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# Recent advances in optical fiber grating sensors for detection of organic substances

Xiaoyu Zhong<sup>a,1</sup>, Quanhua Xie<sup>a,1</sup>, Yang Liu<sup>a</sup>, Yuanyuan He<sup>a</sup>, Nianbing Zhong<sup>a,*</sup>,
Zhouxiang Zhang<sup>b</sup>, Hassan Karimi-Maleh<sup>c,*</sup>, Xiaoling Peng<sup>a</sup>, Eric Lichtfouse<sup>d,*</sup>

## ABSTRACT

The continuous development of society has led to the flourishing of environmental chemistry and biomedical fields, as well as to the research and development of novel functional materials and sensors. Some novel materials have excellent optical and biological/chemical properties and can be coupled with the optical fiber grating. The coupling between the materials and optical fiber grating can improve the performance due to the advanced materials can amplify the changing signals generated during the reaction with analytes of organic compounds. Despite the great achievements in the development of bio- and chemical-optical fiber grating sensors, there is no publicly available systematic review of the recent advances in the use of materials combined with optical fiber grating in organic substances sensing available in the literature since January 2016. Thus, we summarized the sensing principles of the bio- and chemical-optical fiber grating with and without surface plasmon resonance. We discussed the performance parameters such as sensitivity, selectivity, detection range, and low limit of detection for detection of organic substances including organic pollutants, life organics (glucose, hemoglobin, cancer markers, antibody-antigen, biotin and streptavidin, DNA, enzymes), microorganisms, and other organic substances. We analyzed the advantages and limitations of optical fiber grating sensors and presented challenges in their practical applications and future research directions.

### Keywords:

Fiber gratings sensors

Bio- and chemical-sensors

Sensing principles

Organic substances detection

## 1. Introduction

Organic substances including organic pollutants, life organics (glucose, hemoglobin, cancer markers, antibody-antigen, biotin and streptavidin, DNA, enzymes), and microorganisms, are an essential part of the biogeochemical carbon cycle [1–3]. The presence of organic substances in nature can be detrimental or beneficial. For example, in the water industry, drinking water is most likely contaminated by sewage containing the organic pollutants and bacteria [4,5]. Alternatively, human health is closely related to the life organics, such as glucose, hemoglobin, DNA, and enzymes, in blood plasma or other biological fluids [6]. Therefore, to ensure safe drinking water and protect human health, various optical and electrochemical sensors have been developed and widely used in environmental monitoring,

biomedicine, food quality and safety, and other fields [6–10]. Among optical sensors, the performance [11,12] (such as sensitivity, limit of detection, response time, and specificity) of optical fiber gratings-based sensors has been proven to be comparable to the classic optical fiber sensors (e.g., photonic crystal optical fiber sensors [4], interferometric optical fiber sensors [5], U-shaped optical fiber sensors [6]). In addition, the optical fiber gratings have various features such as corrosion resistance, electromagnetic interference resistance, high compactness, miniaturization, multiplexing and remote measurement capability [13]. Currently, the most popular sensors for obtaining high sensitivity and low limit of detection (LOD) include etched fiber Bragg grating (eFBG) [14–20], tilted fiber Bragg grating (TFBG) [21–23], long period fiber gratings (LFPG) [24–26], surface plasmon resonance optical fiber gratings [27–29] and localized surface plasmon resonance optical fiber

<footnote><sup>a</sup> Chongqing Key Laboratory of Optical Fiber Sensor and Photoelectric Detection, Chongqing Key Laboratory of Modern Photoelectric Detection Technology and Instruments, Chongqing Engineering Research Center of Intelligent Optical Fiber Sensing Technology, Chongqing University of Technology, Chongqing 400054, China</footnote>

<footnote><sup>b</sup> University of Electronic Science and Technology of China, Chengdu, Sichuan 611731, China</footnote>

<footnote><sup>c</sup> The Quzhou Affiliated Hospital of Wenzhu Medical University, Quzhou Peoples Hospital, China</footnote>

<footnote><sup>d</sup> State Key Laboratory of Multiphase Flow in Power Engineering, Xi'an Jiaotong University, Xi'an, Shaanxi 710049, China</footnote>

<footnote>* Corresponding authors.</footnote>

<footnote>E-mail addresses: zhongnianbing@163.com (N. Zhong), hassan@uestc.edu.cn (H. Karimi-Maleh), eric.lichtfouse@icloud.com (E. Lichtfouse).</footnote>
gratings [30,31].

With the development of materials science and biotechnology, functional and bio-recognition materials with specificity, high surface area, tunability, and quantum confinement have been developed rapidly [32–37]. The sensitive materials can be assembled on the surface of optical fiber gratings by direct growth, deposition, and coating. Therefore, the combination of optical fiber gratings with biochemical molecular sensitive-materials has been successfully developed for various bio- and chemical-sensors with amazing results such as higher sensitivity, better selectivity, and better low LOD than those of optical fiber gratings [38]. The advantages of the functionalized optical fiber gratings are due to the facts as follows. Selective biochemical reactions occur between sensitive-materials attached to the surface of optical fiber and target analytes, leading to strong changes in the refractive index of the fiber surface [39]. The change in refractive index will result in a spectral change in optical fiber gratings [40]. By detecting the degree of spectral change of optical fiber gratings with an analytical instrument, quantitative detection of the target analyte can be achieved. [41].

2D nano-materials such as MXene [42,43], graphene [44], and derivatives of graphene including graphene oxide [21,24] and reduced graphene oxide [45] as well as metal-organic-framework [46]-coated optical fiber gratings-based sensors have been used to monitor the organic substances. Metal oxides (Ta₂O₅ [25], Cu₂O [47], TiO₂ [48], metal nano-particles and nano-films (Au [28], Ag [49], Au nanoparticles [31]) or a mixture of several materials (TiO₂/SiO₂ [50]) are also used as a signal enhancement element on optical fiber gratings surfaces to optimize refractive index' sensitivity as well as to excite surface plasmon resonance or located surface plasmon resonance for detection of organic substances [25,51]. Furthermore, optical fiber gratings-based sensors assembled with the recognition elements such as antibody [52–54], nucleic acid [55,56], aptamer [23,29,44,57–60], biotin/streptavidin [48,61,62], and enzymes [21,63–65] can selectively detect the life organics. Therefore, optical fiber gratings combined with materials science and biotechnology have made great strides in the recent years. However, the last comprehensive review of optical fiber gratings-based biosensors modified by functional material and biorecognition molecules was published in 2017 [51]. Furthermore, the optical fiber gratings-based chemical sensors were not included in the reported review [51] and has not been systematically reviewed in the past 8 years. In 2023, Du et al. wrote a review of biosensors based on LPFG, but his focus was only on LPFGs and was limited to the field of biosensing [38]. Li et al. reviewed the most common fiber-optic biosensors but do not cover all organic substances [66]. In addition, there are a number of researcher reviews that essentially do not discuss fiber gratings [67] or only discuss a particular type of optical fiber gratings [68–71] or are based only on plasmon resonance principles [72–74] or only include biosensors [75–77].

Here we review optical fiber gratings-based bio- and chemical-sensors modified by materials and biorecognition molecules for organics analysis available in the literature since January 2016. We first discussed the working principle of optical fiber gratings (LPFG, eFBG and TFBG) with and without surface plasmon resonance effect. Subsequently, we presented the sensitivity, selectivity, detection range, response time, and low LOD of optical fiber gratings-based bio- and chemical-sensors for detection of organic pollutants, life organics, microorganisms, and other organic substances. Finally, research results and perspectives of optical fiber gratings-based sensors modified with materials and biorecognition molecules for the detection of organic substances are discussed.

## 2. Working principle of bio- and chemical-optical fiber gratings-based sensors

In general, the sensing principle of bio- and chemical-sensors is based on the selective interaction/reaction between the target analytes and recognition elements on the surface of the optical fiber gratings. The

selective interaction/reaction includes selective adsorption and permeation [78], specific binding of antigen-antibody [79–81], specificity of enzyme to substrate [21,65], specific binding of biotin and streptavidin [48,61], and so on. The results of interaction/reaction are that causes a change in the effective refractive index around the fiber. The optical fiber gratings-based sensors with and without surface plasmon resonance and located surface plasmon resonance can effectively respond to changes in refractive index, thereby achieving quantitative detection of target analytes. Herein, the working principles of bio- and chemical-optical fiber gratings-based sensors with and without surface plasmon resonance and located surface plasmon resonance for detection of object analytes are discussed in detail.

### 2.1. Optical fiber gratings without surface plasmon resonance effect

The standard optical fiber gratings consist of a series of periodic transitions in the refractive index (n) of a fiber core, forming a modulation pattern extending along the fiber axis. This diffraction structure satisfies the phase matching condition between the propagation basic core mode (LP₀₁) and the co-propagation cladding mode. These modes can be back propagation core modes, cladding modes, or radiation (or leakage) modes [51].

Optical fiber gratings can be classified into two types based on grating pitch Λ (i.e., refractive index modulation period), namely short- and long period fiber gratings, i.e., fiber Bragg grating, TFBG and LPFG. For fiber Bragg grating, the grating pitch is typically on the order of hundreds of nanometers and leads to coupling between the fundamental core mode and its corresponding back-propagation mode [82]. This coupling occurs at a defined wavelength, which is called the Bragg resonance wavelength ($\lambda_{res}^{FBG}$). When the light signal reaches the grating area, light that meets the phase matching condition will be reflected, however, the light that does not meet phase matching condition will continue to travel along the fiber. Thus, the grating selects light of a specific wavelength according to Bragg conditions and acts like a reflector or suppression filter:

$$ \lambda_{res}^{FBG} = 2n_{core}^{eff} \Lambda \quad (1) $$

where $n_{core}^{eff}$ is the effective refractive index of the core mode. In the Eq. (1), it can be seen that the Bragg wavelength is affected by changes in the physical or mechanical properties of the grating region. For example, the strain on the fiber changes Λ and $n_{core}^{eff}$ due to the elastic-optical effect, and similarly Λ is also affected by thermal expansion and contraction. Thus, fiber Bragg grating is widely used for measuring physical parameters such as temperature, stress, and strain. However, fiber Bragg gratings are not sensitive to changes in the surface refractive index because the light is more confined to the core of the fiber and only a small amount of radiation is transmitted to the fiber cladding [13,83,84].

To response the change in surface refractive index, one solution is to etch and grind the cladding of the fiber Bragg gratings [45,85] in which case the evanescent wave extends to the outside of the fiber. For the thinned fiber Bragg gratings, their refractive index sensitivity can be significantly increased. The sensitivity of the thinned fiber Bragg gratings is highly dependent on the fiber diameter, the smaller diameter of the fiber, the higher the sensitivity to refractive index but with this comes a reduction in mechanical strength.

