Optical Biochemical Sensors Based on 2D Materials

C H A P T E R

10 Optical Biochemical Sensors Based on 2D Materials B.N. Shivananju1,2, Hui Ying Hoh3, Wenzhi Yu2 and Qiaoliang Bao2 1

State Key Laboratory of Applied Optics, Changchun Institute of Optics, Fine Mechanics and Physics, Chinese Academy of Sciences, Changchun, Jilin, P.R. China 2Department of Materials Science and Engineering, ARC Centre of Excellence in Future Low-Energy Electronics Technologies (FLEET), Monash University, Clayton, VIC, Australia 3College of Electronic Science and Technology and Key Laboratory of Optoelectronic Devices and Systems of Ministry of Education and Guangdong Province, Shenzhen University, Shenzhen, P.R. China

10.1 INTRODUCTION An optical sensor is a device that converts a physical stimulus (strain, pressure, thermal, electrical, magnetism, light, ultrasonic, touch, motion, sound, or biochemical molecules) into an optical output for reading or further processing [1,2]. However, compared to electrical sensing devices [3], optical sensing devices (based on prisms, microscope, spectroscopy, optical interferometers, waveguides, fibers, photonic crystal fibers, and optofluidic) have many desirable advantages, such as ultra-sensitivity, real-time monitoring, no electrical interferences, multiwavelength analysis (i.e., simultaneous response to different analytes), multifunctionality, long-term stability, lightweight, cost-effectiveness, remote measuring capability, lab-on-fiber capability, and in-vivo biochemical sensing applications [4 9]. In optical sensing, critical improvements involve increasing the specificity of label-free sensing and lowering the limit of detection (LOD) [7,8]. Researchers have

Fundamentals and Sensing Applications of 2D Materials DOI: https://doi.org/10.1016/B978-0-08-102577-2.00010-5

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explored optical sensors for various sensing applications such as refractive index [7], temperature [10], light [11 13], strain [14], pH [15], biochemical [16], gas [17], and humidity [18] using a range of nanomaterials [19]. In the last decades, studies on utilizing nanomaterials, including nanowires, nanoparticles, quantum dots, carbon nanotubes (CNTs), metal oxides, and polymers for optical sensing applications have been reported [20]. In comparison, 2D nanomaterials such as graphene, transition metal dichalcogenides (TMDs), topological insulators (TIs), boron nitride (BN), perovskite, and black phosphorus (BP) have attracted significant attention for optoelectronic and photonic devices [21 25]. The richness of the optoelectronic properties of 2D materials has encouraged the development of many optoelectronic and photonic devices, such as ultrafast lasers [26 28], photodetectors [29 32], modulators [33], polarizers [34], light-emitting diodes [35], and plasmonic devices [36,37]. However, we believe that the true potential of 2D materials lies in optical sensors, especially for biochemical sensing in the health-care sector, which could address some of the drawbacks of current sensor technology [38 48]. This book chapter comprehensively and critically reviews emerging optical biochemical sensors based on 2D materials. We first elaborate their optical sensing properties, followed by fabrication of 2D materialcoated optical sensors, biomolecule sensing applications (single cell detection, DNA sensing, and protein sensing), chemical sensing applications (gas sensing, humidity sensing, and ions sensing) and health-care applications (cancer diagnosis, optogenetics, and ophthalmology).

10.2 BIOCHEMICAL OPTICAL SENSING PROPERTIES OF 2D MATERIALS 2D materials have exceptional biochemical optical sensing properties. First, due to the atomic-thin layer structure and large surface area, they are excellent substrates for adsorption of biomolecules via π π stacking. Second, the large surface-to-volume ratio allows high-energy transfer efficiency and fast response time due to ultrafast carrier mobility. Moreover, the 2D materials of interest possess excellent biocompatibility, exceptional fluorescence-quenching ability, broadband light absorption, high chemical stability, outstanding robustness and flexibility [1,48]. Therefore, 2D materials have become widespread in biochemical sensing, diagnostics, and health-care applications [1,48]. The most striking features of optical biochemical sensors based on 2D materials are ultra-sensitivity and ultrafast response, thereby posing the potential to replace some of the current electrical sensors used for biochemical sensing applications [1,46].

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FIGURE 10.1

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Biochemical sensing principle of 2D materials coated optical sensors.

Fig. 10.1 shows the basic mechanism of biochemical sensing based on 2D material-coated optical sensor. When biochemical molecules come into contact with the 2D material, the Fermi level of the 2D material will shift either to p-type or n-type, changing its optoelectronic properties [1]. Thus the conductivity of the 2D material can be easily varied by biochemical doping, which is a very important attribute for any biochemical sensing using optoelectronic and photonic devices. The interaction between the biochemical molecules and the surface of the 2D material results in a change in the electron-hole carrier density of the 2D material, which in turn changes the local refractive index of 2D materials as shown in Eq. (10.1) [1]:   σg;i iσg;r 1=2 1 n2DM 5 2 ; (10.1) ωΔ ωΔ where σg,r and σg,i are the real and imaginary parts of the conductivity of a 2D material, ω is the frequency of light, and Δ is the thickness of 2D material. The optical output (wavelength, intensity) of 2D materialcoated optical device (neff) is determined by the refractive index of 2D material (n2DM), which in turn depends on the interaction of biochemical molecules on the surface of the 2D material [1].

10.3 FABRICATION OF 2D MATERIALS OPTICAL SENSORS 2D material based optical sensors can be fabricated in three ways: (1) transferring chemical vapor deposition (CVD) grown 2D materials onto optical sensors, (2) drop-casting 2D materials solution onto optical sensors, and (3) direct CVD growth of 2D materials onto optical sensors.

