sensors for the detection of biohazardous materials. Finally, the authors discuss machine learning techniques for identifying, classifying, and quantifying biological signals using SERS. They discuss the various machine learning models that have been developed for the trace detection, signal variation, quantification, and identification of SERS signals, including principal component analysis, support vector machines, partial least squares, decision trees, and convolutional neural networks. Considering that SERS is a complex system with many variables, machine learning techniques can help to identify patterns that no experts can identify.

The development of plasmonic biosensors has revolutionized not only the field of biomedical research and diagnostics, but also food safety and environmental research [53]. This Special Issue also contains research on biosensors for food and environmental contaminants such as mercury [54], tetrabromobisphenol A (TBBPA) [55], and thiamethoxam [56]. Kim et al. developed an Au nanoparticle (AuNP)-based immunochromatographic lateral flow assay (ICA) using the light scattering phenomenon of nanoparticles to rapidly detect mercury in rice, with a detection limit of 20 ng g−<sup>1</sup> and a cut-off value of 500 ng g−<sup>1</sup> [54]. The proposed ICA strip can be used to qualitatively determine mercury in rice, and the results agree with conventional instrumental methods such as inductively coupled plasma mass spectrometry. Based on aptamer technology, Yue et al. developed a biosensor system to detect thiamethoxam residues [56]. They developed a colorimetric sensor based on AuNPs to

detect thiamethoxam in tea leaves. To screen ssDNA aptamers that bind specifically to thiamethoxam, they used graphene oxide SELEX (GO-SELEX) technology. They increased the amount of GO in the screening process to increase the screening pressure. The researchers applied the aptasensor system to the actual detection of samples. They achieved recoveries of 96.94% to 105.86% and RSD values of 0.41% to 3.76%, indicating that the aptasensor can be used for the rapid and sensitive detection of thiamethoxam residues. In another work, Dai et al. developed a hydrophobic Cu–Ag chip for SERS detection of TBBPA [57]. TBBPA is one of the world's most widely used brominated flame retardants, but its production and use can affect the environment and human health. Recently, there has been much interest in hydrophobic materials due to their wide range of applications in metal corrosion protection, self-cleaning, oil–water separation, and SERS [58,59]. According to Dai et al., the Cu–Ag chip's hydrophobicity increases the substrate's affinity for TBBPA, allowing TBBPA to approach the surface of the SERS substrate and subsequently combine with it via Ag–Br interactions. First, the researchers fabricated the hydrophobic copper-coated fabric with an ordered micro-nanostructure. Then, they constructed a hydrophobic Cu–Ag chip by introducing Ag onto the Cu surface through an exchange reaction. Combined with ultrasound-assisted extraction (UAE), they succeeded in quantitatively detecting TBBPA in electronics. After UAE, they extracted and determined the TBBPA content in the electronics with a detection limit of 2.0 mg kg−<sup>1</sup> (0.01 mg L−<sup>1</sup> for the TBBPA solution). SERS offers several advantages over conventional methods such as high-performance liquid chromatography (HPLC), including speed, convenience, and sensitivity.

The COVID-19 pandemic, caused by the novel coronavirus SARS-CoV-2, has resulted in a global public health crisis, affecting millions of people worldwide. Rapid, accurate, and sensitive diagnostic methods are critical for optimal patient care. Recently, SERS has emerged as a promising point-of-care testing technique for detecting various analytes, including viruses [60]. Unlike conventional methods such as polymerase chain reaction, which is limited to the analysis of genetic material, SERS offers high specificity and can detect a wide range of analytes. Furthermore, flexible SERS substrates are becoming increasingly important in practical application research due to their robustness and versatility [61]. In this Special Issue, Mousavi et al. evaluated the importance of the flexible SERS substrate for detecting SARS-CoV-2 [62]. They provided an overview of the flexible SERS substrates used to detect different subtypes of the novel SARS-CoV-2 coronavirus and discussed how this method has the potential to become a point-of-care diagnostic tool. In addition to discussing recent advances in SERS-based COVID-19 detection, they discussed the principles of SERS and their amplification mechanisms, the detection of analytes, and the use of multiplex analysis to detect coronaviruses.

In summary, plasmonic biosensors represent a promising technology for the label-free and sensitive detection of biomolecules. The sensitivity, specificity, and label-free detection offered by plasmonic biosensors make them an attractive alternative to conventional detection methods. This Special Issue on "Plasmonic Biosensors", published in *Biosensors*, provides a comprehensive overview of recent developments and applications in the field of plasmonic biosensors. The research articles present their original work in developing plasmonic biosensors for detecting various biomolecules, including DNA. In addition, this Special Issue highlights the significant advances in the development of plasmonic biosensors and their potential applications in various fields. This Special Issue also highlights the importance of interdisciplinary research and collaboration in developing plasmonic biosensors. Researchers from diverse backgrounds, including physics, chemistry, and biology, have contributed to progress in this field.

In conclusion, this Special Issue on "Plasmonic Biosensors" provides an excellent platform for researchers to present their latest work and exchange ideas regarding the future directions of plasmonic biosensors. The articles in this Issue demonstrate the enormous potential of plasmonic biosensors for various applications, and we can expect to see further advances in this field. As researchers continue to explore the fundamental principles of plasmonics and SERS, we can expect more exciting developments in this field in the coming years.

**Funding:** This study was supported by the Basic Science Research Program (Grant#: 2020R1A2C1012109) through the National Research Foundation (NRF) of Korea. This work was also supported by the Korea Medical Device Development Fund grant funded by the Korean Government (the Ministry of Science and ICT, the Ministry of Trade, Industry, and Energy, the Ministry of Health & Welfare, the Ministry of Food and Drug Safety) (Grant#: RS-2020-KD000142) and the Korea Institute of Marine Science & Technology Promotion (KIMST) funded by the Ministry of Oceans and Fisheries, Korea (Grant#: 20210660).

**Acknowledgments:** The authors are grateful for the opportunity to serve as the guest editors of the Special Issue "Plasmonic Sensors: A New Frontier in Nanotechnology", as well as for the contribution of all the authors to this Special Issue. The dedicated work of the Special Issue Editor of *Biosensors* and the editorial and publishing staff of *Biosensors* is greatly appreciated.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

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## *Article* **An Interplay between Lossy Mode Resonance and Surface Plasmon Resonance and Their Sensing Applications**

**Deependra Singh Gaur 1,†, Ankit Purohit 1,†, Satyendra Kumar Mishra 2,\* and Akhilesh Kumar Mishra <sup>1</sup>**

**\*** Correspondence<EMAIL_ADDRESS>
† These authors contributed equally to this work.

**Abstract:** Conducting metal oxide (CMO) supports lossy mode resonance (LMR) at the CMO-dielectric interface, whereas surface plasmon resonance (SPR) occurs at the typical plasmonic metal-dielectric interface. The present study investigates these resonances in the bi-layer (ITO + Ag) and tri-layer (ITO + Ag + ITO) geometries in the Kretschmann configuration of excitation. It has been found that depending upon the layer thicknesses one resonance dominates the other. In particular, in the tri-layer configuration of ITO + Ag + ITO, the effect of the thickness variation of the sandwiched Ag layer is explored and a resonance, insensitive to the change in the sensing medium refractive index (RI), has been reported. Further, the two kinds of RI sensing probes and the supported resonances have been characterized and compared in terms of sensitivity, detection accuracy and figure of merit. These studies will not only be helpful in gaining a better understanding of underlying physics but may also lead to the realization of biochemical sensing devices with a wider spectral range.

**Keywords:** optical fiber sensor; surface plasmon resonance; lossy mode resonance; indium-tin oxide; silver; sensitivity; detection accuracy

#### **1. Introduction**

Surface plasmon resonance (SPR) is generated at metal and dielectric interface by transverse magnetic fields (TM) or p-polarized light [1]. It is impossible to excite the SPR mode by direct light due to the momentum mismatch between the SPR mode and incident light [2]. In order to efficiently excite these modes, we need a momentum matching scheme. Several such schemes have been proposed, e.g., passing light through a high RI prism, using a grating, etc. [3]. Further, there exist two coupling configurations for SPR excitation—Otto and Kretschmann [4,5]. Owing to its ease of implementation, the Kretschmann configuration is preferred often.

The growing field of SPR has attracted significant research attention over the years due to its wide range of applications, which includes nano-antennas [6], imaging [7], biosensing, and so on and so forth [8]. Several extensive theoretical and experimental studies have been conducted on SPR-based sensors in the past [9–13]. A sensor's novelty is determined by the particulars of the plasmonic material used and the design implemented. Waveguide-based sensors have attracted a lot of attention due to their industrial applications. These sensors use a plasmonic material deposited as a thin film around the waveguide (e.g., an optical fiber). These materials can be classified into three categories based on the resonances they support. The first class of materials is plasmonic materials, which support SPR and have a real permittivity that is negative and larger in magnitude than both the imaginary permittivity and the permittivity of the surrounding medium. In the second category of materials, the real part of the material permittivity is positive and greater than both its imaginary part and the permittivity of the surrounding medium. The LMR phenomenon is observed in this category of materials. The third class of materials also exists for which the

**Citation:** Gaur, D.S.; Purohit, A.; Mishra, S.K.; Mishra, A.K. An Interplay between Lossy Mode Resonance and Surface Plasmon Resonance and Their Sensing Applications. *Biosensors* **2022**, *12*, 721. https://doi.org/10.3390/ bios12090721

Received: 30 June 2022 Accepted: 2 September 2022 Published: 4 September 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

real part of the permittivity is close to zero and the imaginary part is large. Such a material supports long-range surface exciton-polariton [14]. The present work only focuses on the first and second classes of materials.

LMR results out of the coupling between lossy mode and evanescent wave at a particular thickness of the thin film [15]. Only a few studies are reported on the application of LMR to sensing because the selection of the appropriate material for the thin film is critical [16–19].

Different types of waveguide structures have been utilized to realize SPR and LMRbased sensors. In particular, plasmonic fiber grating based, U-shaped and D-shaped optical fiber-based biosensor are explored extensively [19–21]. The simultaneous generation of LMR and SPR on the same planer platform has also been reported [22].

Indium tin oxide (ITO) is one of the CMO materials that supports LMR. It is a transparent material with an optical band gap of 3.6 eV, which restricts band-to-band transitions. The electronic and optical properties of ITO can be tuned during fabrication, resulting in a significant variation in its characteristics [23]. This property can be used to shift the resonance wavelength of LMR. Unlike SPR, the excitation of LMR has the advantage of not requiring specific polarization for incident light. Additionally, it is possible to generate multiple dips in the transmission spectrum. LMR dips are usually found in the IR and UV regions but, with the proper optimization of thin-film and with the use of other materials, they can also be observed in the visible region [14]. The IR dip is observed due to oscillations in charge density along the metal-dielectric interface. In contrast, a charge density oscillation along the thickness of the metal film is responsible for UV dip [16].

LMR is also found suitable for sensing applications. The performance and novelty of the sensor are determined by the material and sensing probe design used. Because of the excellent characteristics of optical fibers, these are being used as substrates for depositing ITO thin films to constitute the sensor [24].

Our present study examines the characteristics of bi- and tri-layer fiber optic sensing probes based on ITO that enable simultaneous excitations of SPR and LMR both. In bi- and tri-layer geometries, we investigate ITO + Ag and ITO + Ag + ITO structures, respectively. Applied biosensing, chemical analysis, quality assurance of food, and wavelength filtering are some of the potential applications for the proposed sensing probe.

#### **2. The Model**

To generate LMR, the lossy mode must be coupled with the evanescent wave. At a particular angle or wavelength, the effective index of the evanescent wave matches with the effective index of the lossy mode. The effective RI of the evanescent wave is given by

$$n\_{eff} = n\_p \sin(\theta\_i) \tag{1}$$

where *np* is the RI of the substrate, and *θ<sup>i</sup>* is the incident angle of the light. This relationship shows that the RI of evanescent waves can be controlled by the incident angle of incident light and/or corresponding wavelength.

