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Abstract
The purpose of this work is to propose a simple, portable, and sensitive biosensor structure based on singlemode fiber-multicore fiber-multimode fiber-singlemode fiber (SMF-MCF-MMF-SMF) for the detection of creatinine in the human body. Chemical etching has been used to modify the diameter of the sensing probe to approximately 90 μm in order to generate strong evanescent waves (EWs). The sensor probe is functionalized with graphene oxide (GO), gold nanoparticles (AuNPs), molybdenum disulfide nanoparticles (MoS2-NPs), and creatininase (CA) enzyme. The concentration of creatinine is determined using fiber optic localized surface plasmon resonance (LSPR). While EWs are used to enhance the LSPR effect of AuNPs, two-dimensional (2D) materials (GO and MoS2-NPs) are used to increase biocompatibility, and CA is used to increase probe specificity. Additionally, HR-TEM and UV-visible spectroscopy are used to characterize and measure the nanoparticle (NP) morphology and absorption spectrum, respectively. SEM is used to characterize the NPs immobilized on the surface of the fiber probe. The sensor probe's reusability, reproducibility, stability, selectivity, and pH test results are also tested to verify the sensor performance. The sensitivity of proposed sensor is 0.0025 nm/μM, has a standard deviation of 0.107, and has a limit of detection of 128.4 μM over a linear detection range of 0 - 2000 μM.
© 2021 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
Creatinine is a waste product of muscle metabolism in the human body that is primarily eliminated via the glomerulus. It is a critical clinical biomarker for diabetes, kidney disease, renal failure, and muscle atrophy [1,2]. A healthy person's serum creatinine concentration is 140 μM, however, if a person's serum creatinine level exceeds 800 μM, hemodialysis may be required to maintain normal renal function. As a result, serum creatinine levels must be monitored in patients with chronic kidney disease [3–5]. Previously, some researchers depended solely on colorimetric (Jaffe reaction) or enzyme colorimetric techniques to determine creatinine concentration [6,7]. Gupta et al. developed a fiber optic creatinine sensor based on molecular imprinting technology of MoS2/SnO2 nanocomposites [8]. This type of sensor is not suitable for high concentration detection. However, colorimetric analysis is time consuming and sensitive to metabolites and drugs in biological samples, and enzyme colorimetric analysis is prohibitively expensive for detecting creatinine, as metabolites are abundant [9,10]. As a result, it is critical to select a method that is cost effective, straightforward, rapid, highly selective, and sensitive. In comparison to conventional technologies, biosensors offer the following advantages: rapid response time, ease of operation, high sensitivity, and low cost of analysis [11–14]. Biosensors come in a variety of forms, including optical, electrochemical, spectrophotometer, and chromatographic sensors with several different mechanisms. Chromatography necessitates the use of sophisticated measuring instruments and sample handling procedures [15,16]. Electrochemical types necessitate a time-consuming electrode fabrication process [17,18]. Thus, researchers prefer optical sensors over other types of biosensors due to their unlabeled detection capability, rapid response time, and high sensitivity [19,20].
Kumar et al. developed an improved amperometric biosensor for the detection of creatinine using immobilized CA enzyme, creatinase (CI), and sarcosine oxidase (SOx) on glassy carbon (GC) electrodes [14]. This technique was successfully used to determine serum creatinine levels in healthy individuals and those with kidney and muscle dysfunction. Menon et al. developed a surface plasmon resonance (SPR) biosensor based on the Kretschmann principle for the detection of creatinine using nanocomposite gold films [21]. Topcu et al. developed SPR technology using N-methacryloyl-(L)-histidine methyl ester (MAH) as a functional monomer for unlabeled creatinine detection biosensors using molecular imprinting technique [22]. The optical fiber sensor based on SPR technology requires a small volume, a high resistance to electromagnetic interference, and a high sensitivity [19,23,24]. However, because the SPR fiber sensor is dependent on a uniform metal film, it is challenging to precisely control the metal thickness and uniformity on a non-planar substrate [25]. Due to the above limitations of conventional SPR sensors, their application is limited. The rise of nanophotonic technology and the rapid development of nanomaterial (NMs) synthesis techniques have significantly aided the