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10.3.1 Transferring CVD-Grown 2D Materials Onto Optical Sensors Fig. 10.2 shows the various steps involved in transferring CVDgrown 2D material (graphene) onto optical sensors (microfiber) [49]. First, using CVD a single layer of graphene is grown on the top surface of copper (Cu) foil. Then, poly(methyl methacrylate) (PMMA) polymer is spin-coated at 4000 revolutions per second for 40 s onto the top surface of the graphene/Cu foil, forming a PMMA/graphene/Cu sandwich-like structure. The PMMA polymer film acts as a shielding layer to avoid cracking of the graphene film until it is safely transferred onto the optical sensors [49]. The bottom layer of Cu foil is completely etched from the PMMA/graphene/Cu heterostructure by exposure to 1 M ferric chloride (FeCl3) solution for 60 min. The PMMA/graphene film is washed three to four times with deionized (DI) water to remove any Cu foil residues, then transferred carefully onto the optical sensors, such as optical prisms, waveguides, fibers, etched fibers, D-shaped fibers, tip of fiber connectors, microfibers, photonic crystal fibers, and fiber Bragg gratings (FBGs) [1]. The top layer of the PMMA polymer is next removed by treating it with acetone [49]. Finally, a nanosecond (ns) laser beam through a tapered fiber tip is employed to cut the extra graphene. Upon lifting the optical fiber, the graphene film is wrapped over the optical fiber to form a graphene-coated optical sensor [49].

FIGURE 10.2 The fabrication process of graphene-based optical sensors. Steps depict the transfer of CVD-grown graphene onto optical sensor [49].

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10.3.2 Solution Method of Coating 2D Materials Onto Optical Sensors The drop-casting method of coating 2D material solutions onto optical sensors (tilted fiber Bragg grating [TFBG]) is shown in Fig. 10.3 [50]. The liquid-phase exfoliation method is used to prepare black phosphorous (BP) nanosheets solution and these sheets are then coated onto the optical sensor [50]. The synthesis of BP nanosheets solution is as follows: (1) 25 mg bulk BP crystal was cut and ground into small pieces, which were then added to 25 mL absolute ethanol; (2) the BP dispersion was sonicated by a cell crasher at 25 kHz and 1200 W for 3 h to break the weak van der Waals stack of BP, where the temperature of solution is kept below 277 K with an ice bath; (3) the as-prepared BP dispersion is centrifuged at 5000 rpm for 15 min; (4) the supernatant containing few-layered BP nanosheets is decanted gently [50]. Finally, a BP nanosheets solution is coated onto the optical TFBG sensor using in situ layer-by-layer (i-LbL) technique, which is actually a chemical surface modification [50]. The sequence of the i-LbL technique is as follows: (1) An optical TFBG sensor is rinsed with acetone to remove any impurities from the sensing surface. (2) An alkaline treatment is carried out by dipping TFBG sensor in 1.0 M NaOH solution for one hour to enrich the number of 2 OH groups on the TFBG sensing surface. (3) The sensor is rinsed with DI water and ethanol. (4) A silanization procedure is carried out by incubating the TFBG in freshly prepared 5% (3-aminopropyl) triethoxysilane (APTES) for 2 h to form Si-O-Si bonding. (5) The TFBG sensor is placed into a microfluidic channel where 30 μL BP nanosheet solution is carefully drop-cast onto it. (6) The negatively charged BP nanosheets adhere to positively charged amino groups on the APTES-silanized TFBG sensing surface due to

FIGURE 10.3 Schematic representation of the synthesis of BP nanosheets (A D) and deposition process of BP nanosheets an optical TFBG sensor (E G) [50].

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electrostatic force; (7) Once the ethanol solvent has fully evaporated, the TFBG sensor is dried and coated with BP nanosheets. (8) The above procedures were repeated until the desired thickness of BP nanosheets on the surface of optical TFBG sensor is obtained. (9) The real-time BP deposition on TFBG surface can be monitored by capturing the transmission spectra of BP-TFBG sensor. (10) The optical BP-TFBG sensor is dried and stored in vacuum before further use for biochemical sensing applications [50].

10.3.3 Direct CVD-Grown 2D Materials Onto Optical Sensors Until now there has been no report on the direct CVD growth of 2D materials on optical sensors or optical fibers. Recently, Shivananju et al. demonstrated the direct CVD growth of CNTs on optical FBGs sensors for various applications [11 13], and Jingyu et al. demonstrated the direct CVD growth of graphene on the solid glass [51]. The above two techniques demonstrated a straightforward method of coating optical sensors with 2D materials for various applications, including biochemical sensing. Fig. 10.4 shows a schematic diagram of the CVD approach for the direct growth of uniform and high-quality graphene film on optical FBG sensor. In this process, optical sensors or optical fibers are thoroughly cleaned with DI water, acetone, and ethanol before being loaded into a quartz tube which is placed inside a three-zone high-temperature furnace. The optical sensors inside a quartz tube are flushed with 500 sccm Ar to remove impurities before the temperature of the furnace is increased. The furnace is heated to the desired growth temperature of graphene and allowed to stabilize for approximately 10 min. Typical growth conditions in the quartz tube are optimized with a gas mixture of 500 sccm Ar, 50 sccm H2, and 15.5 2 26.5 sccm CH4, with growth temperatures of 1000 C 2 1100 C for 1 2 7 h [51]. Some optical components

FIGURE 10.4 Schematic diagram of direct CVD growth 2D material onto the optical sensor.