Figure 1 schematically shows the proposed sensing probe in the Kretschmann configuration. The probe consists of a multimode fiber with a core diameter of 400 μm and a numerical aperture of 0.22. The 1 cm fiber cladding has been removed from the fiber probe. On top of the unclad (exposed) core, layers of ITO and Ag have been considered.

At one end of the fiber, light from a polychromatic source is launched, and the spectrometer records the corresponding transmission spectrum at the other end. At a certain wavelength, called the resonance wavelength, the spectrum exhibits a minimum transmitted power. A change in the sensing medium (i.e., RI of the analyte) will alter the resonance wavelength. The sensitivity of the sensor is defined as the shift in resonance wavelength corresponding to the change in the RI of the analyte. Another important characterization parameter, figure of merit (FOM) is defined as the ratio of sensitivity to the full width at half maximum (FWHM) of the transmission dip. Additionally, detection accuracy (DA) is

another important parameter that measures the sharpness of the resonance. The following expression relates these parameters (sensitivity, FOM, and DA) with each other [25].

$$\text{FOM} = \frac{\text{Sensitivity}}{\text{FWHM}} = \text{Sensitivity} \times \text{DA} \tag{2}$$

where DA <sup>∝</sup> <sup>1</sup> FWHM.

Some materials are highly sensitive but have low FOM. In contrast, others have poor FOM. Hence, materials need to be carefully selected.

**Figure 1.** Schematic of the presented SPR setup (TH: tungsten halogen, 3D-MO: microscopic objective with 3D movement).

For multilayer structures, the transfer matrix method is used to calculate the transmission spectrum. Consider a kth layer of thickness dk, having complex RI nk, and dielectric coefficient *εk*. The transfer matrix for N layer system is expressed as

$$M = \prod\_{k=2}^{N} M\_k = \begin{bmatrix} M\_{11} & M\_{12} \\ M\_{21} & M\_{22} \end{bmatrix} = \begin{bmatrix} \cos(\beta\_k) & -i\sin(\beta\_k/q\_k) \\ -i\, q\_k \sin(\beta\_k) & \cos(\beta\_k) \end{bmatrix} \tag{3}$$

where *β<sup>k</sup>* and *qk* are defined as (2*πdk*/*λ*) *ε<sup>k</sup>* − *n*<sup>1</sup> <sup>2</sup>*sin*2*θ*<sup>1</sup> 1/2 and *ε<sup>k</sup>* − *n*<sup>1</sup> <sup>2</sup>*sin*2*θ*<sup>1</sup> 1/2/*εk*, respectively and *θ*<sup>1</sup> is the incident angle of the ray, while *λ* is the wavelength of the incident light. The reflection coefficient *rp* of p-polarized (TM polarized) incident wave through the film is expressed as:

$$r\_p = \frac{(M\_{11} + M\_{12}q\_N)q\_1 - (M\_{21} + M\_{22}q\_N)}{(M\_{11} + M\_{12}q\_N)q\_1 + (M\_{21} + M\_{22}q\_N)}\tag{4}$$

The reflectance, *R*, for TM polarized light is given as

$$\mathcal{R} = \left| r\_p \right|^2 \tag{5}$$

A detailed description of this matrix method is given elsewhere [6,13]. The rays launched within the well-defined range of angle would be guided and the range is given by *θ*<sup>1</sup> = *sin*−1(*ncl*/*n*1) to *θ*<sup>2</sup> = *π*/2.

The transmitted power at the output end of the fiber is given by

$$P\_{\rm trans} = \frac{\int\_{\theta\_1}^{\theta\_2} R\_p^{N\_{\rm ref}(\theta)} n\_1^2 \left(\sin\theta \cos\theta / \left(1 - n\_1^2 \cos^2\theta\right)^2\right) d\theta}{\int\_{\theta\_1}^{\theta\_2} n\_1^2 \left(\sin\theta \cos\theta / \left(1 - n\_1^2 \cos^2\theta\right)^2\right) d\theta} \tag{6}$$

where

$$N\_{ref}(\theta) = \frac{L}{D \tan \theta} \tag{7}$$

The number of reflections occurring in the sensing region is denoted by the Equation (7), where *L* is the length of the unclad region and *D* is the diameter of the fiber. The dielectric constant of the Ag and ITO layer is calculated by the Drude dispersive model expressed as

$$
\epsilon(\lambda) = \epsilon\_r + i\epsilon\_i = 1 - \frac{\lambda^2 \lambda\_c}{\lambda\_p^2 (\lambda\_c + i\lambda)} \tag{8}
$$

and

$$
\epsilon(\lambda) = \epsilon\_r + i\epsilon\_i = 3.8 - \frac{\lambda^2 \lambda\_c}{\lambda\_\mathcal{P}^2 (\lambda\_c + i\lambda)} \tag{9}
$$

respectively, where *λ<sup>p</sup>* and *λ<sup>c</sup>* is the wavelength corresponding to bulk plasma frequency and collision wavelength. In the case of Ag, *λ<sup>p</sup>* = 0.14541 μm and *λ<sup>c</sup>* = 17.6140 μm, whereas *λ<sup>p</sup>* = 0.56497 μm and *λ<sup>c</sup>* = 11.21076 μm for ITO. The Sellmeier equation has been used to determine the RI of the fiber core [13]. We have assumed that above dispersion relations are valid in the whole wavelength range of investigation.

In order to fabricate the sensing probe, we use multimoded plastic clad fibers. The cladding can be removed by a few centimeters (a length of 1 cm of cladding is suitable for sensing applications) and then cleaned in a vacuum chamber using ion plasma bombardment and acetone. The unclad port of the fiber can be coated with metal or ITO after cleaning. Depending on the deposition techniques we have used, the uniformity of the films will vary. In order for the sensing probe to work correctly, the film uniformity must be good. A high-quality film can be achieved using sputtering and e-beam evaporation. The probe can be characterized by injecting light through one of the fiber faces and analyzing its sensing performance using a spectrometer at the other end of the fiber [26–30].

#### **3. Results**

The following two cases have been discussed in this section- in the first case, the ITO layer is deposited directly on the fiber core followed by the Ag layer (bi-layer sensing probe), and in the second case, an additional layer of the ITO is deposited over the Ag (tri-layer sensing probe).

#### *3.1. Bi-Layer Configuration (ITO + Ag)*

In this section, we numerically investigate a bilayer configuration of ITO + Ag coated fiber probe. In the first round of the simulation, the thickness of the ITO layer is fixed at 80 nm, whereas the thickness of the Ag layer is varied from 10 nm to 60 nm.

In the transmission spectrum, we observe two resonance dips for probe configuration with 80 nm ITO and 10 nm Ag for various values of analyte RIs, as shown in Figure 2a. Since the real part of the dielectric constant of the ITO is positive and larger than its imaginary part, at the lower wavelength region the condition of LMR generation is supported. Experimental evidence supports the LMR generation at short wavelengths and SPR excitation at long wavelengths [18]. The resonance dip in the visible region is caused by the LMR phenomena, while the second dip is the result of SPR. We would like to note here that the dielectric constant of ITO remains positive only for shorter wavelengths while at longer wavelengths it becomes negative (see Equation (9)). Therefore, for longer wavelengths, we see SPR resonance while in shorter wavelengths the probe supports LMR. These observations are well documented in the literature [14]. We would also like to note that for SPR excitation we require an interface of materials with opposite permittivity (one positive and other negative). An SPR dip can be observed even for a thin layer of Ag (i.e., 10 nm). Figure 2b illustrates that the SPR dip is more pronounced at a larger thickness of Ag. The thickness of the Ag layer, therefore, plays a critical role in the development of the SPR

dip. In addition, as Ag thickness increases, the SPR dip becomes less sensitive to the RI variations in the analyte, while the LMR dip's sensitivity increases.

**Figure 2.** Normalized transmission spectra of sensing probe with (**a**) ITO (80 nm) + Ag (10 nm) and (**b**) for various thicknesses of Ag for 80 nm ITO and analyte RI *ns* = 1.33.

The thickness of the ITO layer, however, significantly impacts the development of SPR and LMR resonances. As shown in Figure 3a, both resonances emerge with increasing thickness of the Ag layer for the 50 nm thick ITO layer. If the thickness of the ITO layer is less than 50 nm, SPR and LMR resonance dips still grow with the thickness of the Ag layer as shown in Figure 3b. Both of these cases show that SPR dip is insensitive to the RI of the sensing medium (transmitted power plot not shown).

**Figure 3.** Normalized transmission spectra of sensing probe for various thicknesses of the Ag and (**a**) ITO (50 nm) (**b**) ITO (30 nm) and RI of the surrounding medium is 1.33.

In Figure 4a, we have plotted the LMR and SPR dip sensitivity against the thickness of the Ag layer for 80 nm thick ITO film. Both resonances are initially sensitive to changes in RI, but their sensitivities are drastically influenced by the thickness of the Ag layer. Nevertheless, the LMR sensitivity is improved as a result of the thicker Ag layer, while the SPR sensitivity is reduced as depicted in Figure 4a. At a very large thickness of the Ag layer, the SPR dip becomes insensitive to any RI variation of the analyte. This insensitive dip can be used as a reference point for characterizing the sensor's performance. The variation in the resonance wavelength of the LMR and SPR dip is also shown in Figure 4b. As Ag thickness is increased, the SPR resonance wavelength shifts slowly towards the smaller wavelength side. Also, shown is the transmitted power at the resonance wavelengths in Figure 4c. From the figure, we observe that the wavelength that corresponds to the LMR transmitted power minimum decreases to a minimum at a particular thickness of the Ag layer, and then increases. In contrast, the transmitted power for SPR dip is shifted toward the lower wavelength side with an increasing layer thickness of Ag. Previously a similar study is reported in [13], where resonance dip, observed in the visible region, was found useful for sensing applications and the second dip that appeared in the NIR region was

insensitive to the surrounding RI, but we will not focus on this insensitive SPR dip here, since it is already detailed nicely in the literature [13].

**Figure 4.** (**a**) Sensitivity (**b**) resonance wavelength and (**c**) minima of the transmission spectra at resonance wavelength of the LMR and SPR modes with the variation of Ag layer when ITO (80 nm) and RIs of the analyte are 1.33 and 1.34.

Further, we have studied the sensitivity and DA of both the resonances as a function of the analyte RI for different values of the Ag layer thickness as shown in Figures 5 and 6. The sensitivities of both the modes (LMR and SPR) increase with an increasing RI of the sensing medium as shown in Figure 5a,b. These plots also suggest that LMR dip is far more sensitive as compared to SPR dip. Figure 6 shows the corresponding DA variation as a function of the thickness of the Ag layer. DA for LMR dip (Figure 6a) is also relatively large as compared to that for SPR dip (Figure 6b). Also, note the opposite trends in Figures 5 and 6 with variations in the Ag layer thickness. Hence there is a trade-off between optimum values of sensitivity and DA for designing the bi-layer sensing probe.

**Figure 5.** Sensitivity of the (**a**) LMR and (**b**) SPR dip with the variation of RI of the sensing medium for different thicknesses of Ag layer and 80 nm thickness of ITO layer.

**Figure 6.** Detection accuracy of the (**a**) LMR and (**b**) SPR dip with the variation of RI of the sensing medium for different thicknesses of the Ag and 80 nm thickness of the ITO layer.

#### *3.2. Tri-Layer Configuration*

In this section, we investigate the tri-layer configuration (ITO + Ag + ITO). In the following tri-layer configuration, SPR is found to be more sensitive than the LMR dip. The following sub-sections analyze two important cases.