development of LSPR sensors in recent years. LSPR sensors can be used in drug discovery, biological detection, cell labelling, site-specific diagnosis, molecular dynamics research, and disease diagnosis [26–29]. With the advancement of technology for processing metal nanomaterials and nanostructures, as well as the advancement of optical characterization tools, LSPR technology has gradually evolved based on SPR technology. LSPR is an optical phenomenon that occurs when incident light interacts with noble metal nanoparticles (NPs) whose size is less than the wavelength of the incident light [20]. When the incident light has a frequency that is consistent with the free electron oscillation frequency in the NPs, the interaction results in the collective oscillation of the surface electrons [20,28,29]. Due to the unique optical properties and biocompatibility of precious gold nanoparticles (AuNPs), that are similar in size to biomolecules, DNA, proteins, and nucleic acids, LSPR technology has been gradually applied to biosensor research [28,29]. Following that, due to the superior physical and chemical properties such as high electrical conductivity, low noise, and large surface volume, 2D materials have been widely used in the development and research of biosensors [26,30–32]. Due to the biocompatibility of 2D materials, they are particularly well suited for detecting small biomolecules at low concentrations. Biometrics can be used with a variety of biological materials, including enzymes [26,30,33], nucleic acids [34,35], and antibodies [36,37]. The analyte molecules begin to interact with the sensing layer during the biosensor process, causing changes in the refractive index (RI) of the medium. The biosensor's performance is dependent on the chemical and physical conditions in which the experiment is conducted, such as the active material's properties, thickness, temperature, pH value, and stability. The 2D materials GO and MoS2-NPs meet these criteria and are therefore suitable for biosensors [33,38].
In this work, multiple modes are excited when light is transmitted from SMF to MCF in the SMF-MCF-MMF-SMF structure. The excited mode is extremely sensitive to changes in the environment of the MCF cladding, resulting in an ultrasensitive sensor. As a result, immobilization the MCF with NMs provides an ideal platform for avoiding sensitivity limitations and stabilizing the sensing probe. Thus, this work discusses the development of a biosensor based on LSPR for the detection of human creatinine concentration. The evanescent field was enhanced by etching the MCF sensing region, and the sensing region was immobilized with GO, AuNPs, MoS2-NPs and CA to increase sensor's performance.
2. Experimental section
2.1 Materials
To fabricate the sensor fiber structures described previously, standard single-mode fiber (SMF, 9/125 μm), multimode fiber (MMF, 62.5/125 μm), and multicore fiber (MCF, 6.1/125 μm) were purchased from Shenzhen EB-link Technologies Co., Ltd in China, and Fibercore Ltd. in the United Kingdom, respectively. Sulfuric acid (H2SO4), graphite powder, sodium nitrate (NaNO3), potassium permanganate (KMnO4), hydrogen peroxide (H2O2), and hydrochloric acid are required reagents for the synthesis of GO. Trisodium citrate (Na3C6H5O7), hydrogen tetrachloroaurate (HAuCl4), and deionized (DI) water were used to synthesize AuNPs. To prepare MoS2-NPs, MoS2 powder purchased from Sigma-Aldrich was used. The fiber structure was etched with a 40% solution of hydrofluoric acid (HF) acid, and then cleaned with ethanol and nitrogen (N2) gas.
2.2 Instrument and measurement
The developed probe's light transmission characteristics were investigated using a tungsten halogen lamp (HL-2000) and a spectrometer (USB2000+) from Ocean Optics, USA. To fabricate the bare fiber structure, a fusion splicer machine (FSM 100P+, Fujikura, Japan) was used. The general fiber optic cleaver (CT-32) was used for cleaving optical fibers in dynamic scenarios, where no specific length was required. The use of a super large core diameter fiber optic cleaver (CT-106, Fujikura, Japan) to cleave an optical fiber to a desired length. To characterize NPs, the absorbance spectrum of GO, Au-NPs, and MoS2-NPs was measured using a UV-Vis spectrophotometer (U-3310, Japan) to determine the diameter range of the synthesized NPs, in general. High resolution transmission electron microscopy (HR-TEM, Talos L120C, Thermo Fisher Scientific) was used to observe the geometrical shapes of NMs in aqueous solution. Scanning Electron Microscopy (SEM, Gemini and Carl Zeiss microscopy) was used to characterize the surface of the NPs-coated probe, and SEM-EDS was used to determine the composition of the surface substances in the micro-area.