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or optical fibers cannot tolerate a temperature of 1000 C due to low melting points. In this case, we need to adopt low-temperature (400 C 500 C) CVD growth of graphene on optical sensors or optical fibers. The thickness and uniformity of graphene or 2D materials onto optical sensors can be controlled by varying the flow rate of gases, growth temperature, and growth time [1,51].

10.4 BIOMOLECULES SENSING APPLICATION 10.4.1 Single Cell Detection Single cell detection plays a very important role in diagnosis, where a single diseased cell, which differs from the normal cells, carries information about the illness [52]. So there is a need for accurate detection of single cells among a group of normal cells. Currently, flow cytometry and optofluidics methods are widely used for label-free real-time single cell detection. However, these methods suffer from inaccurate singlecell detection and they require high-energy laser source for detection, which may damage the cells [52]. Recently, Xing et al. demonstrated a graphene-based highly sensitive optical refractive index sensor (4.3 3 107 mV/RIU) with a high resolution of 1.7 3 1028, which was used for single cancer cell detection from a group of normal cells [52]. Fig. 10.5A shows the experimental setup used for the ultrasensitive flow sensing of a single cell detection, where a He-Ne laser (632.8 nm, 80 μW) is used as the input light source, adjusted to circularly polarized light by a polarizer and quarter-wave plate. This circularly polarized light is focused onto the graphene-based optical prism sensor (GOPS) platform at the center of a microfluidic channel using an objective lens. The inset of Fig. 10.5A shows a schematic diagram of the GOPS, which consists of a quartz sandwich structure on the prism, a polydimethylsiloxane (PDMS) microfluidic chip, 8.1 nm thick h-rGO, and cell flow. The width and height of the optical microfluidic channel are approximately 20 and 12.5 μm. The diaphragm is used to further improve the detection range by reducing the reflected polarized light beam spot to 1 μm, which passes through the polarization beam splitter (PBS) to separate the TM and TE modes, which are detected by a balanced photodetector [52]. In this single cell detection experiment, Jurkat cancer cells (1%) and normal lymphocytes taken from blood were detected by changes in voltage using the GOPS microfluidic device. The refractive index and size of the Jurkat cancer cells are significantly larger than normal lymphocytes cells. Fig. 10.5B demonstrates the distinct changes in the voltage amplitude signal with respect to time, as the Jurkat cancer cells and

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FIGURE 10.5

Graphene-coated optical prism for ultrasensitive flow sensing of a single cell. (A) Flow-sensing system for a single-cell setup. The inset provides the schematic of the graphene-coated optical single-cell sensor platform, which is a PDMS microfluidic chip/h-rGO/quartz sandwich structure on a prism for single-cell detection. (B) Discrete time-dependent changes in voltage corresponding to mixed lymphocytes and Jurkat cells as they roll across the h-rGO detection window. (C) Enlarged images of the panel for certain positions in which the voltage signals are clearly depicted. The high and low signals represent Jurkat cells and lymphocytes, respectively, as they roll across the optical detection window. (D) Microscopic images of Jurkat cells. (E) Microscopic images of lymphocytes; scale bar is 15 μm [52].

normal lymphocytes flow through the h-rGO optical microfluidic channel at a rate of B7 μL h 1. Fig. 10.5C clearly shows the high and low-voltage signals, which represent the Jurkat cancer cell and normal lymphocyte, respectively, as they pass over the h-rGO optical microfluidic channel, indicating the high resolution and sensitivity of the GOPS, with an ability to detect a single cancer cell among the group of normal cells. The high-voltage amplitude denotes the signal from the Jurkat

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cancer cell due to its larger refractive index and size, compared to the low-voltage amplitude signal from normal lymphocyte cells. Fig. 10.5D E shows the microscopic images, where the Jurkat cancer cells (Fig. 10.5D) are on average larger than the normal lymphocyte cells (Fig. 10.5E). The average sizes of cancer and normal cells are translated into an average refractive index that is accurately detected by the graphene-based optical refractive index sensor [52].

10.4.2 Deoxyribonucleic Acid Sensing Deoxyribonucleic acid (DNA) sensing is very important in diagnostics. The selective and fast detection of DNA molecules at ultralow concentration is vital in numerous fields, such as gene therapy, disease diagnostics, biomedical applications, point-of-care clinical analysis, food quality, safety, and environmental monitoring. Currently, polymerase chain reaction (PCR) is widely used for DNA target and signal amplification [53]. Many new technologies such as silicon-nanowire sensors based field-effect transistor (FET) and surface-enhanced Raman scattering have been shown capable of direct detection of DNA at low concentrations [53]. Recently, 2D materials are being explored for DNA sensing applications, due to their large surface areas and unique optoelectronics properties. Researchers have already demonstrated that graphene and its derivative graphene oxide (GO) can be used for highly sensitive DNA hybridization detection down to a concentration level of few tens pM [53]. In 2014, Loan et al. reported a graphene-MoS2 heterostructure for ultrasensitive detection of DNA hybridization with the concentration level of attomolar (aM) based on optical photoluminescence (PL) [53]. A single layer of MoS2 exhibits high fluorescence-quenching ability, a desirable quality when using PL. The single layer graphene acts as a shielding layer from the surrounding environment to prevent the degradation of MoS2 and serves as a biocompatible interface layer to host DNA molecules on its surface [53]. Fig. 10.6A shows the schematic diagram of a PL experimental setup used for ultrasensitive detection of DNA hybridization based on graphene-MoS2 heterostructure [53]. A confocal optical microscope equipped with 473 nm laser source was used to probe the PL signal of the DNA molecules as they interact with the graphene-MoS2 heterostructure. The spatial PL mappings for the graphene-MoS2 heterostructure immobilized with the DNA solution (40 μL; 10 μM) and hybridized with the complementary DNA solutions (40 μL with various concentrations from 1 to 100 aM) is shown in Fig. 10.6B. The PL mapping intensity increases in the graphene-MoS2 heterostructure with the increase in the concentration of the added target DNA. The PL measurements for