#### 3.2.1. ITO (10 nm) + Ag (10 nm) + ITO (X nm)

This configuration examines the resonance characteristics of a tri-layer ITO + Ag + ITO coated fiber sensor, where a 10 nm layer of ITO is considered on the fiber core that is followed by a 10 nm thick layer of Ag, and then a third layer of ITO with varying thickness. Two resonance dips appear in the transmission spectra when the third layer of ITO is 10 nm thick, as shown in Figure 7a. The resonance dip in the visible range corresponds to LMR and the second dip to SPR. Increasing the thickness of the third ITO layer (40 nm) causes a new resonance dip to appear in the near-infrared region, known as LMR. The newly developed LMR dip (middle dip) is not affected by the change in RI of the surrounding medium, as shown in Figure 7b. Compared to the first LMR dip, the SPR dip (third dip) shows much better sensitivity.

**Figure 7.** Normalized transmission spectra of sensing probe with (**a**) ITO (10 nm) + Ag (10 nm) + ITO (10 nm) and (**b**) ITO (10 nm) + Ag (10 nm) + ITO (40 nm).

The occurrence of the new LMR as a function of the thickness of the third ITO layer is shown in Figure 8a,b, where the sandwiched Ag layers are kept 10 nm and 20 nm thick, respectively. Figure 8a,b illustrates that increase in the thickness of the Ag layer from 10 nm to 20 nm shifts the resonance wavelength of the first LMR and SPR dips toward the longer wavelength side. However, the resonance wavelength of the insensitive LMR dip (central dip) remains nearly unchanged. Also, it seems that with the ITO layer thickness variation, the central dip merges with SPR dip. Alternatively, it also suggests a switch-over behavior between the two dips. This observation requires further exploration.

**Figure 8.** Normalized transmission spectra of sensing probe with (**a**) ITO (10 nm) + Ag (10 nm) + ITO (X nm) and (**b**) ITO (10 nm) + Ag (20 nm) + ITO (X nm).

As the thickness of the third ITO layer increases, the LMR dip's sensitivity decreases, whereas SPR's sensitivity increases as shown in Figure 9. The sensitivity variations of SPR and LMR dips in the tri-layer case are opposite to those in the bi-layer case (see Figure 4a).

**Figure 9.** Sensitivity of the First LMR and SPR dip with the variation of the third ITO layer when the first layer is ITO (10 nm) and the second layer is Ag (10 nm) and RI of the sensing medium varies from 1.33 to 1.36 RIU.

Additionally, the DAs for the first LMR and SPR dip are shown in Figure 10a,b, respectively, for varying thicknesses of the third ITO layer. The DA for LMR decreases with sensing medium RI while the opposite trend is observed for SPR dip in Figure 10b.

**Figure 10.** Detection accuracy of the (**a**) first LMR and (**b**) SPR dip with the variation of third ITO layer for 10 nm thick first layer ITO and 10 nm thick Ag layer.

Furthermore, the variation of the FOM with sensing medium RI for two dips is depicted in Figure 11a,b. The figure clearly shows the better performance of the SPR dip.

**Figure 11.** FOM of the (**a**) first LMR and (**b**) SPR dip with the variation of the third ITO layer thickness where the first layer is ITO (10 nm) and the second layer is Ag (10 nm).

3.2.2. (ITO (50) + Ag (X) + (ITO (50))

This section presents the normalized transmission spectrum for a tri-layer configuration with an Ag layer sandwiched between two ITO layers of thickness of 50 nm each. In this configuration too, two dips of LMR and one of SPR are observed, and the characteristics of these resonances are dependent on the thickness of the Ag layer, as shown in Figure 12a,b. Figure 12a shows transmittance variation with change in the sensing medium RI, while

in Figure 12b, the RI of the sensing medium is kept fixed at 1.33 RIU and the thickness of the Ag layer is varied. Figure 12a depicts that SPR dip is relatively more sensitive and central LMR dip is completely insensitive. Figure 13 indicates that on further increase in the thickness of the Ag layer, the resonance wavelength of SPR dips shifts toward the shorter wavelengths; however, the resonance wavelength of the first LMR dip shifts toward the longer wavelength. The LMR dip observed in the NIR region is insensitive to the analyte RI variations. Moreover, the increased thickness of the Ag layer tends to annihilate the SPR dip as shown in Figure 13.

**Figure 12.** Normalized transmission spectra of sensing probe with (**a**) ITO (50 nm) + Ag (10 nm) + ITO (50 nm) and (**b**) for varying thickness of the Ag layer when both ITO layers are 50 nm thick.

**Figure 13.** Normalized transmission spectra of sensing for varying thickness of the Ag layer when both ITO layers are 50 nm thick.

Further, we study the DA, sensitivity, and FOM of each resonance dip. We demonstrate that the DA, sensitivity, and FOM of the first LMR are improved as the Ag layer thickness is increased, as illustrated in Figure 14a,b,c, respectively.

**Figure 14.** (**a**) DA, (**b**) sensitivity, and (**c**) FOM of the first LMR dip as a function of RI of the analyte.

This investigation is extended to the third resonance dip, as depicted in Figure 15. With the increase in the thickness of the Ag layer, DA, sensitivity, and FOM of SPR dip are improved as shown in Figure 15a,b,c respectively.

**Figure 15.** (**a**) DA, (**b**) sensitivity and (**c**) FOM of the SPR dip as a function of RI of the analyte.

Furthermore, we have plotted the absolute square of the electric field component along the interface across all the thicknesses of probes in two configurations- ITO (10 nm) +Ag (10 nm) +ITO (40 nm) and ITO (50 nm) +Ag (10 nm) +ITO (50 nm). The corresponding wavelength values are given in figure captions. These figures clearly show a large enhancement in the field at SPR resonance (see Figures 16c and 17c). These also corroborate the observed high sensitivity for SPR resonance.

**Figure 16.** Electric field distribution corresponding to (**a**) first LMR (**b**) second LMR (**c**) SPR at their respective resonance wavelength at 561, 1065, and 1532 nm respectively for probe configuration ITO (10 nm) +Ag (10 nm) +ITO (40 nm).

**Figure 17.** Electric field distribution corresponding to (**a**) first LMR (**b**) second LMR (**c**) SPR at their respective resonance wavelength at 631, 995, and 1798 nm respectively for probe configuration ITO (50 nm) +Ag (10 nm) +ITO (50 nm).

We would like to note that since in the present study a plastic-clad highly multimode fiber is considered, the proposed probe is only good for room temperature applications. Although slight variations in temperature do not influence the sensor performance, at high temperatures the fiber cladding will melt down.

#### **4. Discussion and Conclusions**

In conclusion, ITO + Ag-based bi-layer and tri-layer fiber-optic sensors have been studied. In the case of bi-layer geometry, two modes of resonances are possible. These are called LMR and SPR, and these resonances can be used for sensing purposes. In this bi-layer configuration, the LMR dip shows better sensitivity compared to the SPR dip. DA of the LMR dip is also far better than that of the SPR dip. As the thickness of the Ag layer increases further, the SPR dip becomes insensitive and only the LMR dip can be used for sensing. We suggest that at this thickness, the SPR dip can work as a reference, and this turns the sensor into a self-referenced sensor. By choosing the appropriate thickness of Ag, this configuration can be used in chemical, and bio-sensing, whereas the same configuration can also be utilized in wavelength filtering.

Further, two configurations of tri-layer geometry are explored, wherein one SPR and two LMR dips have been observed. In the first tri-layer configuration, the thickness of the outmost ITO layer was varied, the first LMR dip that arises in the visible region is found less sensitive as compared to the SPR dip. The other LMR dip (middle dip) that appears in the NIR region is found insensitive to any change in analyte RI. This insensitive LMR dip appears if the thickness of the third layer of ITO is increased. In the second configuration, the thickness of the Ag layer was varied. Particularly, it has been shown that the resonance wavelength of SPR dip shifts toward the shorter wavelength side; however, the resonance wavelength of the LMR dip shifts toward the longer wavelength side. Furthermore, we have plotted the electric field component along the interface across all the thicknesses of probes in two tri-layer configurations to demonstrate field enhancement. The observation of the insensitive second LMR dip and its manipulation with ITO layer thickness variation are the main contribution of this work as this suggests a switching of resonance type between LMR and SPR. Also, this work provides design rules of ITO-based bi- and tri-layer structures which support the excitations of LMR and SPR. The results of the study are summarized in Table 1. We see from the table that SPR dip exhibit a very high sensitivity of 14 μm/RIU and good DA and FOM.

**Table 1.** Summary of the results obtained in bi layer and tri layer configurations.

**Author Contributions:** A.K.M. and S.K.M. discussed the plan and agreed to it. The calculations were done by D.S.G. and A.P. The data was analyzed by all authors. The original manuscript was written by D.S.G. and A.P., S.K.M. and A.K.M. revised and finalized it. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data that support the findings of this study are available from the corresponding author upon request.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Communication* **Wafer-Scale LSPR Substrate: Oblique Deposition of Gold on a Patterned Sapphire Substrate**

**Kihyeun Kim 1,†, Ki Joong Lee 2,†, Na Rae Jo 2, Eun-Jung Jo 3, Yong-Beom Shin 2,4,\* and Min-Gon Kim 3,\***

<sup>3</sup> Department of Chemistry, Gwangju Institute of Science and Technology (GIST), Gwangju 61005, Korea<EMAIL_ADDRESS>

**Abstract:** Label-free detection of biomolecules using localized surface plasmon resonance (LSPR) substrates is a highly attractive method for point-of-care (POC) testing. One of the remaining challenges to developing LSPR-based POC devices is to fabricate the LSPR substrates with large-scale, reproducible, and high-throughput. Herein, a fabrication strategy for wafer-scale LSPR substrates is demonstrated using reproducible, high-throughput techniques, such as nanoimprint lithography, wet-etching, and thin film deposition. A transparent sapphire wafer, on which SiO2-nanodot hard masks were formed via nanoimprint lithography, was anisotropically etched by a mixed solution of H2SO4 and H3PO4, resulting in a patterned sapphire substrate (PSS). An LSPR substrate was finally fabricated by oblique deposition of Au onto the PSS, which was then applied to label-free detection of the binding events of biomolecules. To the best of our knowledge, this paper is the first report on the application of the PSS used as an LSPR template by obliquely depositing a metal.

**Keywords:** patterned sapphire substrate; localized surface plasmon resonance; oblique deposition; nanoimprint lithography; wafer-scale

#### **1. Introduction**

Localized surface plasmon resonance (LSPR), which occurs in metal nanostructures, has attracted considerable attention from various optical- and electrical-based research fields, such as optical biosensors, color engineering, spectroscopy, metamaterials, and plasmonic optoelectronics [1–5]. Thus, large-scale, reproducible, high-throughput fabrication methods of wavelength-tunable LSPR substrates are in high demand not only for research but also for commercialization [6–8]. To fabricate LSPR substrates, two primary methods have been used until now: (i) the synthesis of metal nanoparticles (NPs), and covalent attachment of the NPs on substrates through self-assembled monolayers (SAMs) [9,10]; and (ii) the patterning of metal nanostructures by nanolithography techniques [7,11,12].

Regarding the former (i), various shapes of gold (Au) NPs were synthesized in a solution, such as spheres, rods, triangles, and spiked structures, which could easily tune LSPR properties depending on the shape [13,14]. However, the formation of SAMs on a substrate was usually unstable, which lowered reproducibility for the fabrication of LSPR substrates [15,16]. Moreover, it was difficult to align metal NPs on a substrate [17]; instability issues when metal NPs were immobilized on a substrate were also noted [18,19]. As for the latter (ii), various nanolithography techniques, such as colloidal, e-beam, and nanoimprint have been applied to fabricate metal nanostructures on substrates [20–22]. The limitations of e-beam and colloidal lithography are low-throughput

**Citation:** Kim, K.; Lee, K.J.; Jo, N.R.; Jo, E.-J.; Shin, Y.-B.; Kim, M.-G. Wafer-Scale LSPR Substrate: Oblique Deposition of Gold on a Patterned Sapphire Substrate. *Biosensors* **2022**, *12*, 158. https://doi.org/10.3390/ bios12030158

Received: 17 February 2022 Accepted: 1 March 2022 Published: 3 March 2022

**Publisher's Note:** MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

and non-elaborate patterning, respectively [23]. Although nanoimprint lithography enabled a high-throughput process and elaborate patterning, it also encountered difficulties in the fabrication of various shapes of master molds, leading to limitations in the tuning of LSPR properties [24].