In an LPFG, the grating period is of the order of 100 µm to 1 mm which leads to a coupling between the fundamental core mode and the forward propagating cladding mode. In this case, several attenuation bands centred at discrete wavelengths appear in the transmission spectrum, each corresponding to the coupling with a specific cladding mode. The resonance wavelength $\lambda_{res(m)}^{LPFG}$ is given by the following relationship:

$$ \lambda_{res(m)}^{LPFG} = (n_{core}^{eff} - n_{clad(m)}^{eff}) \Lambda \quad (2) $$
where $n_{clad(m)}^{eff}$ is the effective refractive index of the mth cladding mode, which depends on the surface refractive index of the fiber cladding. Thus, the change of resonance wavelength depends on the refractive index of the surrounding medium. As the surface refractive index increases, a shift in the resonance wavelength occurs.

Another surface refractive index sensitive optical fiber gratings is the TFBG. TFBG likes the regular fiber Bragg grating, also has periodic refractive index modulation along the fiber axis. However, the second mode coupling is between forward-propagating core modes and backward-propagating cladding modes. Therefore, the transmission spectrum of TFBG has many resonance peaks. The first is the resonance peaks corresponding to core mode coupling. The other is a series of discrete resonance peaks corresponding to core-cladding mode coupling. Unlike the fiber Bragg grating, there is a certain tilt angle between the grating plane and the fiber cross-section in the TFBG, leading to more complex mode coupling. Due to the presence of tilt angle, there are two coupling modes in TFBG. The first mode coupling is between forward-propagating and backward-propagating core modes. also present in the short wavelength range. For TFBG with small inclination angles, ghost modes generated by the mode coupling between the core and some low cladding modes may also appear. Among them, the resonance of the cladding mode can be used to detect organic substances. Because when the environment in which the TFBG is located changes within the evanescent field sampling region, the resonance positions of the corresponding cladding modes change accordingly [86]. The light remaining inside the core is insensitive to changes outside the fiber except for temperature and strain, and thus can be used as a thermometer to calibrate temperature and mitigate temperature cross-sensitivity [87]. For TFBG, to consider the resonant wavelength of each mth cladding mode, the Bragg condition given in Eq. (1) is modified as follow:

$$ \lambda_{res(m)}^{TFBG} = (n_{core}^{eff} + n_{clad(m)}^{eff}) \frac{\Lambda}{\cos\theta} \quad (3) $$

where $n_{clad(m)}^{eff}$ is the effective refractive index of the mth cladding mode

and $\theta$ is the tilt angle of the grating. Like LPFG, the resonance wavelength varies depending on the refractive index of the surrounding medium. Conversely, the resonance wavelength shifts to longer wavelengths, i.e., red shift, as the surface refractive index increases. The three types of fiber grating structures are shown in Fig. 1(a).

## 2.2. Optical fiber gratings with surface plasmon resonance effect

Wood [88] first observed surface plasmon in 1902 in the form of anomalous reflection patterns originating from metallized diffraction gratings. It was not until 1968 that Otto explained this phenomenon and used the Otto configuration to achieve the excitation of plasmon at the surface of metal interfaces [89]. In the same year, Kretschmann and Raether [90] used an optical device to demonstrate the excitation of surface plasma excitons. The functional layers of surface plasmon resonance sensors, which consists a optical-fiber with high refractive index, a thin metal layer and a dielectric layer (sensing medium) with a lower refractive index, are categorized as shown in Fig. 1(b). For the optical fiber gratings with surface plasmon resonance effect, firstly, the evanescent wave was generated by a transverse magnetic polarized light incident at the fiber-metal layer interface under the attenuated total reflection condition. Secondly, the produced evanescent wave penetrates through a nanoscale-thick metal layer and excites surface plasmon polariton at the metal-dielectric interface when the phase matching condition is fulfilled [90,91]. However, the phase matching condition will change with the variation in the refractive index of the sensing medium, resulting in a change in the spectral response of the structure [92]. The change of the phase matching condition in the optical fiber gratings can be observed in the coupling angle, coupling wavelength, phase, and intensity of the optical wave.

## 2.3. Optical fiber gratings with located surface plasmon resonance effect

In addition to covering the surface of the optical fiber with a thin film of metal (such as 40–50 nm thick Au film and 30–50 nm Ag film), it is also possible to modify nanostructured metals, such as nanoparticles and

<figure><img src="image_1.png" /><figcaption>Fig. 1. (a) Three different optical fiber gratings (eFBG, LPFG, TFBG) for refractive index sensing and their respective coupling mechanisms. (b) Schematic diagram of the fiber optic coupling technique used to excite the surface plasma. (c) Schematic illustration of located surface plasmon resonance excitation for gold nanosphere. Note: bio-layer (Bio-L), metal layer (ML), buffer layer (BL), supporting tube (ST), waveguide layer (WL), prism (P), metal (M), dielectric (D), incident light (IL) and superstrate (Ss).</figcaption></figure>
nanorods, on their thin surfaces. When the incident light interacts with the metal nanomaterials, the electromagnetic field of the light induces collective coherent oscillations of the surface-conducting electrons of the metal nanomaterials and resonates with the frequency of the light, a phenomenon known as located surface plasmon resonance [93]. Fig. 1 (c) illustrates the excitation of located surface plasmon resonance for a gold nanosphere. Located surface plasmon resonance is sensitive to changes in the localised medium environment, such as refractive index changes caused by biomolecular interactions on the surface of nano-structures, which is the basis of refractive index sensors. Changes in the local environment can be sensed by measuring the located surface plasmon resonance wavelength drift [94,95]. Therefore, this can be used as bio- and chemical-sensors.

### 3. Optical fiber gratings-based sensors for detection of organic pollutants

Currently, the combination of recognition elements and optical fiber gratings with and without surface plasmon resonance has become a common configuration for various bio- and chemical-sensors. Especially with the development of materials science and biotechnology, significant progress has been made in new materials and coating technologies. The advanced materials and coating methods provide important support for the development of bio- and chemical- optical fiber gratings-based sensors to accurately detect organic pollutants. However, the optical fiber gratings with and without surface plasmon resonance as well as nanomaterials have been cross coupled. Therefore, to clearly describe the research progress in optical fiber gratings-based sensors for detection of organic pollutants, instead of classifying them based on the way the nanomaterials and optical fiber gratings, they will be classified according to organics here.

#### 3.1. Detection of volatile organic compounds

Volatile organic compounds are a class of chemicals with high vapour pressure at room temperature and atmospheric pressure [96], which can indicate environmental changes that may affect human health. Therefore, manufacturing low-cost, portable, accurate and real-time volatile organic compounds sensors is of considerable significance in the field of biomedical [97], agricultural and food production [98,99], and environmental monitoring [100,101].

Metal organic framework materials are three-dimensional hybrid crystalline nanomaterials formed by the self-assembly of metal cations

and organic ligands. meta organic frameworks with unique properties, such as high specific surface area, tunable pore and inner cavity shapes and sizes, make them attractive for volatile organic compounds optical fiber gratings-based sensors applications (detection of methanol, ethanol, acetone, butane, etc.) [102]. Wu et al. [46] homogeneously deposited 300-nm-thick zeolitic imidazolate framework on the LPFG by in situ crystallisation to detect ethanol vapour (Fig. 2(a)). When the LPFG coated with zeolitic imidazolate framework was exposed to ethanol at a concentration of 9.8 ppm, the resonant wavelength of the sensor shifted from 1535.08 nm (0 ppm) to 1535.04 nm; The sensor response exhibits high linearity with the ethanol of concentration in the range of 9.8–540 ppm, and the sensor is more sensitive to high concentrations of ethanol vapour. Sensitivity can be further improved by adjusting the grating period so that the sensor operates precisely at the phase-matching turning point. Hromadka et al. [103] proved that when the zeolitic imidazolate framework was deposited on the surface of LPFG and the sensor was working at phase-matching turning point, the sensitivity respectively reached to both ethanol and acetone vapours of 18 ± 1.5 pm/ppm and 15 ± 1 pm/ppm. The low LOD of the sensor for ethanol and acetone reached to 5.56 ppm and 6.67 ppm, respectively (Table 1).

TiO₂ with high refractive index (The refractive index of rutile type and anatase type are 2.76 and 2.55, respectively) has been widely used as the coating materials of the optical fiber gratings-based sensors. For example, Shadab et al. [45] showed that when the eFBG was coated with 30 ± 2 nm-thick reduced graphene oxide/TiO₂ coating, the sensitivity of the reduced graphene oxide/TiO₂ coated eFBG was superior to that of uncoated and TiO₂-coated eFBG sensors. Because the thin film of TiO₂ can improve the effectiveness and integrity of the eFBG and the reduced graphene oxide particles can further enhance core mode perturbation.

Furthermore, the nanoscale atactic polystyrene with high refractive index was also coupled with optical fiber gratings for detection volatile organic compounds. Esposito et al. [104] showed that when the LPFG was coated with atactic polystyrene coating, it can be used to selectively monitor butane gas. Because the wavelength shift of LPFG due to polymer layer changes caused by the adsorption of gas molecules to the atactic polystyrene coating when the sensor is exposed to butane. The sensitivity, LOD, and response time of the sensor are 2.2 nm/vol%, 0.18 vol%, and 180 s, respectively.