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FIGURE 10.6 (A) The schematic illustration of the DNA detection method using an optical microscope and the graphene/MoS2 heterostructure sensor. (B) The photoluminescence (PL) peak area mappings of the graphene/MoS2 heterostructure hybridized with the complementary target DNA. (C) The PL spectra and integrated PL peak area in the presence of target DNA (1, 10, 100, and 1000 aM) [53].

the graphene-MoS2 heterostructure is carried out in a dry state, where the heterostructure was rinsed with DI water and dried after each probe DNA immobilization and target DNA addition. Fig. 10.6C summarized the integrated PL peak area over 1.7 1.95 eV for each condition, where we can clearly see a positive correlation between integrated PL peak area (in an arbitrary unit) and the concentration of the added complementary DNA from 1 aM to 1 fM. The PL color mapping clearly demonstrates graphene-MoS2 heterostructure responds to the target DNA with the detection limit of aM concentration [53].

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10.4.3 Protein Sensing The biomedical industry needs highly sensitive, ultrafast and compact biosensors that allow protein sensing to be monitored in real time. Methods such as enzyme-linked immunosorbent assays (ELISA), mass spectrometry, radial immunodiffusion, and western blotting have been reported for protein sensing applications [54]. Currently, there is a significant interest in the development of label-free protein sensing based on optical fiber methods. Fig. 10.7A shows a schematic of an optical fiber sensor for C-reactive protein (CRP) sensing, in which anti-CRP antibody (aCRP)-GO is coated on an etched fiber Bragg grating (EFBG) [54]. CRP is an indicator for cardiovascular diseases and chronic inflammations and plays a major role in diagnosing a patient. Currently, various techniques are used to detect CRP, such as the standard ELISA (0.3 10 mg L 1), quartz crystal microbalance (0.13 μg L 1 5.01 mg L 1), electrochemistry (0.1 50 mg L 1), or surface plasmon resonance (SPR, 2 5 mg L 1) [54]. A GO-coated EFBG device operates in the clinical range of 0.01 100 mg L 1 and work by monitoring the shift in the Bragg wavelength (ΔλB 5 2neffΛ, where neff is a function of ncore  1.47 and nclad 5 nGO  1.7, and Λ  532 nm, which is the grating period) as a function of concentration of CRP, as shown in Fig. 10.7B. It was demonstrated that the shift in the Bragg wavelength for aCRP-GO-coated EFBG is Bfive times higher than for aCRP without GO. This result verifies the sensitivity enhancement due to the presence of GO, which increases the binding between aCRP molecules and the surface of EFBG. Cross-sensitivity was evaluated, along with CRP detection, by introducing different interfering compounds, such as glucose (4000 mg L 1), urea (2000 mg L 1), and creatinine (6000 mg L 1), showing there is no significant shift in the Bragg wavelength. This work demonstrates the specificity of aCRP-GO complex-coated EFBG for CRP detection (0.01 100 mg L 1) in the presence of other compounds (glucose, urea, and creatinine), which is comparable to other standard methods reported previously [54]. The above results show the potential of 2D material based optical sensors for protein sensing applications.

10.5 CHEMICAL SENSING APPLICATIONS 10.5.1 Gas Sensing The sensing and measurement of gas molecules such as CO2, CO, NO2, NH3, H2, and O2 at room temperature are important for both understanding and monitoring a variety of occurrences such as

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FIGURE 10.7 (A) Schematic image of GO-coated EFBG sensor [45]. (B) The shift in the Bragg wavelength as a function of CRP concentration in (i) aCRP (purple open circles) and (ii) aCRP GO complex (olive green squares) coated EFBG sensors. Inset: Shift in the Bragg wavelength as a function of the concentration of only interfering factors such as urea (blue open squares), glucose (magenta open circles), and creatinine (red open stars) [54].

industrial processes, environmental changes, and hospitals [1,17]. Over the past few decades, researchers have demonstrated gas sensing at low concentrations (ppm) using various nanomaterials, such as CNTs, nanoparticles, nanofibers, nanosheets, and quantum dots [17]. Recently, 2D materials based gas sensors are showing superior performance in terms of ultralow concentration (ppm or ppb) detection and ultrafast

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sensing, arising from their atomic-thin layer structure, large surface area, large surface-to-volume ratio, large gas-adsorption capacity, fast response time to dynamic gas molecules due to its ultrafast carrier mobility. Fig. 10.8A shows a schematic diagram of the experimental setup of graphene-based 2D material-coated D-shaped optical fiber for multiple gas (NH3, H2O, and xylene) sensing applications [55]. A tunable laser (1510 1590 nm) with a power of 12 dBm was used as the input light

FIGURE 10.8 (A) Schematic diagram of the experimental setup of graphene-coated D-shaped fiber (GDF) for gas sensing. Spectral shifts of the GDF exposed in (B) NH3 gas, (C) H2O vapor, and (D) xylene gas. (E) GDF’s sensitivities for three types of gas molecules [55].