Herein, we report on a fabrication method for wafer-scale, elaborate LSPR substrates using high-throughput methods, such as nanoimprint lithography, wet-etch processing, and thin film deposition. In detail, the LSPR substrates were accomplished by oblique deposition of Au onto a (triangular-pyramid shape) patterned sapphire substrate (PSS) which was obtained through a wet-etch process, resulting in the formation of Au nanostructures on one side of the PSS. Thus, unlike typical strategies for fabrication of LSPR substrates that directly fabricated patterned metal nanostructure using nanolithography [20–22], the strategy suggested in this paper for LSPR-substrate fabrication was to pattern a substrate first, followed by oblique deposition of Au on the patterned substrate. This approach can variously tune LSPR properties according to the shape of PSS controlled by the wet-etch condition, deposition angle and/or direction of metals, which would be highly advantageous for obtaining wavelength-tunable LSPR substrates. Therefore, such a fabrication method could be an inspiring approach for nanofabrication and/or LSPR-based research fields.

#### **2. Materials and Methods**

#### *2.1. Formation of a SiO2-Nanodot on the Sapphire Wafer*

SiO2-nanodot was formed on a sapphire wafer (2 inches in diameter) via nanoimprint lithography, as reported in our previous study (Figure S1) [23]. In brief, a nanodot-patterned master stamp was used to pattern a thermoplastic resist layer that had been coated on the sapphire wafer. Then, Cr was obliquely deposited on the resist pattern to form the Cr hard masks only on the top of the resist layer that was not stamped by the master stamp. Next, O2 plasma was treated to etch the resist layer where it was not covered by the Cr hard mask. A SiO2-nanodot pattern was formed by the deposition of SiO2 using an e-beam evaporator. Finally, the resist layer was completely removed using acetone. Such nanoimprint lithography is a highly advantageous technique for fabricating nanosized patterns on substrates because of its high speed and high-throughput, as well as high reproducibility and fidelity [23]. The diameter, height, and pitch of the SiO2 nanodots were 130, 50, and 300 nm, respectively.

#### *2.2. Fabrication of PSS via Wet-Etch*

The SiO2 nanodot-deposited sapphire wafer was annealed using a furnace under air at 750 ◦C for 12 h to densify the SiO2. Then, the sapphire wafer was soaked in a beaker containing a mixture of H2SO4 and H3PO4 (1:3 *v*/*v*), and heated for 1 h at 310 ◦C to etch the sapphire wafers [25,26]. After finishing the wet-etch process, the beaker was cooled to 20 ◦C, and the PSS was washed with deionized (DI) water.

#### *2.3. Formation of Metal Nanostructures on the PSS*

Au nanostructures were formed on the PSS by oblique deposition (45◦) of Au using a thermal evaporator. The deposition thickness and rate were set to 20 nm and 0.5 Å/s, respectively.

#### *2.4. LSPR Shift of the Au Nanostructure-Formed PSS Due to Biomolecule Attachment*

An LSPR shift was observed by exposing it to streptavidin (STA), biotin-bovine serum albumin (BSA), and STA, in turn. To attach STA onto the Au surface, 100-μM biotin-HPDP was attached to the Au nanostructure on the PSS overnight, followed by washing of the Au/PSS with EtOH and deionized (DI) water. Then, STA (50 μg mL−1) and biotin-BSA (50 μg mL–1) were attached in turns; and at each step, absorbance was measured after washing.

#### **3. Results and Discussion**

To fabricate a patterned sapphire substrate (PSS), SiO2 nanodots—which played a role as hard masks—were well-orderly formed on the entire surface of the sapphire wafer (Figure 1a–c). Figure 1a shows blue light scattering observed when the sapphire wafer was exposed to light due to the well-ordered SiO2 nanodots on the wafer. In addition, SEM images proved that precisely ordered SiO2 patterns were present on the sapphire wafer; the diameter, height, and pitch of the pattern were 130, 50, and 300 nm, respectively (Figure 1b,c).

**Figure 1.** (**a**) A digital image and (**b**,**c**) SEM images of the SiO2 nanodot array on the sapphire wafer (2 inches) ((**b**) top and (**c**) tiled view). (**d**) Schematic of the fabrication of the patterned sapphire substrate (PSS) via wet-etch. (**e**) A digital image of a piece of the PSS. (**f**–**h**) SEM images of the PSS for top, 45◦ tilted, and side views, respectively.

The SiO2 nanodot pattern on the sapphire wafer was annealed using a furnace under air at 750 ◦C for 12 h to densify the SiO2 nanodots before proceeding with the wet-etch process. This process was essential because as-deposited SiO2 nanodots using the ebeam evaporator were not able to act as hard masks under high temperatures and acidic conditions (such as a mixture of H2SO4 and H3PO4), as shown in Figure S2. After the annealing process, the sapphire wafer was wet etched by a mixed solution composed of H2SO4 and H3PO4 at an elevated temperature of 310 ◦C for 1 h (Figure 1d). The composition of a wet-etching solution was 3:1 *v*/*v*.

As a result of the wet-etch, a PSS was fabricated, which was analyzed by digital (Figure 1e) and SEM images regarding top, 45◦ tilted, and side views (Figure 1f–h). Triangular-pyramid shape patterns in the PSS were observed across the entire wafer (Figure 1f,g). The triangular pyramid could have been formed because the SiO2 hard mask prevented direct contact between the single crystal sapphire wafer and acidic etchant (H2SO4 and H3PO4). The height of the triangular pyramids on the PSS was approximately 360 nm (Figure 1h), which varied depending on the wet-etch temperature (Figure S3). This process occurred because although H2SO4 and H3PO4 can etch sapphire as follows (Equations (1) and (2)):

$$\rm Al\_2O\_3 + 3H\_2SO\_4 \to Al\_2(SO\_4)\_3 + 3H\_2O\_2 \tag{1}$$

$$\text{Al}\_2\text{O}\_3 + 2\text{H}\_3\text{PO}\_4 \to 2\text{AlPO}\_4 + 3\text{H}\_2\text{O},\tag{2}$$

the roles of both etchants are different in terms of etching direction (H2SO4 and H3PO4 for perpendicular and lateral, respectively) [25]. In addition, these different behaviors also depended on the temperature. Thus, shape control was achieved by controlling the composition of the etchant. In addition, after the first etch of the sapphire wafer, the second etch of the wafer using a different composition of etchant resulted in sharp or broad triangular-pyramid shapes on the PSS (Figure S4).

Wafer-scale Au nanostructures were fabricated by oblique deposition of Au (20 nm) on the PSS using an evaporator. Oblique deposition (45◦) of Au on the PSS resulted in Au nanostructures being only formed on only half of the available surface area of the triangular pyramids on the PSS, described as red lines in Figure 2a; moreover, a shadowing effect was observed due to adjacent triangular pyramids (red arrow in Figure 2b). For better visualization, a schematic illustration was demonstrated for obliquely deposited Au (45◦)/PSS (Figure 3c). Whereas, vertical deposition resulted in nearly full coverage of Au on the PSS (Figure S5). In the case of oblique deposition of Au, the shape of the Au nanostructures was determined by the shape of triangular pyramids on the PSS, which can be controlled by the composition of etchant and etching temperature (Figures S3 and S4). Consequently, the shape of Au nanostructure can vary not only depending on the shape of the triangular pyramids on the PSS but also the deposition angle and/or direction with respect to the PSS.

**Figure 2.** (**a**,**b**) SEM images after oblique (45◦) deposition of Au onto a PSS ((**a**): top view, (**b**): 45◦ tilted view). Scale bars in the images are 300 nm. Red lines in (**a**) and arrows in (**b**) indicate Au nanostructure and the shadowing effect due to adjacent triangular pyramids, respectively. (**c**) Schematic illustration of obliquely deposited (45◦) Au on PSS.

**Figure 3.** (**a**) Absorbance of obliquely and vertically deposited Au on PSSs, and vertically deposited Au on flat glass and sapphire substrates. (**b**) LSPR shift of the PSS with obliquely deposited Au: (i) biotin-HPDP, (ii) STA, (iii) biotin-BSA, and (iv) STA. (**c**) Schematic illustration for step-by-step biomolecule detection using an Au/PSS. The stages of (i), (i + ii), (i + ii + iii), and (i + ii + iii + iv) correspond to those in Figure 3b.

The phenomenon of a localized surface plasmon resonance (LSPR) of the Au nanostructures on the PSS was observed by measuring absorbance (Figure 3a). LSPR peaks were observed in the cases of Au deposition on a PSS, but not on flat substrates. The LSPR peak appeared in the vertically deposited case because the vertical deposition between the sample and Au sources in the thermal evaporator used in this work was not exactly vertical (Figure S6). Regarding the obliquely deposited case, the LSPR peak was red-shifted compared to that from the vertically deposited case; this was because the shapes of Au nanostructures in each case were different. Different shapes of metal nanostructure induce various LSPR properties [27]. Therefore, the location of the LSPR peak can be controlled by the shape of triangular pyramids on the PSS and deposition angle and/or direction of the PSS.

One of the most popular applications of the LSPR phenomenon is in biosensing fields because the LSPR shift occurs by perturbation in the presence of target molecules located near the metal nanostructures (~10 nm), which is a label-free tool that monitors molecular bindings in real-time, in small-volume samples [28]. Hence, Au nanostructures on PSS were exposed to biomolecules (STA, biotin-BSA, and STA, in turn), and LSPR shifts were observed (Figure 3b). A schematic illustration is depicted for the procedure for biomolecule detection (Figure 3c). In detail, to attach STA on the surface of the Au nanostructures, biotin-HPDP was, at first, covalently functionalized on the Au surface through the thiol-Au interaction [29]. As a result of the step-by-step attachment of STA/biotin-BSA/STA, the LSPR peak of the Au nanostructures was red-shifted step-by-step due to the attachment of the biomolecules (detailed values of LSPR are shown in Table S1). The redshift phenomenon occurred because the biomolecules attached to the Au nanostructures caused a change in the refractive index near the Au surface. Thus, our Au nanostructure on PSS could have an application as LSPR biosensors by manipulating the structures of the PSS and/or oblique deposition conditions. Furthermore, a thorough investigation of LSPR properties by controlling the PSS and deposition variables, such as the plasmonic substrate, could be further applicable to color engineering, spectroscopy, metamaterials, and plasmonic solar cells.

We also fabricated Au/PSSs where Au was deposited at different angles: 30◦, 45◦, and 60◦ (Figure S7). Among them, the LSPR peak obtained from Au (45◦)/PSS showed the sharpest. This work is a preliminary study demonstrating the application of PSS for LSPR substrate by obliquely depositing Au on the PSS, and thus, LSPR properties of the Au nanostructure on PSS can be improved further, depending on Au deposition angle, the thickness of deposited Au, the direction of PSS under oblique deposition of Au, and the shape of PSS.