#### 3.2. Detection of semi-volatile organic pollutants

Compared with volatile organic compounds, semivolatile organic

<figure><img src="image_2.png" /><figcaption>**Fig. 2.** (a) (I) Optical power distribution of the basic mode (LP01, core mode) and higher-order mode (LP08, cladding mode). (II) Schematic diagram of the sensor [46]. (b) Schematic diagram of the Ta₂C-M/Au/TFBG sensor design. (c) Diagram of the optical detection system [42]. Note: zeolitic imidazolate framework (ZIF-8), broadband source (BBS), polarization controller (PC) and optical spectrum analyzer (OSA).</figcaption></figure>
<table><caption>**Table 1**
Optical fiber gratings-based sensors for detection of organic pollutants.</caption><thead><tr><th>Analyte</th><th>Sensor</th><th>Sensitivity</th><th>Detection range</th><th>RT (s)</th><th>LOD (ppm)</th><th>Ref.</th></tr></thead><tbody><tr><td>Ethanol</td><td>LPFG</td><td>18 ± 1.5 pm/ppm</td><td>61–666 ppm</td><td>NG</td><td>5.56</td><td>[103]</td></tr><tr><td>Acetone</td><td>LPFG</td><td>15 ± 1 pm/ppm</td><td>49–543 ppm</td><td>NG</td><td>6.67</td><td>[103]</td></tr><tr><td>Butane</td><td>LPFG</td><td>0.22 pm/ppm</td><td>1800–10,000 ppm</td><td>180</td><td>1800</td><td>[104]</td></tr><tr><td>Phenol</td><td>LPFG</td><td>0.0492 nm·(mg·L<sup>-1</sup>)<sup>-1</sup></td><td>0.0075–100 mg/L</td><td>395</td><td>7.5 × 10<sup>-3</sup></td><td>[107]</td></tr><tr><td>DNB</td><td>FBG</td><td>10.46 ± 0.06 pm/ppb</td><td>NG</td><td>60</td><td>0.0358</td><td>[109]</td></tr><tr><td>Glyphosate</td><td>LPFG</td><td>3.5 nm/µM</td><td>0–4 µM</td><td>NG</td><td>3.38 × 10<sup>-3</sup></td><td>[30]</td></tr><tr><td>BHT</td><td>TFBG</td><td>232.75 nm/RIU</td><td>10<sup>4</sup>–10<sup>9</sup> pM</td><td>NG</td><td>2.2 × 10<sup>6</sup></td><td>[42]</td></tr><tr><td>Chloroform</td><td>eFBG</td><td>1.5 pm/ppm</td><td>10–100 ppm</td><td>7</td><td>9.22</td><td>[110]</td></tr><tr><td>EEs</td><td>TFBG</td><td>0.433 dB/(ng/mL)</td><td>0.01–100 ng/mL</td><td>~375</td><td>1.5 × 10<sup>6</sup></td><td>[22]</td></tr><tr><td>BPA</td><td>LPFG</td><td>1.5 nm/(g/mL)</td><td>0.01–10<sup>5</sup> ng/mL</td><td>1440</td><td>2.32 × 10<sup>-3</sup></td><td>[113]</td></tr></tbody></table>

Note: response time (RT), low limit of detection (LOD), not given (NG), 1,3-dinitrobenzene (DNB), fiber Bragg gratings (FBG), 2,6-di-tert-butyl-4-methyl phenol antioxidant (BHT), environmental estrogens (EEs) and bisphenol A (BPA).

pollutants are more difficult to degrade and exist for a longer period of time, and are widely present in air, water, soil and other environments. Furthermore, semivolatile organic pollutants are not easy to be degraded, posing a serious threat to ecological and environmental safety and human health [105,106]. Therefore, it is crucial to detect and accurately quantify the concentration of semivolatile organic pollutants in the environment.

The TiO₂ composite coupled with optical fiber gratings can also be used for the detection of semivolatile organic pollutants. Zhong et al. [107] proposed a novel photochemical phenol sensor integrated with a photocatalytic LPFG. The sensor was composed by an etched LPFG, a UV-visible light-driven Er³⁺: YAlO₃/SiO₂/TiO₂ coating, and a polymer film (PEBA2533 doped with β-cyclodextrin) with high phenol permeation selectivity. The Er³⁺: YAlO₃/SiO₂/TiO₂ photocatalyst was used for the photocatalytic degradation of phenol. The sensor achieves selective detection of phenol by shifting the resonance wavelength due to the change in the effective refractive index of the solution caused by the photodegradation by-products of phenol. The results showed that when the Er³⁺: YAlO₃/SiO₂/TiO₂ coating thickness was 500 nm, the polymer film thickness was 25 µm, and the Er³⁺: YAlO₃/SiO₂/TiO₂ coating was irradiated with UV-visible light in the wavelength range of 260–800 nm, the sensitivity, response time, low LOD, and detection range was 0.0492 nm·(mg·L⁻¹)⁻¹, 395 s, 7.5 µg/L, and 7.5 µg/L–100 mg/L, respectively (Table 1). Also detecting phenol compounds, Mishra et al. [108] coated LPFG with tyrosinase-immobilized polyacrylamide gel, which produces complex products when phenol compounds come into contact with tyrosinase molecules. The produced products alter the refractive index of the gel on the surface of LPFG, which in turn affects the resonance wavelength shift of LPFG [108]. The results showed that the sensor has the maximum sensitivity to catechol in the concentration range of 0–1000 µM, 0.0088 nm/µM, with the low LOD of 7.6 µM, and has the advantage of being fast, sensitive, and reproducible.

Eisner et al. [109] described an evanescent field eFBG with doped and non-doped nanoporous silica sol coatings for detection of the airborne nitroaromatic explosive 2,4-dinitrotoluene [109]. The coating doped with 3-aminopropyltrimethoxysilane and a tetrabutylammonium hydroxide can form a charge transfer complex with 2,4-dinitrotoluene, which changes the surrounding refractive index when the airborne 2,4-dinitrotoluene is adsorbed in the silica-sol receptor coating. Experimental results show that the low LOD of 2,4-dinitrotoluene in air is 35.8 ± 0.9 ppb. A lower LOD can certainly be obtained by further etching of the fiber Bragg gratings, but the mechanical properties of the sensor will be weakened. Therefore, other more sensitive sensors such as TFBG, LPFG and surface plasmon resonance can be chosen. Heidemann et al. [30] developed a glyphosate sensor by coating a cysteamine functionalized gold nanoparticle film on a LPFG. The principle is based on the reaction between glyphosate molecules and cysteamine, which alters the effective refractive index of the cladding mode of the LPFG. The low LOD of glyphosate in water is about 20 nM. Furthermore, Yang et al. [42] covered a continuous homogeneous Ta₂C-MXene coating on gold-plated

TFBG for qualitative and quantitative analysis of two carcinogenic food additives saccharin and 2, 6-di-tert-butyl-4-methyl phenol antioxidant (Fig. 2(b) and (c)). The detection limit of the sensor was as low as 10 pM for both carcinogenic food additives.

In 2017, Ahad et al. [110] developed for the first time a novel fiber optic sensor combining polyaniline and FBG for chloroform detection. The polyaniline successfully synthesized by using aniline as monomer, sodium dioctyl sulfosuccinate as dopant, and ammonium persulfate as oxidant. Functionalized FBG can respond to the refractive index change on the surface of the FBG suffered by hydrofluoric acid etching and the charge carriers and -NH⁺ reactive sites carried by the polyaniline chains with the chloroform. The interaction changes the surface refractive index of the FBG surface, leading to the drift of the Bragg wavelength. The sensitivity and detection range was 0.9 pm/ppm and 10–100 ppm. In 2018, Ahad et al. [111] further increased the sensitivity of the sensor to 1.5 pm/ppm with low LOD of 9.22 ppm by adjusting the synthesis parameters of polyaniline. The sensor had the maximum sensitivity when the polymerization temperature was 0 °C and the aniline: sodium dioctyl sulfosuccinate was 5: 5.

### 3.3. Detection of other organic pollutants

Recently, the group of Liu developed a TFBG surface plasmon resonance based biosensor for the detection of environmental estrogens in water [22]. Gold-plated TFBG was first functionalized with desthiobiotin-polyethylene glycol-thiol (DTB-PEG-SH). Then TFBG was put into bovine serum albumin solution to do the blocking treatment, the schematic diagram of the detection system is shown in Fig. 3(a). With the help of molecular dynamics, the 17β-estradiol-streptavidin (E₂-STV) complex was designed and synthesized to act as a competitor for environmental estrogens, releasing more E₂-STV once more environmental estrogens were present in the experimental samples. The E₂-STV could bind to desthiobiotin on the gold-plated TFBG, and the blocking treatment enhanced the specific streptavidin and desthiobiotin interaction. This interaction changed altered the surface refractive index on the gold film surface, leading to a change in the surface plasmon resonance signal, the sensing mechanism of the sensor is shown in Fig. 3(b). The detection limit of the sensor for environmental estrogens can be reduced to 1.5 × 10⁻³ ng mL⁻¹ estradiol equivalent concentration level.

Bisphenol A is an endocrine disrupting compound that can cause serious illness. It is widely used in the production of polycarbonate plastics, epoxy resins and thermal paper [112]. Therefore, accurate monitoring of bisphenol A in the environment is important for the protection of human health. In 2023, Huang et al. [113] proposed an LPFG sensor modified by a DNA aptamer for label-free detection of trace bisphenol A molecules. Firstly, amine groups were functionalized on the surface of the optical fiber. Subsequently, the bisphenol A aptamer was immobilized on the LPFG surface by a dehydration condensation reaction. When the sensor surface comes into contact with a bisphenol A molecule, the conformational transition of the bisphenol A aptamer
<figure><img src="image_3.png" /></figure>

<figure><img src="image_4.png" /><figcaption>Fig. 3. (a) Environmental estrogens detection system. (b) Sensing mechanism of TFBG with surface plasmon resonance for the detection of environmental estrogens in water [22].</figcaption></figure>

leads to a drastic change in the surface refractive index of the fiber, which in turn leads to a change in the transmission spectrum of the LPFG sensor. The sensing platform has high selectivity and excellent bisphenol A detection with low LOD of 0.232 ng/mL (Table 1).

In summary, compared with eFBG and TFBG, LPFG is more suitable for the detection of organic pollutants. Because the fiber Bragg grating is inherently insensitive to the refractive index of surrounding medium, eFBG is faced with low mechanical strength, and TFBG needs to be combined with surface plasmon resonance. The LPFG sensors exhibited extremely high sensitivity to organic pollutants when the period of the LPFG operated at the phase matching turning point and the sensitive materials had high refractive index.

## 4. Optical fiber gratings-based sensors for detection of life organic substances

Chemical optical fiber gratings-based sensors have reasonably good stability and sensitivity, but their selectivity is often limited. Compared

to chemical optical fiber gratings-based sensors, optical fiber gratings-based biosensors tend to be more selective by the specificity of the recognized biomolecules. However, optical fiber gratings-based biosensors are less stable in storage and operation, so the development of optical fiber gratings-based biosensors with good stability, reversibility and reproducibility are still the goal today. This section introduces optical fiber gratings-based biosensors based on biomolecular modifications for the detection of life organics substances. The organics substances include the glucose, microorganisms, proteins, nucleic acids, enzymes, antibodies-antigens, drugs and some other biomolecules such as vitamin D.

### 4.1. Detection of glucose

Glucose sensors play an important role in scientific and biological research, especially in medicine, human glucose levels are related to many diseases, such as diabetes, cancer, etc [114,115]. OTherefore, the rapid and accurate detection of glucose concentration is crucial for the
medical diagnosis and food safety supervision. We know that the refractive index of a glucose solution increases linearly with increasing glucose concentration and the concentration can be measured with a refractive index sensitive sensors. However, the difference in refractive index between glucose concentration levels in humans is small, and any other dissolved substances can also cause changes in refractive index. Therefore, in the field of life sciences, glucose concentration sensors need a low glucose concentration range, high biocompatibility, high stability, and low LOD. To fulfill the above objectives, the glucose sensors, which are based on the coupled with optical fiber gratings and biomaterials such as glucose oxidase [116] and haemoglobin [117] that can bind specifically to glucose molecules, were developed.

