1. Introduction

It is important to note that the long-period fiber grating (LPFG)-based biosensor is applied to measure the wavelength shift caused by the RI change of the device surface due to the selective adsorption for target molecules on the surface, rather than measuring the wavelength shift caused by the change of the RI of the bulk surrounding medium. Therefore, the functionalization of the LPFG-based biosensor is a fundamental step in realizing biochemical applications. Generally, two-part functional layers are deposited onto the surface of the LPFG-based biosensor: one part is the biocarrier layer used as the immobilization of the bioreceptor, and the other is the bioreceptor layer used as the recognition element (enzymes, proteins, antibodies, and so on) to selectively capture the target. Therefore, various methods have been employed for the immobilization of the bioreceptor layer onto the optical fiber. The comparison between the bioreceptor, target and performance of the different functionalization methods of LPFG-based biosensors can be seen in Table 1.

The most effective method for LPFG-based biosensor functionalization is based on covalent immobilization, due to its permanent attachment to the bioreceptor. The 3-aminopropyltriethoxysilane (APTES) silanization is a common and ideal method used in most chemical modifications of silica substrates. This method was successfully realized for the covalent immobilization of protein ^[27][28][101,102], DNA ^[6][29][67,68], antibodies ^[5][30][31][17,82,103] and so on.

In this case, the optical fiber requires a pretreatment step to form silanol groups (Si-OH), by immersing in KOH/NaOH, acid, or piranha solution. Similarly, ethoxy (–OCH₂CH₃) groups existing in the APTES molecule can also form the Si-OH through a hydrolysis reaction in aqueous environments ^[32][104]. Then, the condensation between Si-OH leads to the formation of a siloxane (Si-O-Si) bond, allowing the APTES molecules to immobilize onto the fiber surface. In addition, the adjacent APTES molecules can form a polymer matrix through condensation, resulting in the formation of free amino-functional (–NH₂) surfaces of silica substrates ^[33][105]. After the silanization is accomplished, there is a step of activation of thecarboxyl groups on antibodies or enzymes, with the aid of 1-ethyl-3- (3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) and N-hydroxysuccinimimide (NHS). Then, the antibodies or enzymes can be bonded to the –NH₂ groups via forming amide bonds, hydrogen bonding, or electrostatic interaction ^[34][106].

The group of Anjli proposed an enzymatic biosensor based on LPFG through the stable covalent binding of the lipase enzyme for the detection of triacylglycerides ^[1][80]. APTES silanization to immobilize glucose oxidase was also performed by Wu in ^[2][81]. In this work, glucose oxidase was immobilized onto the S-shaped LPFG by APTES silanization technology and used as the bioreceptor for the detection of glucose. The transmission loss variation was used as a measurement associated with glucose oxidase and glucose-specific binding. The experimental results show that the proposed sensor performed a sensitivity of 6.229 dB/wt% in a range from 0∼1 wt%. More recently, Gan et al. ^[3][39] developed an LPFG-based sensor based on egg yolk antibody (IgY) covalently immobilized by APTES silanization for the detection of Staphylococcus aureus. The detection test could be completed in about 20 min, and the detection of Staphylococcus aureus was performed down to 33 CFU/mL. Therefore, the developed device for the detection of Staphylococcus aureus was expected to be applied to the field of medical and food detection.

Although the amino functionalized fiber is widely proposed, the single functional group makes it unable to bond to other kinds of biological receptors, thus limiting its application. Different research groups have focused their attention on GO; the fiber surface is chemically or physically deposited with GO nanosheets after silanization. The GO is rich in oxygen-containing functional groups, such as epoxy, hydroxyl and carboxyl, which provide GO the capability to covalently bond various biomolecules ^[35][107]. Moreover, the GO is also endowed with the capability to adsorb biomolecules by noncovalent immobilizing, such as electrostatic, interaction hydrogen bonding and π–π stacking ^[36][37][108,109].

Chen et al. ^[9][46] reported a dual-peak LPFG deposited with GO nanosheets to immobilize the IgG and the IgG/anti-IgG as a bioconjugate pair for immunosensing. They adopted a new strategy to deposit GO that relied on chemical bonding followed by physical adsorption. Chemical bonding occurred between the APTES-silanized fiber and GO. In the meanwhile, GO nanosheets were physically adsorbed on the fiber surface, along with the evaporation of water. Finally, the GO-deposited LPFG-based sensor was immersed into IgG solution and covalently bonded together via EDC/NHS cross-linking chemistry. The detection of anti-IgG was obtained down to a concentration of 7 ng/mL in PBS buffer. The reusability of the sensor was also carried out by stripping off bound anti-IgG. Successively, the same group developed the detection of hemoglobin, based on a GO nanosheet-functionalized LPFG-based sensor ^[10][85].