#### **4. Conclusions**

In this paper, we have suggested a fabrication method of wafer-scale LSPR substrates by engineering a single crystal substrate and obliquely depositing Au, using highthroughput and high-reproducibility techniques. As a single crystal substrate, sapphire wafers were chosen because of their transparency, which is highly advantageous for optical applications, such as optical biosensors, metamaterials, and solar cells. The PSS was fabricated using a nanoimprint lithography technique, thin film deposition, and wetetching (using H2SO4 and H3PO4). The oblique deposition (45◦) of Au on the PSS enabled the formation of Au nanostructures on the PSS due to the three-dimensional structure (triangular-pyramid shape) of the PSS. The Au nanostructures formed by the oblique deposition exhibited a red-shifted LSPR peak compared to those prepared by vertical deposition, implying that LSPR properties can be controlled not only by the shapes of the PSS but also by the deposition conditions, such as angle and/or rotation of the PSS. The Au nanostructure on PSS was applied to detect biomolecules by observing the LSPR shift; the LSPR peak was red-shifted by attaching STA, biotin-BSA, and STA, in turn.

PSS has been widely used as a substrate for light-emitting diode, so far due to its transparency, wafer-scale patterning, and high-throughput production. Here, we applied such a PSS to the fabrication of an LSPR substrate by obliquely depositing Au on the PSS. Further technical approaches could improve our fabrication strategy to expedite its application for commercially available products. As an example of the approaches, some experimental variables—Au deposition angle, the thickness of deposited Au, the direction of PSS under oblique deposition of Au, and the shape of PSS—could be addressed. In addition, due to the high expense of the sapphire wafer, the PSS might be used as a master mold to fabricate the reverse-shaped polydimethylsiloxane (PDMS) that could be used as patterned PDMS substrate for the angle deposition of Au.

Thus, such a fabrication method of wafer-scale plasmonic substrates could pave the way for use in LSPR biosensors by manipulating the structures of the PSS and/or oblique deposition conditions. This technology is also expected to be useful in other optical applications, such as color engineering, spectroscopy, metamaterials, and plasmonic optoelectronics.

**Supplementary Materials:** The following supporting information can be downloaded at: https://www.mdpi.com/article/10.3390/bios12030158/s1, Figure S1: Schematic of the formation of a SiO2 nanodot array on a sapphire wafer, Figure S2: SEM image of the SiO2 nanodot (not annealed) patterned sapphire wafer after the wet-etch, Figure S3: SEM images of a PSS, Figure S4: SEM images after the second wet-etch of PSS, Figure S5: SEM images of vertically deposited Au on PSS. Figure S6: Schematic illustration of vertical and oblique depositions of Au on PSS, Figure S7: Analysis of Au/PSSs where Au was deposited at different angles, Table S1: LSPR-shift observation of obliquely deposited Au on PSS when exposed to biomolecules.

**Author Contributions:** Conceptualization, K.K., Y.-B.S. and M.-G.K.; methodology, K.K., K.J.L. and N.R.J.; software, N.R.J.; validation, K.K., K.J.L. and M.-G.K.; formal analysis, K.K. and K.J.L.; investigation, K.K. and K.J.L.; resources, K.K., Y.-B.S. and M.-G.K.; data curation, K.K. and K.J.L.; writing—original draft preparation, K.K., K.J.L. and E.-J.J.; writing—review and editing, K.K., K.J.L., E.-J.J., Y.-B.S. and M.-G.K.; visualization, K.K. and K.J.L.; supervision, Y.-B.S. and M.-G.K.; project administration, Y.-B.S. and M.-G.K.; funding acquisition, K.K., Y.-B.S. and M.-G.K. All authors have read and agreed to the published version of the manuscript.

**Funding:** This work was financially supported by a National Research Foundation grant funded by the Ministry of Science, ICT, and Future Planning as the MidCareer Researcher Program (NRF-2021R1A2B5B3001417); and Nano·Material Technology Development Program through the National Research Foundation of Korea (NRF) funded by Ministry of Science and ICT (Grant number NRF-2021M3H4A4079382). This research was also supported by a grant of the Korea Health Technology R&D Project through the Korea Health Industry Development Institute (KHIDI), funded by the Ministry of Health & Welfare, Republic of Korea (HI21C0957).

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The authors confirm that the data supporting the findings of this study are available within the article and its Supplementary Materials.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

## *Review* **GLAD Based Advanced Nanostructures for Diversified Biosensing Applications: Recent Progress**

**Sarjana Yadav 1, Sneha Senapati 2, Samir Kumar 3, Shashank K. Gahlaut 1,† and Jitendra P. Singh 1,\***

	- New Delhi 110016, India
	- <sup>3</sup> Department of Electronics and Information Engineering, Korea University, Sejong 30019, Republic of Korea
	- **\*** Correspondence<EMAIL_ADDRESS>
**Abstract:** Glancing angle deposition (GLAD) is a technique for the fabrication of sculpted micro- and nanostructures under the conditions of oblique vapor flux incident and limited adatom diffusion. GLAD-based nanostructures are emerging platforms with broad sensing applications due to their high sensitivity, enhanced optical and catalytic properties, periodicity, and controlled morphology. GLAD-fabricated nanochips and substrates for chemical and biosensing applications are replacing conventionally used nanomaterials due to their broad scope, ease of fabrication, controlled growth parameters, and hence, sensing abilities. This review focuses on recent advances in the diverse nanostructures fabricated via GLAD and their applications in the biomedical field. The effects of morphology and deposition conditions on GLAD structures, their biosensing capability, and the use of these nanostructures for various biosensing applications such as surface plasmon resonance (SPR), fluorescence, surface-enhanced Raman spectroscopy (SERS), and colorimetric- and wettability-based bio-detection will be discussed in detail. GLAD has also found diverse applications in the case of molecular imaging techniques such as fluorescence, super-resolution, and photoacoustic imaging. In addition, some in vivo applications, such as drug delivery, have been discussed. Furthermore, we will also provide an overview of the status of GLAD technology as well as future challenges associated with GLAD-based nanostructures in the mentioned areas.

**Keywords:** glancing angle deposition (GLAD); biosensors; biomedical detection; GLAD substrates; plasmonics

### **1. Introduction**

Biosensors have emerged in recent decades as a quintessential tool in the fields of healthcare, bioscience, the food industry, the chemical and biomedical industries, defense and security, and environmental monitoring [1]. Any system or chip consisting of a surface that is capable of binding to a specific biological substance that recognizes it, together with the ability to translate the biophysical or chemical reaction into a quantifiable signal, can be referred to as a biosensor [1]. The biosensor recognizing element may be nucleic acids, proteins, carbohydrates, lipids, antibodies, or an entire cell. These elements may be attached or immobilized on the sensor using various methods such as adsorption, bond formation, encapsulation, or entrapment [2]. Selectivity, specificity, reproducibility, stability, and sensitivity are various characteristics of a typical biosensor. Recently, the scientific community has witnessed vast development in optical, chemical, and physical transducers along with a specific affinity for biochemical interaction. This has led to vast progress in the development of biosensors in various fields. The most widely and extensively used biosensors are those that are based on nano-regimes due to their low detection limits, high stability, easy operation, low cost, short response time, and low power requirement [2]. Metallic-based nanostructures, therefore, possess specific unique features that ensure high

**Citation:** Yadav, S.; Senapati, S.; Kumar, S.; Gahlaut, S.K.; Singh, J.P. GLAD Based Advanced Nanostructures for Diversified Biosensing Applications: Recent Progress. *Biosensors* **2022**, *12*, 1115. https://doi.org/10.3390/ bios12121115

Received: 27 October 2022 Accepted: 23 November 2022 Published: 2 December 2022

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**Copyright:** © 2022 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (https:// creativecommons.org/licenses/by/ 4.0/).

sensitivity of the biosensors compared to other planar sensor configurations due to their nanoscale dimension and periodicity, thus ensuring strong binding of the target analyte and increased sensing properties [3].

The rapid advancement in the field of biosensing utilizing plasmonics has led to an extensive study of metallic nanostructures and their sensing applications [4–6]. One such emerging tool for fabricating metallic nanostructures is the glancing angle deposition (GLAD) technique [7]. It is a physical vapor deposition (PVD) process in which the gaseous flow of depositing atoms impinges on the substrate in a vacuum at an oblique angle such that the deposited film results in a columnar morphology [7,8] as a result of the shadowing effect. The tilting of the as-deposited column is controlled by the angle of incidence. The morphology of the columnar structure can be modified using a mobile substrate whose position can be managed during deposition [9,10]. The GLAD technique is considered better than the other chemical growth methods in terms of ease of fabrication of nanostructures due to controlled size, shape and spacing of columns in single step [11] Additionally, the nanostructures grown by this method possess high surface area to volume ratio [12], high porosity [13], crystallinity, and uniformity [11,14] making it the most suitable method for sensing applications. These structure has been widely investigated in designing different sensor devices [15], energy devices [16], and optical devices [17]. The GLAD technique has been extensively studied and developed over the past 15 years, but there is still much to explore.

The last few decades have witnessed a high increment in gold and silver-based nanostructures for catalytic [18] and biomedical applications such as immunotherapy, nanomedicine, and in vitro diagnostics [19–22]. Silver- and gold-based nanostructures are most widely used for sensing applications due to their high plasmonic properties [23–25], particularly their extinction band lying in the visible region [26] and low dispersion losses. Silver has a strong bactericidal effect due to the interaction of Ag ions with the thiol group and antimicrobial activity against microbes [27,28]. Gold has an excellent affinity for binding to antibodies [29], and hence, is most suited for antigen–antibody surface plasmon resonance [30,31] (SPR)-based detection. Silver- and gold-based nanostructures are shown to be the preferred transducers for LSPR-based sensing [32], surface-enhanced Raman scattering (SERS)-based sensing [33], as well as fluorescence resonance-based energy transfer [34] (FRET) sensing [35]. The highest SERS enhancement obtained using Ag nanorods prepared using the GLAD technique is 109 [36], so GLAD-based nanostructures combined with other techniques provides variations in metallic nanostructure shape, pattern, and period, extending the scope of their sensing abilities. The colorimetric- and wettability-based properties of these metallic nanostructures are found to be highly sensitive towards certain chemical and biological processes, revealing their use in biosensing detection [37–39]. The ongoing research in biosensor technology highlights an increase in the use of nanostructures and nanoparticles as a receptor or chip to enhance detection sensitivity and capability. Nanostructure-based biosensors or biochips fabricated using PVD methods exhibit advantages over conventionally prepared nanomaterials in terms of their high surface-to-volume ratio, fast response time, high versatility, multiple-fold increase in sensitivity, and better target-specific element recognition [40].

In this review, the biosensing and other imaging properties of GLAD-fabricated nanostructures are evaluated. We shall discuss the fabrication of these nanostructures with different types of substrate structures along with their implications in biosensing and other biomedical applications. The abilities of these nanostructures in various biosensing fields have been explored. A comparison of GLAD-based nanostructures with conventionally used nanoparticles and chips will be highlighted along with their advantages over the latter. There are various advantages of GLAD-based deposition over other chemical route depositions; for example, no precursor materials are needed in GLAD, which provides safety from toxic precursors and by-products [41]; atomic-level control of chemical composition and thickness can be afforded in this deposition [42]; and a higher temperature is not required and heat-sensitive substrates can be fabricated using this technique. GLAD

is a green fabrication technique which produces very little pollution [43], promising high uniformity and repeatability in the structures [44–46].

#### **2. Recent Strategies for Fabrication of Nanostructures Using the GLAD Technique**

PVD methods such as sputtering, electron-beam vapor deposition, and thermal evaporation are most commonly used to fabricate thin films. When the vapor flux is normally incident over the substrate, thin-film formation takes place. However, if the angle of incidence is large with respect to the normal of the substrate surface (>70◦), columnar structures are formed. This type of PVD deposition, carried out using either thermal evaporation or e-beam evaporation, is referred to as oblique-angle deposition (OAD) [3]. This extended OAD combined with azimuthal substrate rotation, giving various shapes and size of nano-columns, is referred as GLAD [14,47] or as dynamic oblique-angle deposition by some researchers [48]).