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source, which propagates through the graphene-coated D-shaped fiber (GDF) placed inside a sealed gas chamber, and the output transmitted light is collected by using an optical spectrum analyzer (OSA). When gas molecules (NH3, H2O, and xylene) come into contact with graphene, the permittivity of the graphene film will vary, causing a change in the refractive index of the sensing region of the optical D-shaped fiber, which in turn results in a shift in the output transmission spectrum. Fig. 10.8B D shows the transmission spectrum shift of GDF for different gas molecules (NH3, H2O, and xylene). The gas molecules were injected into the gas chamber with varying concentrations of 0, 200, 500, and 1000 ppm. GDF shows a maximum sensitivity of B3 pm ppm 1 for NH3, B1 pm ppm 1 for H2O, and B0.6 pm ppm 1 for xylene gas molecules, respectively (Fig. 10.8E). It is found that the sensitivity of GDF is 10 times better than without graphene [55]. These results forecast the potential of 2D materials coated optical sensors for ultrasensitive and highly selective gas molecules sensing applications.

10.5.2 Humidity Sensing Humidity sensors are used to measure and monitor the amount of water present in the surrounding air. These sensors are widely used in industries such as semiconductor, biomedical, textiles, food processing, pharmaceuticals, meteorology, microelectronics, agriculture, structural health monitoring, and environment monitoring [18]. Materials such as metal oxides, polymers, hydrogel, nanoparticles, and CNTs are explored for humidity sensing [18]. Recently, researchers are trying to explore 2D materials such as graphene, BP, and TMDs for humidity sensing applications [56 58]. These 2D materials have a high surface-to-volume ratio which supports a high sensitivity to variation in humidity. Various sensing techniques such as electronic (resistive, capacitive), thermally conductive, gravimetric, and optical methods are available for humidity sensing. However, optical sensors for humidity sensing gain popularity due to many advantages such as compact size, lightweight, inexpensive, real-time monitoring of humidity in hazardous environments, and remote humidity sensing capability [56]. Luo et al. have demonstrated a novel 2D Tungsten disulfide (WS2) material-coated optical side-polished fiber (WS2CSPF) for use as a highly sensitive and fast response humidity sensor [56]. Fig. 10.9A shows the experimental setup which consists of a 1550 nm laser source, 1 3 2 coupler, humidity chamber (humidity adjusting range: 35%RH to 95%RH and temperature adjusting range: 210 C to 100 C), and optical power meters or OSA. A commercial humidity/temperature meter is inserted into the chamber to monitor the actual humidity and

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FIGURE 10.9 (A) Experimental setup for humidity sensing based on 2D tungsten disulfide (WS2) material-coated optical side-polished fiber (WS2CSPF). (B) Variation of relative output optical power through WS2CSPF. (C) Relative output optical power of WS2CSPF as a function of relative humidity and SPF’s data [56].

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temperature. An input laser (1550 nm) is coupled into an unpolished single-mode fiber to monitor the power fluctuation of the laser. The output laser power of SPF (control measurements) and WS2CSPF (humidity sensing) are monitored using an optical power meter or OSA. The relative humidity (RH) inside the chamber was first increased from 35%RH to 85%RH in steps of 5%RH and later decreased in similar steps. The response of the WS2CSPF optical sensor has been investigated by exposing the sensor sequentially to a range of different humidity conditions at a constant room temperature (25 C). Fig. 10.9B shows the performance of the WS2CSPF optical humidity sensor to varying humidity from 35%RH to 85%RH in steps of 5%RH, both increasing and decreasing order. The basic mechanism of humidity sensing based on WS2CSPF optical sensor is as follows: whenever humidity is increased in the chamber, the concentration of H2O molecules increases. These molecules interact with the 2D WS2 material and charge transfer takes place from WS2 to H2O molecules. According to orbital mixing theory, the conductivity of WS2 material decreases due to the reduction of major conducting electrons following a humidity increase, which in turn decreases light absorption. The transmitted loss of an SPF coated with WS2 film decreases while the transmitted optical output power increases. Hence the humidity sensing function enhancement can be achieved by using an SPF coated with a 2D WS2 material, as seen in Fig. 10.9C, which shows the relative output power of WS2CSPF optical humidity sensor as a function of RH in comparison with SPF optical sensor without the 2D WS2 material. This novel WS2CSPF optical humidity sensor can achieve a linear correlation coefficient of 99.39%, a sensitivity of 0.1213 dB/%RH, and a humidity resolution of 0.475%RH [56]. It is clear from these data that the sensitivity of a WS2CSPF optical humidity sensor is enhanced about 15 times compared to that of the bare SPF without 2D WS2 material. Therefore, we can see that the presence of the 2D WS2 material greatly improves the sensitivity and stability of the optical SPF humidity sensor [56].

10.5.3 Heavy Metal Ion Sensing Heavy metals are metallic chemical elements (mercury, arsenic, thallium, chromium, cadmium, and lead) with a relatively high density and are usually toxic and carcinogenic. They are natural components of the earth’s crust which cannot be destroyed and degraded and enters into our bodies via air, drinking water, and food. Heavy metal (e.g., lead [Pb]) at a high concentration in drinking water can lead to toxic biochemical effects in humans which in turn cause problems in the kidneys, heart, nervous system, gastrointestinal tract, joints, and