In 2012, Deep et al. [118] first used a glucose oxidase modified optical fiber gratings-based sensor to detect glucose concentration, but the use of the toxic material 3-aminopropyl-triethoxysilane to connect the fiber to the glucose oxidase was unfavourable for biocompatibility and environmental friendliness. In 2018, Jiang et al. [21] used a label-free biosensor based on graphene oxide and glucose oxidase functionalized large tilted fiber gratings for low concentration glucose detection. The detection results showed that the resonance wavelength of the sensor had a linear response to glucose concentration in the range of 0–1.5 mg/mL with a response coefficient of about 1.33 nm/(mg/mL). The introduction of graphene oxide and glucose oxidase can greatly improve the sensitivity, and graphene oxide instead of 3-aminopropyl-triethoxysilane can solve the safety hazards caused by 3-aminopropyl-triethoxysilane toxicity. Therefore, graphene oxide/glucose oxidase functionalized optical fiber gratings for detection of glucose has been favoured by researchers. However, the production process of TFBG is complicated and costly, compared with it, the graphene oxide/glucose oxidase modified LPFG sensor has a greater advantage in industrial production because of the low production cost and simple process of LPFG. In 2019, Xu et al. [64] developed graphene oxide/glucose oxidase functionalized LPFG and successfully constructed a label-free glucose biosensor. The biosensor maintained high sensitivity and showed a linear response in the concentration range of 0–1.2 mg/mL with a response coefficient of 0.77 nm/(mg/mL).

However, another researcher indicated that the glucose enzymes for may generate irreversible inactivation when their exposed to high temperatures and organic solvents [90]. Therefore, Liang et al. [119] proposed to encapsulate the enzyme with meta organic frameworks, which increase the enzyme's thermal stability and tolerance to organic solvents. Furthermore, Zhu et al. [116] also presented zeolitic imidazole-based metal organic frameworks/glucose oxidase composite functionalized LPFG for glucose detection, which can achieve high protein encapsulation efficiency and exhibit high activity and stability under harsh conditions. The sensor exhibited a linear response in glucose solutions of 0.19 mg/mL with a sensitivity of about 2.77 nm/(mg/mL) (Table 2).

Despite the excellent sensitivity and selectivity of the enzyme-based optical fiber gratings-based sensors, the inherent fragility of the enzyme leads to poor accuracy and reproducibility and a lack of immunity to temperature and pH changes. Therefore, surface plasmon resonance

TFBG sensor was developed by Lobry et al. [28]. TFBG was first plated with gold and then covered with polydopamine to enhance the bonding between the gold layer and concanavalin A before the functionalisation of concanavalin A. The surface plasmon resonance TFBG glucose sensor shows high sensitivity in the range of 1 µM to 0.1 mM with the low LOD of 0.01 µM. Dyankov et al. [117] controlled the LPFG to work near the phase-matching turning point and for the first time used the refractive index change caused by the selective binding of glucose to hemoglobin for glucose detection. The haemoglobin was deposited directly on the surface of optical fiber gratings-based on the matrix-assisted pulsed-laser evaporation (MAPLE) method, instead of chemical methods to immobilize the specific recognition molecule. The experimental results demonstrate that the MAPLE technique preserves the biological activity of haemoglobin, and the low LOD of the glucose LPFG biosensor reached to 0.1 mmol/L.

## 4.2. Detection of hemoglobin

Hemoglobin is an important substance within the blood of animals and is a key chemical related to gas transportation in the body [120], so it is crucial for the detection of hemoglobin content in the blood.

The earliest use of optical fiber gratings-based sensors to detect hemoglobin concentration dates back to 2007. The group of Chen corroded the LPFG with hydrofluoric acid to make it work close to dispersion turning point, and was able to achieve a 0.005 % change in protein concentration without modifying any of the biorecognition molecules [121]. Chen et al. [122] using a functionalized LPFG coated with graphene oxide nanosheets. The sensing principle was based on the resonance intensity change caused by the adsorption of graphene oxide on hemoglobin molecules. The functionalized LPFG sensor has an ultra-high sensitivity of 1.9 dB/(mg/mL) for hemoglobin detection, with a detection limit of 0.05 mg/mL (Table 2). Later in 2020, Wang et al. [123] prepared a microtapered LPFG via a CO₂ laser heating source, also with a graphene oxide film deposited on top of the fiber surface. A sensitivity of 2 nm/(mg/mL) was achieved with a minimum detection limit of 0.02 mg/mL at hemoglobin concentrations from 0.0 mg/mL to 2.0 mg/mL in deionized water. Furthermore, Sun et al. [124] investigated an excessively TFBG-based hemoglobin biosensor coated with graphene oxide nanosheets for the detection of hemoglobin biomolecules. The experimental results show that the sensitivity and detection range of the sensor were adjustable by adjusting the amount of graphene oxide coating. The sensitivity of the sensor increased with the increase of graphene oxide doping content, but the range of sensor detection hemoglobin decreased with the increase of graphene oxide doping content.

Due to the excellent light absorption, high surface-to-volume ratio and molecular adsorption ability of black phosphorus [125]. Zhang et al. [126] attempted to deposit black phosphorus films on LPFG for bovine hemoglobin detection. Compared with bare LPFG, LPFG deposited with black phosphorus significantly improved the detection of bovine hemoglobin with a sensitivity of 5.6 dB/(mg/mL) and a detection limit of 0.05 mg/mL.

<table><caption>**Table 2**
Optical fiber gratings-based sensors for detection of life organics.</caption><thead><tr><th>Analyte</th><th>Sensor</th><th>Sensitivity</th><th>Detection range</th><th>RT (s)</th><th>LOD</th><th>Ref.</th></tr></thead><tbody><tr><td>Glucose</td><td>LPFG</td><td>2.78 nm/(ng/mL)</td><td>0.18–1.4 ng/mL</td><td>NG</td><td>0.18 ng/mL</td><td>[116]</td></tr><tr><td>Hemoglobin</td><td>LPFG</td><td>1.90 × 10<sup>-6</sup> dB/(ng/mL)</td><td>5 × 10<sup>-4</sup>–10<sup>-6</sup> ng/mL</td><td>NG</td><td>5 × 10<sup>-4</sup> ng/mL</td><td>[61]</td></tr><tr><td>Cancer cell</td><td>TFBG</td><td>0.005 dB/(cells/mL)</td><td>10–10<sup>-6</sup> cells/mL</td><td>300</td><td>10 cells/mL</td><td>[128]</td></tr><tr><td>CK17</td><td>TFBG</td><td>15 dB/(ng/mL)</td><td>10<sup>-3</sup>–0.1 ng/mL</td><td>580</td><td>3.75 × 10<sup>-5</sup> ng/mL</td><td>[129]</td></tr><tr><td>IgG</td><td>LPFG</td><td>1.33 × 10<sup>-3</sup> nm/(ng/mL)</td><td>3 × 10<sup>-3</sup>–0.02 ng/mL</td><td>2400</td><td>23 ng/mL</td><td>[134]</td></tr><tr><td>IgM</td><td>LPFG</td><td>11 nm (ng/mm<sup>2</sup>)<sup>-1</sup></td><td>0.0156–1 ng/mL</td><td>600</td><td>1.5 × 10<sup>-8</sup> ng/mL</td><td>[61]</td></tr><tr><td>Exon 20</td><td>FBG</td><td>0.04 nm/nM</td><td>13.5–100 nM</td><td>2400</td><td>13.5 nM</td><td>[142]</td></tr><tr><td>Thrombin</td><td>TFBG</td><td>0.19 dB/(ng/mL)</td><td>0.182–6.142 ng/mL</td><td>5400</td><td>0.182 ng/mL</td><td>[58]</td></tr><tr><td>Pepsin</td><td>LPFG</td><td>0.26 nm/(ng/mL)</td><td>1–1000 ng/mL</td><td>300</td><td>25.79 ng/mL</td><td>[144]</td></tr></tbody></table>

Note: cytokeratin 17 (CK17), response time (RT), low limit of detection (LOD), not given (NG), immunoglobulin G (IgG) and immunoglobulin M (IgM).
In addition, there are also researchers who have made changes in the
sensor structure. Yang et al. [127] cascaded the LPFG with a single-mode
coreless single-mode (SNCS) fiber. By analyzing the opposite offset of
the interference depression formed by the LPFG and SNCS, the surface
refractive index change can be measured and temperature compensation
is achieved. The experimental results showed that the sensor had a
temperature compensation coefficient of -0.727 mg/mL and a detection
sensitivity of 0.0165 nm/(mg/mL) for porcine hemoglobin.

4.3. Detection of cancer markers

While cancer markers are not sufficient to be used alone to confirm a
cancer diagnosis, they can also be used as an adjunct tool to help phy-
sicians with risk assessment, monitoring of treatment efficacy,
screening, and early detection of patients. HER2 protein and
mammaglobin-A are relevant biomarkers used in the diagnosis of breast
cancer. Cytokeratin 17 is associated with lung cancer. Therefore, rapid
and accurate detection of cancer marker components and levels is
essential for human health management and disease treatment.

Applying optical fiber gratings to the diagnosis of breast and lung
cancer, the Lobry team has done a lot of work [31,59,60,128-130]. In
2020, Lobry et al. [128] first deposited a 50-nm-gold film on TFBG for
stimulating the surface plasmon resonance effect. Subsequently,
aptamers (MAMA2) was selected to biofunctionalize the gold-plated
TFBG for specific detection of mammaglobin-A. The schematic dia-
gram of the sensor is shown in Fig. 4. The experimental assay showed
low LOD results up to 100 cells/mL [128]. After that, to improve the
sensor performance and robustness, Lobry et al. [59] have made the
following three improvements. (1) The single polarization fiber was
selected because of single polarization fiber with a very convenient P-
polarised state propagation to enhance the surface plasmon resonance
signal. (2) The user-friendly demodulation technique of the tracking of
the surface plasmon resonance feature envelope fit was adopted. (3)
Aptamer-protein-antibody sandwich bioassays was developed. Experi-
mental results have shown that the sensor produces a wavelength drift of
800 nm, which is 2.5 times higher than under conditions of aptamer-
HER2 (Human Epidermal Growth Factor Receptor-2) interactions [59].

Zhu et al. [60] also noted that the optical biosensors bring an
evanescent wave in the surrounding medium, which makes them sen-
sitive to both surface refractive index variations (useful biosensing
signal) and to bulk refractive index changes (non-useful signal for bio-
sensing). To discriminate both volume and surface refractive index
changes and increase the signal-to-noise ratio, one part of the TFBG
grating was used as volume refractive index sensing and the rest was
gold-plated and biofunctionalized to measure the surface refractive
index changes upon binding to HER2 [59]. To further improve the low

LOD and signal-to-noise ratio of the sensor, in 2023, Lobry et al. [29]
have developed an electro-plasmonic TFBG biosensor. A 40-nm-thick
gold layer was deposited on the surface of TFBG to drive the current
and polarize the TFBG probe. The electro-plasmonic TFBG biosensor
exhibited a good low LOD of 10<sup>-12</sup> g/mL for HER2 biomarker in solution
which shows a 100-fold enhancement than that of the conventional
TFBG-based biosensing method. Furthermore, the electro-plasmonic
TFBG biosensor also showed a high signal-to-noise ratio. These advan-
tages are attributed to the electro-plasmonic TFBG can attract mole-
cules, make cells closer from the sensing surface, and analyse their
binding in real time.