#### *2.1. Setup Configuration of GLAD*

Figure 1a shows the schematic of a GLAD setup where α is the angle subtended by the incident vapor with the substrate normal, called the deposition angle. Φ is the azimuthal angle that measures rotation about the substrate normal. The vapor flux falls on the substrate at a particular glancing angle (>700), initially leading to the formation of random nuclei, which is supported by the Volmer–Weber growth mechanism [49]. These further grow into isolated columns as a result of competition between shadowing and adatom diffusion due to larger nuclei and adatom surface diffusion. The columns lead to shadowing with size distribution and the screening of smaller neighboring nuclei from the coming flux. Therefore, the growth of the smaller nuclei is suppressed and ultimately ends due to complete shadowing [7]. This process of column extinction continues throughout the film's growth.

**Figure 1.** (**a**) Schematic of GLAD setup where α is the angle subtended by the incident vapor flux with respect to substrate normal and Φ is the azimuthal angle. (**b**) Mechanistic view of GLAD-based growth. (i) Incident vapor flux, shown by full arrows at an angle (α), resulting in random nucleation. (ii) Nucleus growth casting a nearby shadow, as shown by dotted lines. (iii) Column growth in progress, suppressing neighbors' growth via shadowing. (iv) Full column growth at a certain specific angle (β) with respect to the normal.

More vapor flux deposition takes place as the nucleus growth continues developing into isolated columns. Finally, the isolated columns are formed, and are tilted towards the source of the vapor flux. β is the column tilt angle. This process is shown in Figure 1b. It must be noted that β is different from deposition angle α, and β = 0 for the columns

parallel to the substrate. The relationship between angles α and β is given by the tangent rule, given by Niewenhuizen and Haanstra [50], as

$$
\tan \alpha = 2 \tan \beta \tag{1}
$$

when the deposition angle increases obliquely with respect to the normal, in which case, the tangent rule is not followed.

$$\beta = \arcsin\left(\frac{1-\cos\alpha}{2}\right) \tag{2}$$

#### *2.2. Effect of Various Growth Parameters* 2.2.1. Tilt Angle

As discussed earlier, the dependence of the tilt angle cannot be fully described in terms of angle of incidence by a single relationship. The general relationship is given by the tangent rule (Equation (1)), and Tait's rule [51] (Equation (2)) is satisfied only under certain conditions and must be cross-verified for a specific material under the given conditions such as deposition rate, pressure, substrate type, and temperature. All these factors influence the tilt angle β. Nakhodkin and Shaldervan investigated the nature of films of different materials (Ge, Si, GeTe) as a function of tilt angle [52]. They found that most of the films are formed of crystallites inclined to the substrate due to obliquely incident vapor flux. They classified and correlated the structures based on material, cross-section, and tilt angle up to a thickness of 300 nm. Salazar et al. used GLAD-fabricated Cu electrodes at a 65◦ tilt angle to study their sensing ability for the detection of glucose in blood samples [53]. Most of the recent reports showing GLAD-based biosensing applications have employed tilt angles ranging from 65◦ to 85◦ with high sensitivity and reproducibility [53,54].

#### 2.2.2. Azimuthal Rotation (φ)

This represents the rotation of the substrate about the normal to the substrate. The azimuthal angle φ measures the rotation of the substrate, exploiting the growth direction steering the columnar structure along the growth path. The substrate position and the effect of azimuthal rotation on growth is depicted by α and φ and their variations. Control of both of these angles in a desirable manner during substrate rotation is the most crucial part of GLAD technology. If α and φ remain constant throughout the deposition, then it means that the substrate is stationary. α and φ may also change periodically at a given angle, but otherwise, may remain stationary [7]. There are six to nine basic GLAD structures that can be formed by different combinations of α and φ, which, along with their implications in biosensing, are discussed in detail in the next section. This variety of microstructures provides more enhanced properties that have a strong effect on the binding of biomolecules to the substrate and cause a change in the plasmonic properties of these substrates. Zhang et al. investigated the dependence of α and Φ on the biosensing ability of a Au nanostructured array fabricated over closely packed polystyrene balls and obtained the highest SPR and LSPR sensitivity for optimal values of α and Φ [55]. This regime as a function of tilt angle or azimuthal angle, due to the corresponding change in the substrate morphology and column growth, has not yet been fully explored and still has a long way to go.

#### *2.3. GLAD-Based Nanostructures*

The GLAD deposition technique can be used to create thin films of different morphology and sizes. As described in the previous section, the vapor flux falls on the substrate at a certain glancing angle (α > 70◦), resulting in the formation of microstructures because of competition between the shadowing effect and adatom diffusion. The formation and categorization of various GLAD structures depend on the variation in α and ϕ during the deposition process.

The formation of the plane metallic film takes place when the substrate is held perpendicular to the vapor flux (α = 0◦). However, this type of thin film is not very sensitive to biomolecule detection. Agrawal et al. has shown, using Rigorous coupled-wave analysis (RCWA) simulation, how nanohole array in a gold film can be employed for enhanced biosensing compared to planar gold film [56]. Some roughness over the surface of the nanostructure increases the sensitivity of biomolecules multiple-fold [57,58]. Slanted nanorods are the most commonly used GLAD structure for biosensing. These are formed when the substrate is held in the stationary position. α is taken to be greater than 70◦ (~86◦). However, other desired micro- and nanostructures can be sculpted by varying the position of the substrate, and hence, the direction of the incoming vapor flux [59]. High sensitivity and plasmonic enhancement have been obtained using slanted nanorods with single arms and multiple arms [60]. Slanted nanorods have been utilized and studied widely in various biosensing applications using different materials (Ag, Si, Ni, TiO2) [60–62] due to their ease of fabrication, lower material requirement, lower complexity, and high optimized sensitivity [36].

Chevron or zig-zag multi-armed nanorods are formed by a series of slanted nanorods. The substrate is initially held at a specific angle (α, ϕ) for the first arm, then, rotated by 180◦ (α, ϕ + π) for the second arm, and again by 180◦ (α, ϕ + 2π) for the third arm, etc.; that is, every time, the direction is rotated by 180◦ [7]. Square spirals can also be fabricated in the same way as chevrons with a rotation of 90◦. The optimized slanted nanorods have shown enhanced biosensing and have been widely employed for SPR-, SERS-, colorimetric-, and wettability-based biosensing and detection. Sensitivity comparisons of various bioanalytes using various types of these GLAD-fabricated nanostructures have been performed by different groups [63]. Enhanced sensitivity, which can be employed for biological sensing using vertically standing nanoporous zig-zag or chevron nanorods providing hot spots with high SERS enhancement, has been shown by researchers [64,65]. Vertical nanorods are formed when the substrate is rotated at a constant rate during deposition. If the substrate is rotated at a constant slow rate, a helical columnar structure is formed with a constant pitch. Similarly, high and low morphology can be formed with a continuous change in ϕ while changing α in a discrete manner [66]. These structures form a Bragg stack and can be highly suitable for optical biosensing applications. Recently, Kumar et al. reviewed the fabrication and recent biochemical applications of various GLAD-fabricated substrates for SERS. They discussed, in depth, the advantages of these substrates and how these overcome the various limitations of conventional substrates [67], (Figure 2a–e). These GLAD-fabricated nanostructures have recently been employed in various biosensing detection settings. Recently, Kim et al. reported OAD-based LSPR substrates with a very high throughput by fabricating Au nanostructures (at 45◦) over patterned sapphire substrates (PSS) [68] (Figure 2f). Very high LSPR sensitivity was obtained for various biomolecules and bioanalytes.

**Figure 2.** Different GLAD-structured thin films. (**a**) Slanted-post films; (**b**) zig-zag (chevron) films with 4 arms; (**c**) AgNR/SiO2 SCL/SiO2/PCL/Ag multilayer substrates; (**d**) helical film; (**e**) vertical post structure (reproduced with permission from [67]); (**f**) 45◦ tilted top view of Au mushrooms (reproduced with permission from [68]).

Hence, GLAD is a sophisticated technique used to design thin films with specific geometries. The geometries can be tailored in a controlled and desired manner to provide specific advantages for detection over conventional nanoparticles and nanofilms. The analyte binding sites, as well as the hotspots and scattering points (typically for optical sensing), can be increased in the desired way using this technique. The substrates can be fabricated in a controlled manner using multifunctional materials of different kinds by varying the substrate orientation and deposition conditions. The detailed bio-detection and sensitivity of various geometries and structures are yet to be explored in depth and could provide high scope in the future; they may provide certain specific binding sites for the proteins, amino acids, etc. present in biomolecules typically suited for SERS or fluorescence detection or antigen–antibody binding sites for SPR and LSPR.

#### **3. Designing Substrates and Chips for Biosensing Applications**

The sensitivity, binding, limit of detection, and other essential biosensing activities of a biosensor strongly depend on the substrate. However, there are only a few methods that can be used to develop uniform, reproducible, robust, stable, and cost-effective substrates. Recently, the GLAD technique has been employed to fabricate columnar thin-film substrates with high sensitivity [60,69]. Various studies have been conducted to study the sensing ability of the columnar Ag film [70,71]. Zhao et al. studied the s- and p-polarization absorbance spectra of Ag nanorod arrays with different heights and topologic shapes (cylindrical, needlelike, periodic, L, and Y shapes) in detail [72,73]. Additionally, the nanorods with complicated shapes, such as periodic, L, and Y shapes, had more hot spots for optical sensing applications than perfectly aligned nanorods. Thus, it is more suitable to observe the biosensing abilities of a small number of molecules on irregular nanorods than on a perfectly aligned nanorod. Recently, various types of GLAD-based nanostructures have been tested, optimized, and fabricated to increase the sensitivity of bioanalytes and expand their applications in the biomedical field. These substrates offer numerous advantages over

the conventional GLAD-based substrates as they are more sensitive, durable, and flexible, and bind bioanalytes more effectively. The biosensing applications of the various newly fabricated substrates in the field of biomedicine and biosensing are discussed below.

#### *3.1. Low-Temperature Biosensing Chips and Arrays*

Aligned AgNR array films have been fabricated at low temperatures (LT; 140 K) and also at higher temperatures (HT; 373 K) [74,75]. The effects of temperature, from low to high, on the deposition of nanostructures and the morphology of films have already been discussed in the earlier section. Singh et al. deposited tilted AgNRs on glass slides at a substrate temperature of T = 140 K. The surface morphology changed drastically for the LT-deposited Ag nanostructured film. The aligned AgNR array SERS substrates deposited at a low-temperature oblique angle exhibited better SERS enhancement with good optical transmission compared to the standard room-temperature (RT)-deposited AgNR SERS substrates. The LT-deposited Ag nanostructured film was more porous, whereas the RT-deposited Ag nanostructured films appeared to be solid and rod-like. The nanorods with lengths of 1 μm, corresponding to a d = 2 μm QCM reading, deposited at room temperature showed the maximum spectral intensity, which is advantageous for optical (SERS, fluorescence, SPR) biosensing applications. However, for AgNR, the LT AgNR substrates, along with the maximum response, also provided substantial cost savings. A recent study by Yadav et al. demonstrated that magnetic surface-enhanced Raman spectroscopy (M-SERS) was effective in detecting hemozoin, a metabolic by-product of malarial parasites with paramagnetic properties. Using the GLAD technique, the SERSactive AgNRs were fabricated over neodymium substrates (0.3 T) kept at low temperature (120 K). Both hemozoin and hemoglobin were measured via a magnetic field-based SERS method on M-SERS substrates. In the absence of an external magnetic field, the intensity of the SERS signal was reportedly ten times greater than that of conventional SERS substrates. Figure 3a,b shows a SEM image of AgNR arrays fabricated over glass substrates at room temperature and at low temperature (120 K), respectively. Bunches of uniform nanorods were formed along the direction of vapor flux with slight variations in their lengths. The SEM images of AgNRs grown on magnetic substrates at a 120 K substrate temperature indicate the zone A type of morphology. A comparison in the XRD pattern between the crystalline phase of AgNRs at room temperature and at low temperature is shown in Figure 3c (JCPDS no. 04-0783). AgNR arrays exhibit a polycrystalline structure, as indicated by the XRD patterns. As a result of the differences in peak intensities and FWHM values between the two structures, a small difference in crystallinity can be observed.