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reproductive system. Cell functions such as the synthesis of hemoglobin can also be affected. According to guidelines of the World Health Organization (WHO), the maximum permissible limit of lead in drinking water is 10 ppb; exceeding this concentration will affect the human health. Heavy metal ion detection in drinking water is therefore of vital importance [50]. Recently, Liu et al. demonstrated ultrasensitive 2D material (BP) integrated tilted fiber grating (TFG) optical sensor for detecting heavy metal (Pb) ions with an LOD of 0.25 ppb [50]. Fig. 10.10A shows a schematic diagram of the experimental setup used for heavy metal sensing, where a broadband light source with TM polarized resonance was directed onto the optical BP-TFG microfluidic sensor and the output signal was monitored by using an OSA. Solutions with different concentrations of Pb21 (0.1, 1.0, 10, 100, 1000, 1 3 104, 1 3 105, 1 3 106 and 1.5 3 107 ppb) solutions were injected into the BP-TFG microfluidic channel for 120 s and then the solution was carefully withdrawn from the microfluidic channel, such that the BP-TFG optical sensor was exposed for 180 s. The

FIGURE 10.10 (A) Experimental setup for heavy metal chemical sensing. (B) Transmission spectra of BP-TFG showing a clear upshift as Pb21 ions concentration increases. (C) The resonant intensity of BP-TFG change with increasing Pb21 ions concentration [50].

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transmission spectrum was captured by using an OSA for each Pb21 concentrations. Before using the BP-TFG optical sensor for the next measurement (different concentration), it was completely washed with ethanol to remove the adsorbed Pb21 ions [50]. Fig. 10.10B shows the transmission spectrum shift of the BP-TFG optical sensor at different Pb21 concentrations. We can clearly see that by increasing Pb21 concentration, the transmission peak intensity decreases along with wavelength red-shift, indicating that strong optical absorption occurs between BP 2D material and Pb21 ions, which in turn changes the effective refractive index of TFG cladding. Fig. 10.10C shows the transmission spectrum intensity of BP-TFG optical sensor as a function of Pb21 concentration, revealing a nonlinear relationship, with ultra-sensitivity of 0.5 3 1023 dB ppb 1, 7.7 3 1027 dB ppb 1, and 2.3 3 1028 dB ppb 1 for Pb21 concentration ranges of 0B100 ppb, 103B105 ppb, and 106B107 ppb, respectively. These results show the potential of BP-TFG optical sensor for ultrasensitive detection of Pb21 over a wide range of concentrations, from 0.1 ppb to 1.5 3 107 ppb, which is a few orders of magnitude larger than that of other electrical based lead sensors [50].

10.6 HEALTH-CARE APPLICATIONS 10.6.1 Photothermal and Chemotherapy for Cancer Diagnosis Cancer is currently the deadliest disease among humans. Cancer cells grow and reproduce uncontrollably, creating tumors in a specific part of the body and spreading to other areas, where they attach to and destroy surrounding healthy tissues and organs and ultimately cause death. The spread of cancer from one part of the body to another is called metastasis. Today, we can identify distinct types of cancers based on where they start growing, such as breast cancer, lung cancer, kidney cancer, liver cancer, and each cancer requires different approaches for diagnosis and treatment [59,60]. One common diagnostic methods is optical imaging (fluorescence, molecular, and Raman spectroscopy) or biopsies, which give information on cancer initiation, progression, and metastasis. Treatments can include radiotherapy, chemotherapy, biological therapy, hormone therapy, stem cell transplantation, and surgery. However, most of these diagnosis and treatment methods are expensive, invasive, and inaccurate. Hence there is a need for cost-effective, minimally invasive, and accurate methods of detecting cancer. New approaches are possible only with the recent developments in the field of biosensors, especially using nanomaterials such as graphene and its derivatives [59]. GO and graphene-quantum dots (GQDs) can be used as fluorescence probes in

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PL imaging due to their biocompatibility, photostability, nontoxicity, and chemical inertness, which are desirable attributes for designing fluorescence-based biosensors. Moreover, graphene-based fluorescence probes have two vital functionalities in PL: specific detection of cancer cells and as therapeutic agents. Light-induced thermal therapy has been validated clinically for treating and curing cancer tumors. In photothermal therapy, near-infrared light penetrates the target human tissues with limited depth and provides safety delivery of drugs. However, simply using phototherapy treatment can result in the incomplete killing of cancer cells. Therefore, the combination of photothermal and chemotherapy is a more effective treatment. Chemotherapy can enhance the efficiency of photothermal therapy by targeting surviving cancer cells or by inhibiting regrowth of damaged tumor blood vessels. More recently, light-activated nanoparticles that release their payload in response to light irradiation have been developed, achieving improved drug bioavailability with superior efficiency [59,60]. Recently Liu et al. demonstrated effective drug delivery with PEGylated MoS2 2D nanosheets for combined photothermal and chemotherapy targeting cancer cells [59]. 2D MoS2-PEG nanosheets have very strong NIR wavelength absorption and are a promising candidate for photothermal therapy. The nanosheets also possess high surface area to mass ratio, enabling efficient loadings of therapeutic molecules, such as chemotherapy drugs doxorubicin (DOX), 7-Ethyl-10hydroxycamptothecin (SN38), and a photodynamic agent chlorin e6 (Ce6). Fig. 10.11A shows the schematic diagram of the fabrication of MoS2-PEG which is subsequently used for drug loading. 2D MoS2 nanosheets were synthesized by chemical exfoliation and then functionalized with lipoic acid modified PEG (LA-PEG) using a thiol reaction to increase their physiological stability and biocompatibility for drug loading. Fig. 10.11B F shows the MoS2-PEG nanosheets after loading with DOX, which can be used for combined photothermal and chemotherapy for in vivo cancer treatment. First, 1 3 106 murine breast cancer 4T1 cells (40 μL PBS) were injected into the back of Balb/C female mice. When the cancer tumor volume reached B50 mm3, these mice were randomly separated into five groups and intratumorally injected with 20 μL of PBS, DOX, MoS2 PEG, and MoS2-PEG/DOX ([DOX] 5 0.5 mg kg 1, [MoS2-PEG] 5 0.34 mg kg 1). Next, the infected mice were irradiated with NIR (808 nm, 0.35 W cm 2) for 20 min and temperature changes of these mice were monitored by using an IR thermal camera. In Fig. 10.11D, we can clearly see a significant change in temperature, from room temperature to 44 C 45 C in the case of mice injected with MoS2-PEG/DOX as compared to PBS and DOX. Mice with intratumoral injection of MoS2-PEG/DOX but without laser irradiation were also studied as the control experiment. After various treatments, the length