Recently, the detection limits and deposition method have been further improved by Wang et al. in ^[11][86]. In this work, a micro-tapered LPFG was deposited with GO nanosheets for the detection of hemoglobin. After completing the chemical bonding followed by physical adsorption, they took advantage of the optical tweezer effect to further enhance the interaction between GO nanosheets and the fiber. The GO thickness of 203.6 nm was immobilized onto the micro-tapered LPFG. The LOD of 0.02 mg/mL in various interfering compounds was obtained. More recently, the same group developed a biosensor with the same structure for bovine serum albumin (BSA) detection ^[12][87]. The sensing mechanism relies on measuring the wavelength shift caused by the covalent bonding between the GO and BSA. The LOD of 0.043 mg/mL, 0.029 mg/mL and 0.032 mg/mL were achieved in DI water, urea and glucose, respectively. Similarly, they also proposed a micro-tapered LPFG functionalized by GO/polydopamine nanocomposites for cobalt ion sensing ^[13][88]; the nanocomposites were deposited onto the micro-tapered LPFG surface on account of APTES silanization followed by the optical tweezers effect. The proposed sensor showed a sensitivity of 2.4 × 10⁻³ dB/ppb in the cobalt ion concentration range from 1 ppb to 10⁷ ppb, and a detection limit of concentrations as low as 0.17 ppb was achieved. The micro-tapered LPFG-based sensors functionalized by GO have also been implemented for Na⁺ and Mn²⁺ ion detection in ^[38][110] and Ni²⁺ ion detection in ^[14][89].

The method of covalent modification of fibers introduced above has good stability; however, it is more complicated to control the thickness, such as controlling the reaction concentration and time. The realization of functional coating with a controllable thickness is also a factor to be considered when optimizing the sensing sensitivity of the LPFG-based device. The layer-by-layer (LbL) assembly provides a promising way to precisely deposit a functional coating with a nanometer-scale thickness ^[39][40][70,111], driven by electrostatic interactions between oppositely charged polyelectrolytes ^[41][42][112,113].

The group of Tian ^[15][90] developed a salinity sensor based on LPFG coated with ionic-strength responsiveness of chitosan (CHI)/poly (acrylic acid) (PAA) polyelectrolyte multilayers by the LbL assembly method. The entire deposition process was repeated 20 times by immersion of the device into a polycation CHI and polyanion PAA, respectively. Interestingly, the LPFG resonance wavelength shift changed from red to blue with increasing salt concentration. It could be explained by the de-swelling or swelling of the coating in response to a different range of NaCl concentration. The sensitivity of 36 nm/M was obtained in the range of 0.5–0.8 M. This research was also expected to apply to biomedicine and drug delivery. Similarly, the same group also proposed a salinity sensor which coated with ionic strength-responsive hydrogel onto the LPFG ^[16][91]. The two-component polyelectrolytes deposited by LbL assembly might cause some problems, i.e., pH cross-sensitivity and nonlinear relations between the resonance wavelength shift and the concentration of salinity.

Other polyelectrolyte functional coatings were also deposited onto LPFG-based biosensor by the LbL assembly method. An LPFG-based biosensor coated with nano-assembled thin film of poly (diallyldimethyammonium chloride) (PDDA) and tetrakis (4-sulfophenyl)porphine (TSPP) via the LbL technique for ammonia gas detection was fabricated by Lee et al. ^[17][92]. The group of Abd-Rahman fabricated a PDDA/poly (sodium-p-styrenesulfonate) (PSS)-Au nanoparticle coating layer onto an LPFG surface by using the LbL technique for mercury (II) ion sensing ^[18][93]. The designed sensor had an excellent performance in the mercury (II) ion concentration range of 0.5 ppm to 10 ppm. Liu et al. ^[19][64] developed an LPFG-based biosensor coated with poly (allylamine hydrochloride) (PAH)/gold-coated silica nanoparticles via the LbL method for the detection of streptavidin and immunoglobulin M (IgM). Ni et al. ^[20][94] investigated an LPFG-based sensor with a coating of poly (ethylenimine) (PEI) and poly (acrylic acid) (PAA) for pH sensing; the coating layer improved the dispersion and the adhesion ability of multi-walled carbon nanotubes. The group of Tian ^[21][95] considered PAH/PAA as a polyelectrolyte functional coating deposited by LbL assembly to bond specific antibodies for the detection of Staphylococcus aureus. The polyelectrolyte functional coating could facilitate the bacterial adhesion, and the detection with a LOD of 224 CFU/mL was demonstrated in PBS.

The method of optical fiber functionalization depends on the application environment to a certain extent. The functionalization method of optical fibers applied in air is more concise and simpler than those applied in aqueous solution. For example, the dip-coating technique, which is simple to operate and makes it easy to control the thickness of sensitive film, is widely used in gas sensing.

The group of Feng ^[22][96] reported the molybdenum sulfide/citric acid composite films that were deposited by using the sol–gel and dip-coating techniques onto an LPFG for measuring trace hydrogen sulfide gas.

In addition, the metal organic frameworks (MOFs), because of their excellent properties of tunable porosity, large internal surface area and organic functionality, have been widely applied for functionalizing LPFG-based sensors for gas sensing and other sensing. MOFs are hybrid crystalline nanomaterials composed of metal cations and organic ligands ^[43][44][45][116,117,118]. The methods of functionalizing optical fibers with MOFs mainly focus on in situ crystallization ^[46][47][119,120]. The group of Korposh developed an LPFG-based organic vapor sensor functionalized by zeolitic imidazole framework-8 (ZIF-8) films ^[48][121].