#### *3.2. Flexible Substrates*

Compared to conventional hard and rigid substrates, flexible substrates and chips have several advantages. They are able to conform to the underlying object, allow efficient extraction of the target molecule from complex surfaces, have highly physicochemical properties, and can be integrated into wearable devices that provide bio-detection and health-monitoring services. Polydimethylsiloxane (PDMS) and polyethylene terephthalate (PET) sheets are the most common substrates used for the growth of AgNR on flexible substrates [76–78]. Singh et al. prepared a mechanical-strain-resistant AgNR on flexible PDMS and PET substrates [76]. The SERS response of these flexible SERS substrates was comparable to the AgNR on conventional glass slides. Flexible SERS substrates are capable of withstanding high tensile strain (ε) (30%) without losing their SERS performance. However, similar bending can reduce SERS performance by up to 13%. Kumar et al. fabricated flexible and robust SERS-active substrates by embedding the AgNR into the PDMS. Peipei and Young presented a flexible nanoprobe for high-performance real-time SPR biosensing. They constructed plasmonic optical fiber by transferring the metallic nanostructure from patterned templates to optical fiber using epoxy and reported a high figure of merit for SPR sensors.

**Figure 3.** (**a**) SEM images of AgNRs over glass substrates at room temperature (300 K). (**b**) SEM images of AgNRs at 120 K deposited using GLAD technique. (**c**) XRD pattern for the two AgNR SERS substrates. (**d**) Comparison of the Raman and SERS spectra of BPE analyte over the three different substrates (c-SERS and M-SERS substrates, and 200 nm thin Ag film). Reproduced with permission from [63].

#### *3.3. Multilayer Substrates*

Increased sensitivity and strong field enhancement can be achieved in metal–dielectric– metal multilayer structures [79]. Suzuki et al. fabricated Ag and Au nanorods with quasi-parallel principal axes on a SiO2 shape control layer (SCL) with a strongly anisotropic surface morphology [80,81]. They also demonstrated that the morphological and optical properties could be tuned by the deposition conditions without any pre- or post-treatment. Additionally, these structures supporting surface waves such as Bloch surface waves (BSW) could be used for advanced biosensing applications with enhanced sensitivity for protein interaction detection [82]. Multilayered structures have been used not only to tune the plasmon resonance but also to increase the number of hotspots [83,84]. Rajput et al. fabricated four-arm zig-zag Ag nanostructures of Ag and Al, with Al deposited on the edges of the zig-zag structure during growth [64]. The researchers etched less stable Al using a 2.5 wt% HCl solution from the Ag–Al zig-zag structure to generate intraparticle gaps, as depicted in Figure 4. These gaps at the elbows of the nanorods acted as additional hotspots while maintaining the zig-zag morphology. In their study, the SERS performance increased with the number of the increasing arms, which was attributed to the increased hot spots along the edges of the elbows of the zig-zag structure. The SERS performance further improved in the generation of nanoporous zig-zag Ag nanostructures obtained via de-alloying, and a high enhancement factor of ∼106 was observed. These Ag structures are further under study for their applications in enhanced biosensing and early disease detection. Sun et al. constructed a sensitive biosensor based on self-assembled

multilayer film for organophosphate pesticide using CdTe as a fluorescence probe. This multilayer film showed 100 times the sensitivity of CdTe in aqueous solution. Bhalla et al. created a new fabrication approach by exploiting the reactive plasma environment of SF6 to assemble gold nanoplasmonic structures [85]. They developed mushroom-like gold-topped nanostructures on SiO2 substrate. They used it as an LSPR chip for generic bioassay characterization and applications. A uniform protein coating was applied over the LSPR chip, achieving an LOD of 66 zM for biomolecules. These recent multilayered fabricated chips and nanostructures, therefore, are reported to show good stability and easy regeneration, along with increased bio-sensitivity. These can be employed as a promising platform for point-of-care diagnostics.

**Figure 4.** (**a**–**e**) An illustration of the fabrication of nano-porous zig-zag AgNR arrays; SEM images of zig-zag AgNR arrays on silicon substrate with: (**f**) 1 arm; (**g**) 2 arms; (**h**) 3 arms; and (**i**) 4 arms, respectively. Yellow lines show the approximate arm positions. (**j**) BPE SERS spectra (0.1 mM) acquired over zig-zag AgNR arrays with different numbers of arms. (**k**) Enhancement factor for the BPE SERS peak at 1200 cm−<sup>1</sup> as a function of arm number. Calculated enhancement factor values for (**l**) 2-arm zig-zag AgNRs and 3-arm nanogap AgNR arrays and (**m**) 3-arm zig-zag AgNRs and 3-arm nanogap AgNR arrays. Reproduced with permission from [64].

#### *3.4. Patterned Substrates*

Patterned substrates have been demonstrated by various groups; periodically arranged nanostructures might result in enhancement in the localized electric field of plasmonic metals compared to that of randomly arranged nanostructures employed for optical sensing applications [86]. Studies have been performed to couple the periodicity of the grating with the metal nanostructure [87], localize the deposition within a micro-well, [88,89] formation of a hexagonal lattice of nanorods, [90,91] and couple discrete metal nanoparticles to photonic-crystal surface-resonant modes [92]. Researchers have demonstrated the feasibility of patterned substrate formation by combining GLAD and electron beam lithography; thus, they have designed numerous patterned structures in which structure period and structure diameter can be adjusted in a controlled manner, hence controlling the sensitivity [86]. Therefore, this type of substrate can be useful for numerous biosensing applications where patterned structure dimensions play an important role. Anisotropic organometallic halide perovskite nanowalls fabricated using the GLAD technique have recently been reported [93]. Their high plasmonic biosensing ability can be extracted and employed for early and lowcost efficient detection and analysis. Lee et al. [94] developed a unique 3-dimensional patterned substrate employed in ultra-thin-layer chromatography (UTLC) and SERS-based sensing consisting of Si nanowire arrays fabricated using a facile photolithographic and chemical etching technique, and further decorated it with Ag nano-dendrites. These patterned substrates can be further employed for SERS-based study for the distinction and detection of diseases and their co-infections. These multifunctional substrates could simultaneously be used for target localization, mixture separation and label free detection, and therefore, could be readily employed for biochemical assays and food safety. Kumar et al. proposed a flexible nanoslit-based sensor using patterns on periodic optical disks employing imprint lithography for periodic pattern printing on polydimethylsiloxane (PDMS) [95]. Patterned PDMS was obtained using a DVD-R disk and Ag was deposited over the patterned substrates using dynamic oblique-angle deposition at an angle 76◦ The sample was rotated by 180o after 5 nm deposition to form a nanoslit in the grating valley as a result of the shadowing effect. Films of thickness 50–400 nm were finally obtained. Figure 5 shows the SEM images (top view) of Ag nanoslit samples of varying thickness. They further investigated the SERS-based biosensing capability of these substrates using a bilirubin solution. Therefore, patterned substrates can be used as highly sensitive multipurpose substrates in biochemical sensing.

#### *3.5. Coated and Co-Deposited Nanostructured Thin-Film Substrates*

The practical application of Ag nanostructures has been limited by their chemical instability due to the oxidation, sulfuration, and etching of Ag. Alloyed or coated substrates made via the co-deposition of two or more materials have been studied for increased performance and shelf life [95]. There are various reports on increasing shelf life without compromising the sensing performance of a substrate either by coating the Ag with an ultra-thin layer of graphene, SiO2, [96] Al2O3, [97] or by fabricating bi-metallic nanostructures [95,98–100]. The use of low-temperature atomic layer deposition (ALD) to wrap AgNR arrays with a very thin but dense Al2O3 layer can improve their robustness in morphology up to 400 ◦C, and can provide surface coverage that stabilizes SERS activity in air without reducing their sensitivity to SERS.

**Figure 5.** SEM images of Ag nanoslit substrates (top view) with thicknesses of (**a**) 50 nm, (**b**) 100 nm, (**c**) 150 nm, and (**d**) 250 nm. (**e**) Cross-sectional view of 100 nm thickness. (**f**) Net thickness of the film deposited at different substrate points (black region) with orange dotted lines showing incident flux direction and gray arrows showing columnar growth direction. (**g**) Plot of slit width vs. film thickness with a scale bar of 250 nm. (**h**) SERS spectra of 0.06 mg/mL bilirubin over Ag nanoslit substrate corresponding to 100 nm thickness. Reproduced with permission from [87].

#### *3.6. Recyclable Chip and Substrates*

Most of the chips and substrates fabricated for sensing applications are for one-time use only, and given the value of noble metals, these cannot be fully explored as a routine analytical technique. Much study has, therefore, been focused on developing reusable substrates [101]. Kumar et al. fabricated a reproducible and reusable Ag nanoparticledecorated TiO2 SERS substrate since TiO2 is renowned for its photocatalytic properties. The intensity of the Raman signal decreases rapidly in the presence of UV light because the dye concentration decreases rapidly. Approximately 150 min after illumination with UV light, these Ag-TiO2 nanorod SERS substrates recover completely. Hu et al. [102] reported a highly recyclable and sensitive biosensing platform comprising carbon-doped TiO2 nanotube arrays. These can be refreshed and regenerated photo-catalytically, maintaining high selectivity and sensitivity when irradiated with UV or infrared light. Hunt and Armani [103] developed a new idea to generate recyclable optical microcavity via a dry chemistry method using oxygen plasma treatment. This recycling treatment, therefore, helped in reducing cost and preventing harsh and wet chemistry treatment for recycling.

#### *3.7. Microfluidics-Based Substrates*

There is tremendous interest in developing nanofluidic channels with a large surfaceto-volume ratio that exhibit localized surface plasmon resonance due to the strong interaction between the fluids and the channel walls [104]. The advantages of the nanofluidic and microfluidic platforms are: reduced unwanted and nonspecific binding, and detection time minimization by studying multiple analytes or ligand interactions using a single device [105]. By using these nanofluidic channels with plasmonic nanostructures, SERS signals from molecules trapped near electromagnetic hotspots were enhanced. A monolayer of Ag film over silica nanospheres (AgFONs) was fabricated on a glass surface and covered with a polydimethylsiloxane (PDMS) slab. This amplifies SERS by a factor of ten or more, allowing label-free detection of important neurotransmitters in the nanomolar range, rather than using only plasmon resonance.

#### **4. Biosensing Applications of GLAD-Fabricated Nanostructures**

#### *4.1. SPR-Based Detection*

Surface plasmon resonance is a powerful analytical method for detecting the high sensitivity of biomolecules, and is based on plasmonic materials' optical properties. The SPR method has shown great promise with large sensitivity for biomolecular applications [106], especially for real-time biomolecule interaction, protein action, antigen, and nucleic acid detection [107–109]. SPR-based sensors are useful for label-free detection because of the very high sensitivity of the surface plasmon polaritons (SPP) to changes in the local environment in the measurement of both refractive indices and dielectric constants [110,111]. SPR-based biosensors possess the advantage of high versatility and can be tailored for the detection of a vast number of analytes since this type of detection does not demand any special characteristic of the bioanalyte, such as fluorescence properties, or absorption or scattering bands. Additionally, SPR-based bio-detection does not require analytes or the biomolecules to possess fluorescent or radioactive labels.