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FIGURE 10.11 (A) A scheme showing the fabrication process of MoS2-PEG and the subsequent drug loading. (B) A scheme showing in vitro targeted combination therapy with MoS2-PEG-FA/DOX. (C) Scheme of combination therapy based on intratumorally injected MoS2-PEG/DOX. (D) Infrared (IR) thermal images of 4T1 tumor-bearing mice recorded by an IR camera. The doses of DOX and MoS2-PEG were 0.5 mg kg 1 and 0.34 mg kg 1, respectively. Laser irradiation on the tumors was conducted by using 808-nm near-infrared (NIR) laser at a power density of 0.35 W cm 2 for 20 min. (E) The temperature change of the tumors was monitored by an IR thermal camera in different groups during laser irradiation as indicated in (C). (F) Tumor volume growth curves of different groups of mice after various treatments (five mice for each group) [59].

and width of the tumors were monitored every two days for the next 21 days with a digital caliper. The tumors injected with either PBS or free DOX grew quickly within 21 days, suggesting that free DOX at this low dose was not effective in inhibiting the tumor growth. However, the tumor growth in the group treated with MoS2 PEG/DOX and exposed to the NIR irradiation was dramatically inhibited after the combined photothermal therapy and chemotherapy. The above results show that DOX-loaded MoS2 PEG nanosheets used in the combined photothermal and chemotherapy can achieve an outstanding synergistic effect in inhibiting tumor growth in animal model experiments [59].

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10.6.2 Optogenetics Neuroscience is a multidisciplinary science that focuses on the study of the structure and function of the nervous or brain system. Optogenetics is a tool in which individual neuronal circuits in the brain are manipulated by light [61,62]. These circuits relate to different aspects of our behavior and personality as demonstrated in an animal module (Fig. 10.12A). Optogenetic tools play a vital role in health-care applications to resolve some of challenging neural disorders such as Parkinson’s, depression, addiction, autism, anxiety, schizophrenia, drug abuse, memory loss, spinal cord injury, eyesight loss, and stroke. Microelectrocorticography (micro-ECoG) neural interfaces are widely used to access the brain neurons externally to detect and record high-quality neuron signals from the brains of patients suffering from multiple neural disorders, to provide suitable therapy. The current micro-ECoG arrays are made up of indium-tin-oxide (ITO), which is not ideal for the neural interface due to many disadvantages such as brittleness, high-temperature processing required, inability to conform to the cortical surface, and most importantly limited light transmission at the ultraviolet (UV) and infrared wavelengths, which is a critical consideration in optogenetics and neural imaging. graphene- based micro-ECoG neural interfaces can overcome all these drawbacks in addition to its many advantages, such as biocompatibility, flexibility, high mechanical strength, broadband wavelength transmission, high thermal and electrical conductivity, and tunable optoelectronic properties [62]. Fig. 10.12B shows a graphene-based carbon-layered electrode array (CLEAR) device used for optogenetics and neural imaging. The CLEAR device consists of four-layer graphene with a minimum sheet resistance of 76 Ω per square, retaining broadband wavelength (300 1500 nm) with B90% transmission and also possessing mechanical strength superior to ITO and ultrathin metals. Fig. 10.12C shows a schematic diagram of an optogenetics experiment conducted on a mouse by placing a CLEAR device on the cerebral cortex of the mouse brain and stimulating the neurons by shining a 473 nm blue laser (100 mW) through a 200 μm optical fiber and recording in parallel the electrical output signal. From Fig. 10.12D we can clearly see the blue light stimulus being delivered through a CLEAR device embedded on the cortex of a Thy1:: ChR2 mouse. The basic mechanism of optogenetics is as follows, first, a light-sensitive protein (channelrhodopsin-2) is extracted from archaebacteria and algae; this protein produces an electrical current in the form of ions in response to blue (473 nm) light. The DNA extracted from this light-sensitive protein (channelrhodopsin-2) is inserted into specific neurons in the (Thy1::ChR2 mice) brain, and these neurons communicate by “firing,” that is, an electrical signal is created by opening and closing

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FIGURE 10.12 Graphene-based carbon-layered electrode array (CLEAR) device for neural imaging and optogenetics applications. (A) Optogenetics implemented in the animal module [61]. (B) The fabrication process of the CLEAR device. (C) Schematic drawing of the experimental setup, showing the CLEAR device implemented on the cerebral cortex of a mouse, with an optical fiber delivering blue light stimuli to the neural cells. (D) Image of a blue light stimulus being delivered via an optical fiber, through the CLEAR device implanted on the cortex of a Thy1::ChR2 mouse. (E) Optical evoked potentials recorded by the CLEAR device; x-scale bars: 50 ms, y-scale bars: 100 μV. (F) Maximum intensity projection (MIP) of OCT angiogram showing cortical vasculature visible through the CLEAR micro-ECoG device (FOV 2.8 3 2.8 mm2 and 1.1 3 1.1 mm2, respectively). (G) Doppler blood flow velocity image showing the directionality of blood flowing through the vasculature below the CLEAR device (FOV 2.8 3 2.8 mm2 and 1.1 3 1.1 mm2). Red represents blood flowing towards the lens, and green represents blood flowing away [62].