Furthermore, In 2016, Ribaut et al. [129] developed an immuno-
sensor to detect cytokeratin 17 using a 50-nm-thick gold film on TFBG to
excite the surface plasmon resonance effect and modified anti-
cytokeratin 17 antibody used to selectively bind cytokeratin 17. The
developed sensor successfully performed non-invasive in vitro mea-
surements on human tissue after clinical lobectomy and provided a
positive biosensor response in tumour samples with a detection limit of
1 pM [129]. In 2019, Loyez et al. [130] also designed another minimally
invasive catheter based on TFBG sensor embedded biosensor, which
enables fine positioning of the probe in the tissue and was validated in
the porcine lung [130].

To obtain a high-performance optical fiber gratings-based biosensors with surface plasmon resonance effect for detection of protein, the gold coating on optical fiber gratings is usually done using evaporators or sputter coaters under a space under vacuum conditions. Therefore, the thickness of the metal deposited on the fiber cannot be determined directly and was usually measured by using the external quartz micro-balance or scanning electron microscope in off-line. To obtain the measurement of metal layer on the fiber in on-line, the electroless plating technique has gained the attention of researchers because it allows online monitoring of the fiber signal. In 2018, the group of Lobry used the electroless plating technique to excite the located surface plasmon resonance effect by depositing Au nanoparticles on it and modified with anti-cytokeratin 17 antibodies [31]. The sensor reached a detection limit of 14 pM for cytokeratin 17, which is slightly lower than using an evaporator or sputter coater [31].

In 2019, Loyez et al. [131] combined the electroless plating gold plating technique with sputtering gold plating. Firstly, a gold coating of about 4 nm was deposited on both sides of the TFBG by sputtering, and then the fiber were placed directly into the gold plating solution. This allowed the fiber to elicit the electroless plating reaction without the need for silanization and the addition of gold nanoparticles. They compared four different configurations of TFBGs. (1) bare TFBG, (2) gold plating of TFBG using a sputterer, (3) electroless plating method for gold plating of TFBGs and (4) a hybrid coating combining sputtering and

<figure><img src="image_5.png" /><figcaption>**Fig. 4.** (a) Schematic representation of target cancer cells interacting with aptamers and functionalised gold nanoparticles immobilised on the surface of an optical fiber. (b) Schematic representation of a specific HER2 bioassay based on the use of anti-HER2 aptamer biofunctionalised with surface plasmon resonance TFBG [128].</figcaption></figure>
electroless plating gold plating, for cytokeratin 17 detection. Of these, the bare TFBG shows the lowest sensitivity, and the TFBG combining non-chemical electroless plating gold plating technology and sputtered gold plating technology exhibits sensitivity comparable to that of a fully sputter gold plating, while greatly reducing the difficulty of production.

## 4.4. Detection of antibody-antigen

Antibodies-antigens are involved in humoral immune responses and their levels increase during the immune system's response to pathogenic infections. The concentration levels of antibodies/antigens reflect disease indicators in the human body, so the development of low cost, stable and reusable antibody-antigen sensors is important for the medical diagnostics [132].

The principle of the detection of antibodies or antigens is basically to use their interaction, so the antibody or antigen becomes a natural biological material used to specifically recognize each other. This interaction leads to the formation of complexes on the surface of optical fiber gratings, which results in a change in the resonance wavelength. However, the problems of poor antigen-antibody binding and small changes in refractive index have affected the sensitivity and low LOD of the sensors. Researchers have improved these problems by developing the functional materials such as graphene oxide, Au nanoparticles, Ag nanoparticles, SiO₂ nanoparticles, TiO₂ polystyrene, etc. and achieved good results.

In 2017, Liu et al. [133] proposed a kind of dual-peak LPFG ultra-sensitive antibody-antigen immunosensor functionalized by graphite oxide nanosheets. The experimental results showed that the low LOD of the sensor was 7 ng/mL, which was 10 times better than that of the uncoated graphene oxidedual-peak LPFG. This is because graphene oxide improves the chemical binding and physical adsorption of bio-recognition molecules, as shown in Fig. 5(a). In addition, Dong et al. [134] used the same method described above to deposit and

functionalize graphene oxide films on a U-bent LPFG that is embedded in a dual-mode fiber. The experimental results showed a high sensitivity of 1.32777 nm/(µg/mL) for IgG and a minimum detection limit of 23 ng/mL (Table 2). In 2023, Dey et al. [135] proposed three kinds of measures to enhance the sensitivity. The first was to make the sensor work in the near mode transition of the cladding mode. The second was to make the sensor work near the phase matching turning point. The third was to enhance the evanescent field of the cladding mode by decreasing the diameter of the cladding. Subsequently, the electrostatic self-assembly technique was used to deposit a nanoscale polymer overlay consisting of poly(allylamine hydrochloride) and sodium poly-4-styrenesulfonate on the sensor surface. The low LOD of the sensor reached to 6.9 × 10⁻⁸ µg/mL for detection of IgG. Considering the cost of LPFG fabrication, Chen et al. [136] developed an IgG sensor based on the resonance of ultra-narrowband cladding modes in fiber Bragg gratings. The coupling between the forward-propagating guided core mode of the multimode fiber and the backward-propagating guided cladding mode of the fiber Bragg gratings, without eroding the cladding diameter, enables the FBG to respond to the changes in the surrounding refractive index and enhances the mechanical properties of the FBG. The experimental results showed that the low LOD of the sensor was 32 pM. Since fiber-based sensors are susceptible to temperature crosstalk, Xiao et al [17] proposed a highly birefringent microfiber Bragg grating biosensor with temperature self-compensation. The IgG molecules were immobilized the surface of the optical fiber to achieve specific detection of anti-IgG molecules. Although the sensitivity and low LOD of the biosensor were several tens of nm/RIU and µg/mL, respectively, this work provided a new idea to resolve the problem of temperature cross-sensitivity of fiber optic biosensors.

Carbonic anhydrase IX is a diagnostic protein biomarker for tumour cell hypoxia. Tyagi et al. [3] modified LPFG with Au nanoparticles and immobilized anti-carbonic anhydrase IX antibodies using 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide and N-hydroxysuccinimide cross-

<figure><img src="image_6.png" /><figcaption>**Fig. 5.** (a) Dual-peak LPFG ultrasensitive antibody-antigen immunosensor [133]. (b) The schematic diagram of the layer-by-layer modification of LPFG and chemical decoration of gold nanoparticles with amine groups was used to attach biotin [61]. (c) (I) The experimental setup of the CM-LPFG sensor for protein detection experiments. (II) The schematic diagram of the CM-LPFG sensor detailing the size of the grating structure [54]. Note: long period grating (LPG), phosphate buffer saline (PBS), immunoglobulin G (IgG), graphene oxide (GO), polycations (PAH), 3-aminopropyltriethoxysilane (APTES), syringe infusion pump (KDS), comb of microchannel long-period fiber grating (CM-LPFG), optical spectrum analyzer (OSA).</figcaption></figure>
linking method. The sensor was experimented in flow cells containing the A549 human lung adenocarcinoma cell line, and 6.25 × 10⁴ cells/mL were detected in cell culture medium.

Human immunoglobulin M levels are clinically important for the diagnosis of diseases. Liu et al. [61] measured immunoglobulin M by coating LPFG working under phase-matching turning point with a layer-by-layer method containing three layers of poly (allylamine hydrochloride)/gold-plated silica nanoparticles, as shown in Fig. 5(b). The silica nanoparticles have a high surface-to-volume ratio and the gold nanoparticles in turn can excite an located surface plasmon resonance effect leading to improve the low LOD (15 pg/mm²).

Most of the microfluidic systems used in general biosensors are straight channels. However, Wen et al. [54] fabricated a comb-shaped microfluidic channel for detecting SK-Hep1, as shown in Fig. 5(c). Utilize specific reactions between anti-histone deacetylase 6 (Anti HDAC6) and SK Hep1. When anti-HDAC6 binds to HDAC6, the surface refractive index will be changed, resulting in a shift in the resonant wavelength of the sensor. The low LOD of the biosensor reached to 1.363 ng/µL.

In terms of TFBG-based plasmonic antibody-antigen sensing applications, in 2018, Wang et al. [137] used graphene oxide and staphylococcal protein A for the detection of IgG. The introduction of graphene oxide and staphylococcal protein A can immobilize more antibodies onto the surface of the gold film, enhance the binding efficiency of antigens and antibodies, and further improve the sensitivity of the sensor. Experimental results showed that the sensitivity and low LOD of the biosensor were 0.096 dB/(µg/mL) and 0.5 µg/mL, respectively.

Furthermore, optimizing the intrinsic structure of optical fiber gratings was also used to improve the performance of sensors. Xiao et al. [138] fabricated a high-order diffraction (HOD)-LPFG with a period number less than 10 and a long grating pitch to detect prostate specific antigen concentration in phosphate buffer saline. Dey et al. [139] demonstrated an etched LPFG biosensor (the cladding was etched to a

diameter of 21 µm) to keep it working near the turn around point. The etched LPFG was functionalized using physisorption deposition of Eudragit L100 copolymer to provide a carboxy functional group (–COOH) for immobilized immunoglobulin G. The 1 ethyl-3(3-dimethyl amino propyl) carbodiimide HCl and N-hydroxy Succinimide bifunctional cross-linker was used to activate the –COOH group. By measuring the dual peak resonance of lowest order cladding mode (LP₀,₂ cladding mode), the low LOD of the biosensor for the detection of anti-immunoglobulin G reached to 0.16 ng/mL.

### 4.5. Detection of biotin and streptavidin

Biotin, also known as vitamin H, is a water-soluble vitamin belonging to the B-vitamin group. Biotin is involved in numerous metabolic processes in living organisms and is indispensable for the normal metabolism of fats and proteins [140]. Streptavidin (STV) is a protein with a very high affinity for biotin, and the streptavidin/biotin binding is considered as one of the strongest non-covalent interactions (Ks ~ 10¹³ M⁻¹) known in nature. Therefore, biotin and streptavidin then form a biological model system as a natural pair of biorecognition elements and analyte.