The primary sensing phenomenon in SPR is related to propagating surface plasmon polaritons. A surface plasmon polariton (SPP) is an electromagnetic wave propagating at the metal–dielectric boundary. The electromagnetic field of an SPP confined at the metal–dielectric interface decreases exponentially in both metals and dielectrics [112]. The dielectric constant of the propagating surface plasmon wave is highly sensitive to changes in the local refractive index of the dielectric. This property is the underlying principle of SPR biosensors. Metal-binding molecules or the recognition element attached to the surface of the metal capture the analyte, resulting in a change in the local refractive index at the metal surface. This change in the local refractive index creates a change in the propagation constant, which can be easily measured using various methods112. Figure 6 shows the underlying mechanism followed during SPR sensing using the Kretschmann configuration.

**Figure 6.** Schematic of SPR-based detection of biomolecules using Kretschmann configuration. The dip in the resonance peak shifts due to the change in the refractive index.

Recently, new strategies, such as the generation of frictional charges using triboelectric surfaces, have been employed for the detection and generation of localized surface plasmon resonance (LSPR) wavelength shifts in large-area Au nanostructured surfaces [113]. These LSPs with frictional charges, using recent methods such as spectroscopic and triboelectric measurements, have been found to show a complex interplay of bioconjugation and biocomplex detection. These budding strategies in combination with GLAD-fabricated nanostructures can be employed for enhanced biosensing detection and applications. GLADfabricated metallic thin films facilitate the development of SPR- and LSPR-based biosensors, possessing the advantage of a considerable increment in sensitivity. These nano-sculptured thin films have been demonstrated, both theoretically and experimentally, to possess high SPR sensitivity [114,115]. Additionally, their periodicity, gap, etc., can be tailored in a desired way, hence the binding of bioanalytes and sensitivity enhancement [115].

#### 4.1.1. Bi-Material-Based SPR Enhancement

The properties of two different materials favorable for experimental sensing studies can be exploited together through the fabrication of biomaterial GLAD nanostructures. Additionally, using biomaterials also helps to eliminate the unfavorable properties of either of the two materials; for example, silver is highly plasmonic and provides a sharp SPR curve with high sensitivity but becomes easily oxidized. The oxidation of silver can be eliminated if a thin layer of some inert material is used over its surface. Graphene, possessing high impermeability [116] for gases and easy adsorption [117] ability for biomolecules, can be used to coat Silver to provide high SPR sensitivity. Wu et al. reported an SPR-based graphene biosensor in which the graphene layer was coated over a gold sheet. The sensitivity of the chip was shown to increase several times through calculations compared to the conventional gold film used for SPR biosensing. In another work, Choi et al. [118] presented a numerical model showing a layer of graphene coated over a 60 nm silver film. The SPR sensitivity was shown to increase several times due to the graphene coating and also to prevent the oxidation of silver. Over the graphene surface, a bimolecular DNA hybridization was modeled as a thin dielectric layer. The group showed improvement in the SPR bimolecular sensitivity of the graphene-coated substrate compared to conventional gold film chips employed for SPR biosensing applications. Graphene adds to the SPR sensitivity in comparison to the gold film due to its better adsorbing properties of biomolecules; this is due to the pi interactions between the biomolecules' carbon ring structure and graphene hexagonal cells [119,120]. Gold/zinc oxide (Au/ZnO), as well as gold/chromium (Au/Cr)-based nanocomposite thin films, have also been studied and were used effectively for enhancing SPR performance for the detection of a carbohydrate antigen, which is a tumor marker for breast cancer. The linear range for the SPR sensor was shown to increase from 1 to 40 U/mL [120]. Therefore, bimetal GLAD structures can be used to increase SPR-based bio-sensitivity along with increasing the inertness of the metallic structures.

#### 4.1.2. Nanocolumnar Thin-Film-Based SPR Chips

As discussed in the earlier section, thin-film nanocolumns fabricated using the GLAD technique provide high sensitivity for the detection of bioanalytes. They are easy to fabricate, possess a high surface –volume ratio, require less material, and are broadly used as SPR chips. The use of these metallic nanocolumnar thin-film biosensors enhances both angular as well as spectral sensitivity due to increased roughness and porosity, and induces SPR broadening due to scattering losses inside the porous metallic film [120,121]. Kaur et al. [122] reported ZnO thin-film-based SPR detection for cholesterol biosensing applications. ZnO was coated over the gold film using the GLAD technique. ZnO/Au prism was used in the Kretschmann configuration. The cholesterol oxidase (ChOx) enzyme had been immobilized over the ZnO surface. PDMS microchannels fabricated over the ChOx/ZnO/Au/prism setup were used to pass different cholesterol concentrations, and SPR reflectance was obtained for both static and running mode, showing high SPR sensitivity. Figure 7a shows the SPR reflectance curves for various cholesterol concentrations over the ChOx/ZnO/Au prism, and Figure 7b shows the calibration curve for the variation in the resonance angle with cholesterol concentrations. However, in most of the work discussed above employing SPR detection using thin metallic films, the sensor chips were fabricated as planar thin films using GLAD-based techniques. The experimental use and applications of GLAD-based columnar thin films using different structural morphologies are still limited for biosensing applications. The effect of porosity on the SPR dip for different metallic nano-columnar thin films (Al, Ag, and Au) fabricated via GLAD using Kretschmann configuration has also been studied. The SPR dip widens with an increase in the porosity of the thin film, and finally, disappears near the TIR region. This group showed the vast possibility of SPR-based biosensors utilizing the porosity of the metals with increased sensitivity compared to nonporous films.

**Figure 7.** (**a**) SPR reflectance curve for ChOx/ZnO/Au prism corresponding to different cholesterol concentrations. (**b**) Calibration curve for variation in resonance angle with cholesterol concentrations. Reproduced with permission from [122].

Shalabney et al. [123] have explained this phenomenon very well both via theoretical calculations as well as experimentally. They carried out SPR measurements in Kretschmann configuration for GLAD-fabricated Ag nano-columnar thin films for different values of β and studied the SPR sensitivity in both angular as well as spectral modes. They indicated that for more porous films, the resonance wavelength shows a clear redshift of about 15 nm with respect to less porous films, and hence, higher sensitivity. Further, a theoretical calculation of the dielectric tensor of the material was carried out. There was an enhancement in the electromagnetic field intensity distribution as there was an increase in the volume interaction of the evanescent field in both the columnar void region and the sample area. These nanocolumnar thin films provide a high increment in sensitivity compared to planar metallic films. Agrawal et al. [114] presented a 2D nano-pillar array with nano-gaps for biosensing. They used rigorous coupled-wave analysis (RCWA) simulations to study the SPR sensitivity. The biosensor performance was studied in terms of sensitivity, characterized by spectral shift Δλ. The effect of polarization of the light incident was studied using simulations, which showed that nano-pillars' biosensing capability is independent of incoming light. The sensitivity of SPR chips can be most suitably defined in terms of differential reflectance amplitude (DRA) since the sensor performance also depends on reflectance contrast along with reflectance change. Agrawal et al. [56] also demonstrated, in another work, the increased sensitivity of the SPR detection of nanohole arrays of gold compared to planar gold film (Figure 8). The SPR sensitivity of the biosensor had been calculated in terms of DRA. They carried out RCWA simulations to study the effect of

variation in the diameter and period of the nanoholes. Byun et al. [124] experimentally confirmed the high SPR bio-sensitivity of a gold nanowire array compared to conventional gold film. They used an angular interrogation scheme to perform biosensor characterization and showed that their results were in agreement with RCWA simulated results. They fabricated gold nanowires with two different periods—200 nm and 500 nm—with each showing increased sensitivity of 44% and 31%, respectively, when evaluated using ethanol of varying concentrations.

**Figure 8.** (**a**) Schematic diagram of circular nanohole array over the gold film. P, D, h, and H represent array period, array diameter, gold film thickness below the nanohole, and gold film thickness without the nanohole, respectively. This array is further coated with a 1 nm thick layer of biomolecules. (**b**) Reflectance spectra calculated for a planar gold film of thickness 60 nm and for gold film structure array with a period (P) and diameter (D) of nanohole array of 50 nm and 30 nm, respectively; gold film thickness without nanohole (H) of 60 nm; and gold film thickness below nanoholes of 40 nm. Reproduced with permission from [56].

Combining the two different ideas discussed above, one may suggest the fabrication of nanocolumnar thin films consisting of GLAD-engineered structures of different morphologies combined with other materials (ZnO, graphene, and gold) to increase bio-sensitivity. This may lead to us to ultra-high sophisticated SPR biosensors with multiple-fold sensitivity increments.

#### 4.1.3. Microfluidics-Based SPR Chips

Microfluidic microchannels and systems are mainly employed for the SPR imaging approach. The advantages of the microfluidic platform are: reduced nonspecific binding and detection time minimization by studying multiple analytes or ligand interactions using a single device [105]. SPR-based microfluidic assays have been developed for various biomolecular interactions such as antibody-binding targets [125], multi ligand/analyte biosensing [126], and foodborne pathogen detection [127]. Malic et.al. [128] reported a multichannel droplet-based SPR detection platform comprising electrowetting on a dielectric microfluidic device for SPR imaging. The device was fabricated using Cr and Au layers as SPR supporters followed by the patterning of 24 um photoresist as a lift-off mask. A 50 nm Au film was etched to define a thin Au layer as ground. So, a spot-by-spot detection achievement was employed. Natarajan et al. [129] reported a 3D microfluidic system for protein sample deposition to discrete spots of a target. They showed exposure of each spot to a larger sample volume due to continuous flow and increased uniformity of the spots, and hence, expanded the protein array applications for protein printing. A compact and robust microfluidic flow cell for SiO2 film followed by a PDMS microfluidic surface has been developed [130]. This biosensor flow has the advantage of operation up to 185 kPa in an aqueous environment. A unique droplet-based SPR biosensing platform coupled to a surface acoustic wave microfluidic system (SAW) has also been reported [131]. Surface acoustic wave excitation in continuous mode for different levels of power generation was studied. Droplet-based microfluidic SPR systems possess the advantage of limiting mass transport in comparison to conventional microfluidic flow.

Though most microfluidic systems focus on on-chip fabrication for SPR imaging, the major challenges of cost-effectiveness and the simplicity and compatibility of the detection and monitoring of SPR signals still require further effort and research. GLADfabricated arrays and chips can be very promising candidates for highly sensitive SPR imaging purposes.

#### *4.2. SERS-Based Biosensing*

Surface-enhanced Raman scattering (SERS) is the enormous amplification of Raman signals from molecules, by several orders of magnitude, when adsorbed on metal colloidal nanoparticles or a rough metal surface, discovered in 1974 by Fleischmann et al. [132,133]. The detection of analyte molecules even at a single-molecule level can be realized by employing the SERS technique, which expands its practical applications [134]. SERS has found applications in various fields, including the detection of trace chemicals [135] such as dye molecules, food additives, [136] and pesticides, [77] and in bioanalysis, medical diagnosis, [36] the detection of biomolecules, [137] cancer diagnosis, [138] in vivo molecular probing in live cells, [139] and explosives detection [140]. SERS substrates are the nanostructured platform that supports plasmon resonance and amplifies Raman signals [141–143] and are broadly classified as random morphology or ordered/periodic SERS substrates. Random morphology SERS substrates (which include roughened electrodes, metallic silver and gold colloids, and metal-island film on a planar substrate) are inhomogeneous and are not highly reproducible [144]. Periodic arrays of metallic nanostructures (using nanolithography and other physical vapor deposition techniques) can overcome this issue by providing uniform and controlled morphology of SERS substrates [145].

SERS has become a powerful technique for chemical and biological sensing applications due to its excellent sensitivity and specificity. Uniform and highly reproducible SERS substrates with batch-to-batch reproducibility that have a SERS signal variation of less than 15% can be fabricated using the GLAD technique [146]. Using SERS multivariate statistical analysis, various pathogen species or even strains can be differentiated. Therefore, extrinsic detection using Raman labels is more widely explored.

#### 4.2.1. Detection of Pathogens