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ion channels. Finally, light-induced responses from the cerebral cortex of the mouse brain are recorded by the CLEAR device for three different light intensities with a stimulus time of 3 ms (Fig. 10.12E). Because of the high transparency of the CLEAR device in the infrared wavelength region, we can also clearly see the structure of the cerebral vascular 3D optical coherence tomography (OCT) angiogram image of the mouse brain (Fig. 10.12F). The typical velocity profile of the blood flow in the brain vessels is easily visible through the CLEAR device as shown in Fig. 10.12G. This study confirms that a 2D materials-based CLEAR device can be easily embedded on the brain surface for superior performance in neural imaging and optogenetics applications [62].

10.6.3 Ophthalmology In the last few years, many wearable devices have been developed based on flexible and stretchable nanomaterials, as well as advances in micro/nanofabrication, smart electronics, and information technology. Currently, researchers are showing extensive interest in wearable devices for biomedical applications such as smart contact lens [63 65]. These electronic wearable devices need flexible, biocompatible, and environmentally stable electrode materials for various sensing and display applications. The recent innovation in 2D materials such as graphene can play a vital role in wearable biomedical devices. Wearable eye contact lenses have been used for diagnosis of glaucoma and diabetes by measuring intraocular pressure and glucose composition of tears [63]. During this process, RF technologies can be employed for signal or power supply; however, electromagnetic wave interference (EMI) from surrounding wireless devices can affect the signal collected. Wearing contact lenses for a long period of time may also cause dry eye syndrome. Therefore there is a need to protect the eye from EMI and retain moisture using a diffusion barrier [63]. Recently, Lee et al. demonstrated a smart eye contact lens based on a CVD-grown graphene 2D material which acts as dehydration protection and electromagnetic interference shielding to prevent eye diseases such as cataracts [63]. Fig. 10.13A and B shows the schematic diagram of working principle of the EMI shielding graphene-based smart eye contact lenses. When the contact lens is not covered with graphene (Fig. 10.13A), the EM wave can easily pass through the contact lens and be directly absorbed by the eyeballs, which may cause thermal damage in the eyeballs leading to cataracts. Once the contact lens is covered by the graphene layer (Fig. 10.13B), the EM wave is partially absorbed by the graphene layer, avoiding thermal damage to the inner eyeballs. Fig. 10.13C shows experimental validation by exposing normal and

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FIGURE 10.13 The schematic and working principle of a graphene-coated contact lens. (A) Electromagnetic (EM) wave passes through contact lens and absorbed by an eyeball, possibly causing heat damage inside. (B) EM energy is absorbed by graphene and dissipated as heat before reaching the interior of the eye. (C) Sample preparation for the microwave oven test. Egg whites on an Si wafer are covered with normal contact lenses and lenses with a graphene coating, respectively. (D) Infrared (IR) camera images showing the elevated temperature of the graphene-coated lens inside a microwave oven, indicating the EM energy is efficiently absorbed and dissipated as heat. (E, F) Dehydration of a contact lens can be reduced due to gas-impermeability of graphene. (G) Schematic of the experimental setup to measure the water evaporation rate through contact lenses. (H) Weight loss measured with time on a hot plate at 38 C [63].

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graphene-coated contact lens to strong EM radiation (120 W for 50 s). This was done by placing them inside a microwave oven, which emits EM of similar wavelength as those from 4 G LTE and Bluetooth (2.45 GHz). The results clearly show that the thermal denaturalization, i.e., the color change of the egg white, shielded by a graphene-coated lens is considerably less than the case with the normal lens. When the graphene layer is exposed to EM radiation, the charge carriers (electrons) in orbital motion induce oscillating magnetic moments in response to the external magnetic field, which efficiently absorbs the EM energy and dissipates it as thermal energy. From Fig. 10.13D, we can clearly see the temperature change (captured by IR camera) from room temperature (B27 C) to above B45 C when strong EM (120 W for 20 s) irradiated the graphene-coated lens, while the temperature of the contact lens without graphene remains almost unchanged [63]. Fig. 10.13E F shows a schematic diagram of how prolonged wearing of contact lenses without a graphene layer leads to dehydration of the eyeballs, which may cause xerophthalmia (Fig. 10.13E), whereas graphene lenses offer protection from dehydration (Fig. 10.13F). The dehydration protection of the graphene-coated contact lens is verified by placing the normal and graphene lens on water-filled vials. The vials are next placed on a hot plate maintained at 38 C (Fig. 10.13G). After one week, the weight of the vial covered by graphene lens decreased by 0.5535 g, while the weight of the vial covered by normal lens decreased by 0.8268 g, showing that graphene acts as a dehydration protection material (Fig. 10.13H). Thus graphene-coated eye contact lens can provide dehydration protection and electromagnetic interference shielding for eyeballs and this serves as a bionic platform for wearable biomedical technologies in the future health-care applications [63].

Acknowledgment We acknowledge support from the National Natural Science Foundation of China (61875139), Shenzhen Nanshan District Pilotage Team Program (LHTD20170006) and Australian Research Council (ARC, FT150100450, IH150100006 and CE170100039). Q. Bao acknowledges support from the Australian Research Council (ARC) Centre of Excellence in Future Low-Energy Electronics Technologies (FLEET).

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