In 2016, SiO₂/Au nanoparticles coated LPFG biosensor was created to detect avidin [62]. SiO₂/Au nanoparticles were electrostatically assembled onto optical fiber using a polycation layer of poly(allylamine hydrochloride), as shown in Fig. 6(a). SiO₂ nanoparticles increased the porosity of the sensitive layer, resulting in a larger surface area and enhanced interaction efficiency between evanescent waves and the sensitive layer. Au nanoparticles were used to stimulate the located surface plasmon resonance effect, and because Au nanoparticles and SiO₂ nanoparticles can guarantee the stability of combination through a covalent bond, the sensor stability and reusability were improved. The LPFG sensor was set to operate near the phase-matching turning point.

<figure><img src="image_7.png" /><figcaption>**Fig. 6.** (a) Schematic representation of layer-by-layer deposition of (PAH/SiO₂:Au)₂ films on LPG [62]. (b) A bilayer (PEI/PAA)₁ mFBG sensor surface structure [56]. (c) Schematic diagram of the structure of the sensor. (II) The SEM of sensor surface [142]. Note: long period grating (LPG), fiber Bragg grating (FBG), microfiber Bragg grating (mFBG), poly(acrylic acid) (PAA), polyethylenimine (PEI) and chitosan (CS).</figcaption></figure>
The sensor exhibited a low LOD of 19 pg/mm². However, the sensor performance was not optimal due to the poor surface coverage and uniformity of the sensor coating and the sensor was near rather than at the phase-matching turning point. Subsequently, in 2018, the silica core gold shell nanoparticles was modified on the surface of the LPFG biosensor by using layer-by-layer method [61]. The LPFG sensor, which uniformly coated with a film comprising three layers of PAH/Si nano-particles with gold shell, showed a high sensitivity of 3.88 (ng/mm²)⁻¹ with a good low LOD and minimum detectable concentration respectively of 0.86 pg/mm² and 1.25 nM for the detection of streptavidin. The improved performance of the biosensor was due to the LPFG operating close to the phase matching condition (Table 2).

To avoid the Au nanoparticles on the surface of the optical fiber gratings being washout during experiments, Dominik et al. [48] chose to deposit titanium oxide films on LPFG to provide mechanical protection for Au nanoparticles. Moreover, the amine group on the titanium oxide film can bind to the peptide bond formed between the carboxyl groups of biotins, thus increasing the amount of biotin immobilization and thus improving the performance of the sensor.

## 4.6. Detection of DNA

As an important biosensor, DNA sensors play an important role in genetic engineering, drug development, epidemic controls, health applications and clinical medical research, etc [141]. The process by which two complementary DNA sequences are paired by Watson-Crick bases to form a stable double-stranded DNA molecule is called DNA hybridization. The process of hybridization of DNA is highly specific and detection of specific sequences of DNA can be used to detect genetic mutations or biomarkers associated with specific diseases.

Sun et al. [56] designed a microfiber Bragg gratings-based biosensor to detect DNA in microfluidic chips. The polyelectrolyte multilayer membrane was prepared by using the polyethylenimine, poly(acrylic acid), and single-stranded DNA and was loaded on the surface of the sensor by using the layer-by-layer technique. The absorption capacity of the biosensor for target DNA chains was enhanced due to the use of (polyethylenimine/poly(acrylic acid))₄(polyethylenimine/DNA)₁ multi-layer films immobilized on the surface of FBG, as shown in Fig. 6(b). The probe uses complementary base-pair reactions to cause a change in the refractive index of the surface fiber, which leads to a wavelength shift. The sensor successfully detected the target single-stranded DNA hybridization process at a concentration of 1 µM in a microfluidic chip.

In addition, Li et al. [142] demonstrated a FBG surface plasmon resonance based DNA hybridization sensor. The sensor can achieve dual surface plasmon resonance in an FBG and implement a three-parameter detection of DNA sequence, temperature and pH using a single fiber, as shown in Fig. 6(c). Particularly, the sensor showed a good low LOD of 13.5 nM and high sensitivity of 0.04 nm/nM for the detection of exon 20 of the epidermal growth factor receptor gene (Table 2).

## 4.7. Detection of enzymes

Enzymes are important substances involved in the life activities of living organisms and help regulate many other physiological processes. Furthermore, the changes in the type and content of enzymes in living organisms often indicate abnormal health conditions. Therefore, the detection of changes in the type and content of enzymes is important for the quick and accurate determination of the dosing strategy and reduce the severity of the disease.

In 2016, Quero et al. [143] developed a reflective LPFG for the rapid detection of AmpC β-lactamases. For the reflective LPFG biosensor, first, the nanoscale atactic polystyrene with a high-RI overlay was coated on the reflective LPFG to improve its sensitivity to surface refractive index. Second, the poly(methylmethacrylate)-co-methacrylic acid (PMMA-co-MA) was coated on the surface of the atactic polystyrene coated reflective LPFG to promote a stable covalent bioreceptors

immobilization. Third, 3-aminophenylboronic acid (3-APBA) was fixed on the surface of the PMMA-co-MA/atactic polystyrene-coated reflective LPFG. The 3-APBA as a biocognitive element specifically recognizes β-lactamase, resulting in a change in the local refractive index, which in turn affects the change in resonance wavelength of the biosensor. The biosensor’s low LOD can reach tens of nM for detection of β-lactamases.

In 2018, Lao et al. [58] proposed a TFBG-based sensor that combines surface plasmon resonance and located surface plasmon resonance for in situ detection of thrombin, as shown in Fig. 7. The 50 nm gold film is first sputtered onto the TFBG, then the sulfhydrylated 15-mer DNA aptamer (Apt15) and 29-mer DNA aptamer (Apt29) are attached to the gold layer by covalent bonding, and finally, the 13 nm Au nanoparticles are bound to the outermost layer. The plasmon resonance coupling between the gold film and Au nanoparticles was used to improve the sensor limit for thrombin detection, and the low LOD of this thrombin sensor reached to 1 nM.

Strong located surface plasmon resonance effect on the surface of optical fiber gratings requires precise control of the diameter and shape of the nanoparticle as well as the thickness and uniformity of the metal deposition. This greatly increases the difficulty as well as the cost of sensor fabrication. In reality, the sensitivity of refractive index change can be enhanced without covering the optical fiber gratings with functional materials. An aptasensor based on etched TFBG was developed by Marzhan et al in 2019 [23] on a single-mode optical fiber for thrombin detection. The sensor surface was not deposited with any functional material, but only the fiber diameter was reduced by etching the TFBG with hydrofluoric acid to adjust the sensitivity, and the surface was functionalized with thrombin aptamer. The results highlighted that the aptasensor showed excellent LOD of 0.075–0.110 nM. However, the mechanical strength of the sensor is compromised by this method due to the reduction of the fiber diameter.

Pepsin test is an important clinical tool to determine the normal function of the stomach and to diagnose gastric diseases. To rapidly and accurately detect pepsin, in 2023, a label-free and sensitive micro-tapered LPFG sensor that functionalized by graphene oxide was proposed for pepsin detection [144]. The amide group formed between the carboxyl group of graphene oxide and the amino group in pepsin allowed pepsin to be firmly adsorbed by graphene oxide, resulting in an alteration of the surface refractive index of the micro-tapered LPFG. The graphene oxide coated micro-tapered LPFG had a high sensitivity to pepsin concentrations in the range of 1–1000 ng/mL with a low LOD of 25.79 ng/mL (Table 2).

Here we found that the biochemical molecular material properties of selectivity, stability, and reproducibility used in the detection of life organics are critical to the bio- and chemical-optical fiber gratings-based sensors. Because the concentration levels of glucose, hemoglobin, DNA and enzymes are extremely low in blood plasma or other biological fluids and the accuracy measurement of the life organics is crucial in the evaluation of human health level and analysis of etiology. Furthermore, to our knowledge, there have been no reports on the TFBG for the detection of DNA, biotin and streptavidin, and hemoglobin. In terms of DNA and cancer marker detection, LPFG sensors should also be developed in the future.

# 5. Optical fiber gratings-based sensors for detection of microorganisms and other organic substances

## 5.1. Detection of microorganisms

Microorganisms are found everywhere and can be detrimental or beneficial in nature. For example, the clean water will be contaminated by formation of microorganism [145], and the humans are more vulnerable than ever to water infections, viruses and bacteria. Alternatively, effective bioremediation and high-efficiency bioenergy require high-quality microorganisms [146]. Therefore, rapid and accurate detection of microorganisms is imperative to control their growth and
<figure><img src="image_8.png" /><figcaption>Fig. 7. (a) Schematic of the optical fiber sensing system for detection of thrombin. (b) Microscopic image of a gold coated fiber optic sensor. (c) SEM image of the prepared gold nanoparticles. (d) Schematic of the detection principle of the TFBG's gold nanoparticle functionalised surface plasmon resonance fiber-optic sensor. (e) Schematic details of surface functionalization [58]. Note: polarization controller (PC), optical spectrum analyzer (OSA) and surface plasmon resonance (SPR).</figcaption></figure>

metabolic activity in different fields [147].

To detect *Escherichia coli* (*E. coli*), in 2016, Dandapat et al. [148] developed a temperature-insensitive LPFG biosensor closest to its turn-around wavelength. The biosensor showed that a spectral shift between 1.3 nm and 2.5 nm occurred in the range of *E. coli* concentrations of 10²–10⁵ CFU/mL, indicating that the sensor can reliably and quantitatively detect *E. coli* in water. In the same year, Koba et al. [148] controlled LPFG to work in the vicinity of phase-matching turning point for the label-free detection of *E. coli* *B*. The sensor was functionalised using phage adhesin enabling it to selectively bind to *E. coli*. The *E. coli* *B* caused a resonance wavelength shift over 90 nm.

Recently, Singh et al. [149] developed a 40-nm-thick-TiO₂ nano-coated etched LPFG biosensor with T4 phage immobilised on the surface of the TiO₂ coating for the detection of *E. coli* at 0 CFU/mL–50 CFU/mL. The sensitivity was 2.55 nm/RIU and the low LOD value was 10 CFU/mL. It has been demonstrated that the addition of TiO₂ coating to the device can improve sensitivity because the TiO₂ coating can enhance the change in surface refractive index of the optical fiber gratings. However, conventional high refractive index metal oxides deposited on LPFG are less resistant to alkaline solutions, which are often used for hydroxyl-ation when functionalizing optical fibers. Piestrzynska et al. [25] showed that tantalum oxide (TaOx) is with the same functionality as high refractive index overlays, such as Al₂O₃ and TiO₂ were deposited by atomic layer deposition methods used to induce mode transition effects, making the LPFG highly sensitive to changes in refractive index. The ultrasensitive TaOx-coated LPFGs for their surface biofunctionalization can be used to detect the *E. coli* and avidin protein. In addition, Cele-banska et al. [150] developed bioinspired sensor based on etched LPFGs and α-mannoside receptor, which was the first time that the combination of bionic carbohydrates and LPFG for the detection of *E. coli*. The experimental results showed a low LOD of 10³ CFU/mL of the target sample and the formation of biofilm by *E. coli* on the surface of mannose was observed.

In addition to *E. coli*, other microorganisms can also be detected by optical fiber gratings-based sensors. Zhong et al. [151] created a microstructured fiber Bragg grating to accurately and simultaneously monitor the cell growth of *Rhodopseudomonas palustris* (*R. palustris*) CQK

01 and temperature in a bioreactor. Yang et al. [152] used a layer-by-layer deposition technique, a nanoporous polyelectrolyte coating of [PAH/PAA]₁₀-PAH was shaped on the surface of LPFG, and then the-antibody of *Staphylococcus Aureus* (*S. Aureus*) was immobilized on it. The size of the pit structure in the PAH/PAA coating was comparable to that of *S. Aureus*, promoting more bacterial adhesion to the surface, which enhanced the sensitivity of the sensor. The experiments were performed in the concentration range between 10⁴–10⁸ CFU/mL and showed that the sensitivity of LPFG modified by [PAH/PAA]₁₀-PAH and *S. Aureus* antibody was 0.478 ± 0.005 nm/log (CFU/mL) and the low LOD was 224 CFU/mL. In addition, Gan et al. [153] immobilised specific yolk antibodies (IgY) purified from sterile eggs in the LPFG grating region (Fig. 8). The bacterial-antibody interactions led to an increase in the thickness and density of the biological overlayer, enabling detection of *S. Aureus* by tracking changes in resonance wavelengths in the LPFG transmission spectra. This antibody has the advantages of low cost, ease of preparation, ease of mass production, and the ability to bind to antigens over a wide pH range. The sensor showed selective and sensitive to *S. Aureus*, and the resonance wavelengths change induced by *S. Aureus* between 10² CFU/mL and 10⁹ CFU/mL were between 0.60 nm and 4.61 nm, and the low LOD was appropriately 33 CFU/mL (Table 3).

In addition to bacteria, researchers have also done measurements of phages, viruses, and fungi using optical fiber gratings-based biosensors. A sensor for the detection of *Newcastle disease virus* (NDV) was developed by Luo et al. [154]. By modifying AuNs on large tilted fiber grating (80 nm) was used to stimulate the located surface plasmon resonance effect, thus amplifying the effect of refractive index fluctuations on the core to co-propagation envelope and improving the sensitivity and detection limit of the sensor. Activated staphylococcal protein A was then attached to gold nanospheres for enhancing the bioactivity of anti-NDV monoclonal antibodies this was able to promote specific binding of the antibodies to NDV. The proposed sensor had a low LOD of 25 pg/mL, which was a 5–10 times increase compared to the case without AuNs coating.

Lee et al. [155] proposed a phase-shifted LPFG with the severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) spiking antibody for selective detection of the SARS-CoV-2. When exposed to different
<figure><img src="image_9.png" /><figcaption>**Fig. 8.** LPFG biosensor for the detection of S. aureus using IgG as the biorecognition elements. (a) Schematic diagram of yolk antibody extraction. (b) Function-alisation of the LPFG surface to immobilise yolk antibodies. (c) Diagram of the detection system [153]. Note: optical spectrum analyzer (OSA), single mode patch cord (SMPC), single mode optical fiber (SMF), micro-channel groove (MCG), supercontinuum source (SS), polymer fiber coating (PFC), silanization treatment (ST) and bacteria detection (BD).</figcaption></figure>

<table><caption>**Table 3**
Optical fiber gratings-based sensors for detection of microorganisms and other organic substances. Note: Response time (RT).</caption><thead><tr><th>Analyte</th><th>Sensor</th><th>Sensitivity</th><th>Detection range</th><th>RT (s)</th><th>LOD</th><th>Ref.</th></tr></thead><tbody><tr><td>S. Aureus</td><td>LPFG</td><td>0.5 nm/log (CFU/mL)</td><td>10<sup>2</sup>–10<sup>9</sup> CFU/mL</td><td>1200</td><td>33 CFU/mL</td><td>[153]</td></tr><tr><td>NDV</td><td>TFBG</td><td>1.627 nm/(ng/mL)</td><td>0.025–0.2 ng/mL</td><td>1800</td><td>0.025 ng/mL</td><td>[154]</td></tr><tr><td>T7 phages</td><td>LPFG</td><td>0.367 nm/log (PFU/mL)</td><td>5 × 10<sup>3</sup>–5 × 10<sup>9</sup> PFU/mL</td><td>100</td><td>5 × 10<sup>3</sup> PFU/mL</td><td>[163]</td></tr><tr><td>Vitamin D3</td><td>LPFG</td><td>–1400 nm/RIU</td><td>1–1000 ng/mL</td><td>1200</td><td>1 ng/mL</td><td>[157]</td></tr><tr><td>Cocaine</td><td>LPFG</td><td>2.31 × 10<sup>–4</sup> nm/(ng/mL)</td><td>0.08–0.24 ng/mL</td><td>NG</td><td>0.08 ng/mL</td><td>[57]</td></tr><tr><td>Dopamine</td><td>TFBG</td><td>0.29 dB/log M</td><td>0.1–10<sup>4</sup> pM</td><td>NG</td><td>0.16 pM</td><td>[44]</td></tr><tr><td>Triacylglycerides</td><td>LPFG</td><td>7.8 × 10<sup>–7</sup> nm/(ng/mL)</td><td>1.278 × 10<sup>5</sup>–4.473 × 10<sup>6</sup> ng/mL</td><td>NG</td><td>1.278 × 10<sup>5</sup> ng/mL</td><td>[63]</td></tr><tr><td>Dalafenib</td><td>LPFG</td><td>1.52 × 10<sup>–5</sup> nm/(ng/mL)</td><td>7.5 × 10<sup>4</sup>–7.5 × 10<sup>6</sup> ng/mL</td><td>450</td><td>74.4 µg/mL</td><td>[158]</td></tr><tr><td>Carboxyfentanyl</td><td>LPFG</td><td>NG</td><td>50–1000 ng/mL</td><td>840</td><td>50 ng/mL</td><td>[159]</td></tr><tr><td>Cortisol</td><td>TFBG</td><td>0.275 nm/(ng/mL)</td><td>0.1–10 ng/mL</td><td>1200</td><td>0.018 ng/mL</td><td>[160]</td></tr><tr><td>Aquaporin-2</td><td>TFBG</td><td>1.4 × 10<sup>–3</sup> dB/(ng/mL)/min</td><td>0–14.2 ng/mL</td><td>900</td><td>1.5 ng/mL</td><td>[53]</td></tr><tr><td>C-reactive protein</td><td>LPFG</td><td>1.92 × 10<sup>–5</sup> nm/(ng/mL)</td><td>1–10<sup>5</sup> ng/mL</td><td>NG</td><td>0.15 ng/mL</td><td>[24]</td></tr><tr><td>Anti-cardiac troponin-I</td><td>tmFBG</td><td>0.28 nm/(ng/mL)</td><td>0.01–10 ng/mL</td><td>300</td><td>2.2 × 10<sup>–3</sup> ng/mL</td><td>[162]</td></tr></tbody></table>

Note: Newcastle disease virus (NDV), fiber Bragg grating (FBG), tapered microfiber Bragg grating (tmFBG), not given (NG), response time (RT) and low limit of detection (LOD).

concentrations of SARS-CoV-2 stinger protein, local changes in surface refractive index induced by antibody-protein binding caused the TiO₂ polystyrene coated LPFG to exhibit wavelength drift, and the low LOD of the biosensor was 0.1 ng/mL. Janczuk-Richter et al. [155] reported a highly sensitive immunosensor of LPFG working in phase-matching turning point for the detection of T7 phages. The fiber surface was functionalized with 3-(triethoxyphenyl) propyl succinic anhydride for covalent binding of anti-T7 antibodies, and the low LOD of the sensor reached to 5 × 10³ PFU/mL. Finally, a sensor for the detection of *Aspergillus Niger* in the air was developed by Gambhir et al. [156]. The sensor surface was functionalized with a monoclonal glucose oxidase IgG for target-specific covalent binding and was able to measure 10³ CFU/mL of pathogenic *Aspergillus Niger* fungal spores in the

environment.

## 5.2. Detection of other organic substances

To detect organic pollutants, glucose, protein, antibody-antigen, biotin and streptavidin, DNA, enzymes, and microorganisms, different FBG, LPFG, and TFBG biosensors with sensitive-materials had been developed. Currently, the LPFG and TFBG coupled with the advanced functional material have also been developed to monitor other organic substances.

For the LPFG-based biosensors, a double cladding fiber LPFG biosensor for vitamin D3 detection was reported by Esposito et al [157]. The sensor surface was coated with nano-sized graphene oxide for
binding as a biorecognition element was 25(OH)D₃, a specific antibody to vitamin D3. The reported low LOD for vitamin D3 detection was 1.0 ng/mL, with a detection range of 1 to 1000 ng/mL (Table 3). Celebenska et al. [57] proposed a cocaine biosensor in which the cocaine aptamer MN6 was immobilized on the LPFG after functionalization with amionopropyltriethoxysilane and p-Phenylene diisothiocyanate. As cocaine induces conformational rearrangement of the aptamer to occur, this changes the refractive index around the LPFG, which in turn leads to a transmission wavelength shift. The low LOD of this sensor was approximately 25 µM. Baliyan et al. [63] developed a lipase enzyme coated LPFG biosensor to detect the concentration of triacylglycerides. The high specific binding of lipase to triacylglycerides gives the sensor high selectivity and high sensitivity to triacylglycerides in human blood of 0.5 nm/mM with a minimum detection limit of 17.71 mg/dL. Furthermore, He et al. [158] combined a dalafenib-specific nanomolecularly imprinted polymer with LPFG to measure trace amounts of dalafenib. The molecularly imprinted polymer bound to 10–750 µg mL⁻¹ dalafenib in fetal bovine serum, leading to the sensor with wavelength shift of 11.87 nm. Liu et al. used carboxyfentanyl nano-molecularly imprinted polymers to functionalise LPFG sensors for detection of carboxyfentanyl drug. The developed sensor exhibited a progressive response when the carboxyfentanyl concentration was increased from 0 to 1000 ng/mL, with a minimum detection concentration of 50 ng/mL, corresponding to a wavelength shift of 1.20 ± 0.2 nm [159].

In the TFBG-based biosensors, Hu et al. [44] reported nanoscale gold-layered TFBG with a single layer of graphene. To improve the low LOD, an appropriate single-stranded DNA was selected as dopamine-specific aptamer modified on the fiber. Single-stranded DNA was stably immobilized on the fiber for highly sensitive detection of dopamine under the action of graphene. The change in the surface refractive index of the fiber was amplified when the single-stranded DNA aptamer encountered a dopamine molecule in solution leading to a change in the conformation of the aptamer molecule. The sensor responds linearly to dopamine concentrations in the range of 10⁻¹³ M to 10⁻⁸ M, with a low LOD of 0.1 pM. Leitao et al. [160] developed TFBG surface plasmon resonance immunosensors encapsulated by thin Au of nanoscale thickness, which were functionalized by anti-cortisol antibodies for cortisol detection. Intriguingly, they used a method based on monitoring the local maxima of the plasmonic signals to improve the sensitivity of the sensor. The biosensor showed a linear trend in the cortisol concentration range of 0.1–10 ng/mL and a sensitivity of 0.275 ± 0.028 nm·ng·mL⁻¹.

Guo et al. [49] designed a TFBG coated with a nanoscale silver film to detect changes in urine protein concentration. The prepared TFBG biosensor was embedded it in a microfluidic chip, therefore, the volume of the biological sample could be precisely controlled [49]. The device distinguishes protein changes in the urine of healthy, diseased and treated mice by monitoring changes in the fiber optic surface plasmon resonance and cut-off resonance spectra, the sensitivity and LOD of the sensor are 5.5 dB/(mg/mL) and 1.5 × 10⁻³ mg/mL, respectively. The team then built on the original urine protein assay by refining the detection target to Aquaporin-2, functionalising the sensor with an antibody to Aquaporin-2 to specifically detect Aquaporin-2 [53]. To reduce the size of the sensor, they modified the sensor to be reflective type by plating a gold mirror on the tip of the TFBG. The biosensor utilised 50 nm gold-coated TFBG to excite the surface plasmon resonance, resulting in improved stability and reusability compared to previous studies using silver films. They measured urine samples from normal and nephrotic rats in the laboratory with a detection limit of 1.5 ng/mL, and the urine samples from the nephrotic rats caused an intensity change of 0.3 dB in 15 min while the normal rats only caused a change of 0.15 dB [53].

Other optical fiber gratings-based immunosensors have also been developed by researchers, including the use of double-clad fiber LPFG to detect C-reactive protein in serum. A thin layer of graphene oxide nanoparticles was coated on the double-clad fiber LPFG to covalently immobilize the functional groups that capture C-reactive protein

antibodies [24]. The experimental results showed that the sensor detected C-reactive protein in the range from 1 ng/mL to 100 µg/mL with a low LOD value of 0.15 ng/mL. Recently, mFBG sensors were developed for detection of cardiac biomarkers (cTn-I). Ran et al. [161] specifically recognized cTn-I proteins by modifying a cTn-I antibody on mFBG and employed harmonic resonance to exclude the effect of temperature changes. The sensor was able to distinguish specific immune responses without affecting temperature changes and showed a low LOD of 13.5 ng/mL in the cTn-I assay. Afterwards, to further improve the sensor performance, in 2023 they [162] developed a unilaterally tapered microfiber FBG-based biosensor modified with Au nanostars and anti-cTn-I for the specific rapid and sensitive detection of anti-cardiac troponin-I. The modification of microfiber Bragg grating by Au nanostars increases the refractive index sensitivity by about a factor of two due to the significant enhancement of the electromagnetic field mediated by the star-shaped tip of the gold nanostar. The sensor had a low LOD of 2.2 pg/mL at a concentration of 10 ng/mL of anti-cardiac troponin-I. The detection time was shortened to less than 5 min because the binding of anti-cardiac troponin-I and anti-cardiac troponin-I was more rapid caused by the electrostatic attraction.

A mFBG biosensor for the quantification of placental growth factor (PLGF) in clinical serum samples is reported by Cao et al [18]. Functional modification of the PLGF antibody on the surface of the mFBG probe using a covalent functionalization technique enabled it to specifically detect the PLGF protein. In addition, an FBG was added for temperature calibration, allowing for dual-parameter detection of PLGF and temperature to eliminate crosstalk. The sensor could respond to PLGF concentrations ranging from 5 to 120 pg mL⁻¹ with a low LOD of 5 pg/mL. The biosensor is also expected to be coupled with microfluidic chips, which will significantly reduce the amount of clinical serum samples required.

Overall we found that the TFBG have different applications in sensing like detection of the microorganisms, organic pollutants, life organics, and other organic substances. To enhance the sensitivity of the TFBG sensors, their surface should be coated with a thin metal film (such as gold, silver, copper, aluminum, alkali metals, and sub-alkali metal film) to generate surface plasmon resonance before coupling with the sensitive material. Furthermore, the TiO₂ nanoparticles with high refractive index can be used to enhance the sensitivity of the LPFG sensors for detection of microorganisms. Therefore, the optical fiber gratings can collaborate with sensitive-materials and nanoparticles with high refractive index (such as GeO₂ with refractive index of 1.991 [163]) and transmittance to prepare highly sensitive sensors for detection of organic substances.

## 6. Conclusion and future prospects

In this review, we systematically summarize recent advances in the detection of organics using optical fiber grating-based sensors from 2016 to present. The principles of sensor' measurement of analytes, the composition of functional materials introduced into sensors, and the mechanisms to improve sensor performance are highlighted.

The review highlights that, in recent years, the environmental pollution problems accompanying the rapid industrial development worldwide have led to an increasing emphasis on environmental monitoring, biomedicine, food quality and safety, and other fields. The combination of functional materials and optical fiber gratings can be used to accurately detect antibodies, blood glucose, drug concentration, cancer biomarkers in the human body, hazardous gases, water pollutants and so on. Despite the great achievements in the development of sensors, there is still much room for research in repeatability, sensitivity, compatibility, and cost.

**Enhancing repeatability.** The strength of the bond between the sensitive material and the optical fiber gratings is one of the factors influencing the repeatability of the sensor. However, most of the surface functionalization schemes make use of chemical bonding, and this
bonding may be weakened by reacting with the analyte during the measurement. Chemical bonding combined with physical adsorption may be a good method to reinforce the attachment of the target identification sensitive-materials to the optical fiber gratings [92,93,100]. Furthermore, the surface of the optical fiber gratings has a suitable roughness, which can also be used to enhance the adhesion strength between the sensitive-materials and optical fiber and is beneficial for enhancing the sensitivity of the sensor. In practical applications, the repeatability of the optical fiber gratings-based biosensors is not only affected by the bonding strength between sensitive-materials and optical fiber, but also by the influence of stability of the sensitive-materials. The most materials and biorecognition molecules are affected by the temperature, humidity and pH, and once beyond the acceptable range of their performance will be drastically down. The excellent thermal and chemical stability of metal-organic frameworks can be used as a protective shell layer. The modifying material is wrapped in it to improve the stability of the material. Of course, poor repeatability of the optical fiber grating sensors is not just caused by the difficulty of repeating sensitive materials. In long-term stability studies, even if the optical fiber lacks a sensitive coating, optical fiber grating sensors exhibited the phenomenon of zero wavelength drift [164]. These may be some future directions worth delving into. In conclusion, the development of materials and biorecognition molecules that maintain stable properties under various environmental parameters should receive more attention.

**Enhancing sensitivity.** The sensitivity of the optical fiber gratings-based orgaincs sensors are controlled by the performance of the sensitive-materials and working modes of the optical fiber gratings. The larger specific surface area and high refractive index bring more opportunities for materials and analytes to combine, which can accelerate the rate and intensity of the reaction and is very beneficial to improve the sensitivity of the sensor. Therefore, optimisation of existing synthesis methods to decrease the size and crystal surface of nanoparticles with high refractive index is a future direction to be worked on. Furthermore, the working modes of the optical fiber gratings directly affects the sensitivity of the sensors and is controlled by their grating period. Therefore, adjusting the grating period of optical fiber gratings to make the sensors operating at phase-matching turning point is another effective measure to improve sensitivity.

**Enhancing compatibility.** Compatibility includes the biocompatibility of the optical fiber gratings-based sensors and the compatibility between the sensors and other devices. Due to the small size of optical fiber gratings-based sensors can be integrated into tiny devices to achieve online minimally invasive detection for organics parameters from the in vivo, in vitro, and biochemical reactions, and etc. Biocompatibility means whether the material will cause toxic effects on human tissue or the environment after implantation. It directly determines whether it can be used in the human body, clean water and microbial communities [165-167]. So, more bio-friendly functional materials need to be developed. Furthermore, to enhance the application fields of the optical fiber gratings-based sensors, trying to combine with other devices is necessary. Combine microlenses and optical fiber can generate photonic nanojets, which can enable three-dimensional manipulation of singlet nanoparticles and biomolecules. The ability to control individual dielectric nanoparticles can be utilized to optimize the nanostructures on the surface of optical fiber gratings, as well as to control the efficient binding of sensitive-materials to biomolecules, thereby improving sensor performance [168,169].

**Decreasing device cost.** Although fiber grating-based bio- and chemical-sensors have the characteristics of high selectivity, microstructure, and in situ decoding gaseous and liquid analytes, they require specialized detection instruments (laser sources, spectrometers, etc.) with high-cost. However, the paper-based sensors are lightweight and portable and can be quickly detected in the field without the support of expensive and complex equipment. Furthermore, electrochemical sensors can share an analysis platform to monitor multiple parameters. However, FBG, TFBG and LPFG detection instruments are not

compatible, and they all require specialized analytical equipment, which is another reason to increase the cost. Therefore, decreasing the complexity and increasing the compatibility of detection devices is another key to promote the practical application of fiber grating-based bio- and chemical-sensors.

### CRediT authorship contribution statement

**Xiaoyu Zhong:** Investigation. **Quanhua Xie:** Investigation. **Yang Liu:** Writing – original draft. **Yuanyuan He:** Writing – original draft. **Nianbing Zhong:** Writing – original draft. **Zhouxiang Zhang:** Writing – original draft. **Hassan Karimi-Maleh:** Writing – review & editing. **Writing – original draft.** **Xiaoling Peng:** Writing – original draft. **Eric Lichtfouse:** Writing – review & editing.

### Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

### Data availability

No data was used for the research described in the article.

### Acknowledgments

The authors acknowledge financial support from the National Natural Science Foundation of China (NSFC) (52176178, 52304321), Major Scientific and Technological Research Project of Chongqing Municipal Education Commission (KJZD-M202201101), Technological Research Project of Chongqing Municipal Education Commission (KJQN202301101), Chongqing Natural Science Foundation Innovation and Development Joint Fund (Municipal Education Commission) Project (CSTB2022NSCQ-LZX0059), Key R&D Projects of Sichuan Provincial Science and Technology Plan (2023YFS0455), Meishan Science and Technology Plan Project (2022ZYZF13), Chongqing Talent Project (cstc2022ycjh-bgzxm0241), Innovation Research Group of Universities in Chongqing (CXQT21035), and Chongqing University of Technology Postgraduate Innovation Project (gzlcx20223058).

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