2.3 Sensing mechanism of the probe
The generation of evanescent waves (EWs) is critical for the advancement of the LSPR sensor. On the one hand, EWs can induce LSPR and generate absorption peaks associated with RI. On the other hand, the evanescent field can produce sufficient overlap with the analyte to be measured in space to enable perception of biological reaction behaviour. However, conventional SMF or MMF cannot generate an effective EWs because light propagates forward through the total internal reflection at fiber core-cladding surface, limiting the evanescent field to the inner boundary of the cladding. Thus, to achieve an effective evanescent field, a particular optical fiber configuration must be used to break the inherent light field modes of SMF and MMF. The sensor structure was designed and developed using LSPR technology by combining the SMF-MCF-MMF-SMF structure with GO, AuNPs, and MoS2-NPs. When light is transmitted from the core of SMF to the core of MCF, a portion of the light enters the central core of MCF while the remaining part enters the other core of MCF due to refraction. The core mismatch between SMF and MCF is conducive to the excitation of several modes of MCF and transmission across its remaining core’s diameters. The loss due to mode field mismatch is expressed as [39,40];
Here, ${I_i}$ and ${I_j}$ are the power distributions of modes of ${i^{th}}$ and ${j^{th}}$. With the change of RI in the external environment, the sensing spectrum will also change, resulting in the shift of peak wavelength. Etching of fiber can increase the optical signal's penetration depth (${d_p}$) into the cladding, thereby stimulating EWs interaction with the surrounding medium. ${d_p}$ is equivalent to [40];
where, $\lambda $ is the wavelength of the optical signal, ${\theta _i}$ is the incident angle of the core-cladding interface, ${n_{core}}$ and ${n_{clad}}$ are the RIs for the core and the cladding, respectively. As previously stated, the light encircled the MCF's core, and thus spliced with the MMF (62.5 μm/125 μm). It also has a higher signal to noise ratio (SNR) due to its larger core diameter and numerical aperture (NA). This contributes to the majority of MCF transmitted light (core and cladding modes) being captured by the MMF core. Finally, the MMF-SMF is used to convert the multimode signal to a single mode signal, that is received and demodulated by the spectrometer.2.4 Fabrication of sensor probe
The sensor is fabricated by splicing a small section of the MCF and MMF to the SMF. The fabrication process of SMF-MCF-MMF-SMF based fiber Mach-Zehnder Interferometer (MZI) structure is shown in Fig. 1. MCF was chosen as the sensing probe due to a number of advantages, including low loss when combined with MMF and SMF, and high sensitivity to small RI changes, that enables the fabrication of a compact structure and high sensitivity sensing probe [26,41]. In comparison to conventional optical fibers, MCF has multiple fiber cores in its cladding layer, that enhances the high degree of freedom of fiber parameters. It is a multifunctional guided wave system, that also enables the development of a wide variety of sensing probes using MCF. Due to the close match between the NA of the MCF and the SMF, the insertion loss between the MCF and the SMF and MMF is typically less than 0.1 dB. The other end of the MCF is spliced to the MMF. The MCF and MMF have a length of 15 mm and 20 mm, respectively. Its operation is based on the modulation of the MCF mutual core's coupling coefficient, that makes the structure suitable for biological applications. The MCF used in the fiber optic sensor structure is composed of seven identical Ge-doped silica cores, one of whom is surrounded by six hexagonally arranged cores. In comparison to the conventional SMF-MMF-MCF-MMF-SMF structure, the proposed SMF-MCF-MMF-SMF structure significantly reduces fiber loss due to MMF fusion. We attempt to use SMF to perform the function of MMF in the SMF-MCF-MMF-SMF structure to obtain core and cladding modes. Partial core mode travels through the center core of the MCF fiber, while the majority of the SMF core mode (due to core mismatch) and SMF cladding mode travel through the remaining six cores of the MCF fiber. This way not only is splicing loss reduced, but also multimode signal transmission is unaffected.
Fig. 1. Schematic of the fabrication process for a fiber structure composed of SMF-MCF (1.5 cm)-MMF (2 cm)-SMF-based sensor structure.
2.5 Synthesis process of GO, AuNPs, and MoS2-NPs
The GO powder was obtained via a modified Hummer’s method [42,43]. Further, it was pulverized in an ultrasonic machine to obtain GO solution. Turkevich method was used to synthesize spherical colloidal AuNPs with a particle size of around 10 nm [44,45]. The AuNPs solution was stored at room temperature. For MoS2-NPs, N-methylpyrrolidone (NMP) was introduced as an organic solvent and used to dissolve 30 mg MoS2 powder using ultrasonic machine [26,46].
2.6 Nanocoating and enzyme functionalization
Acetone was used to clean and smooth the probe sensing area, removing most of the organic dust. The cleaned sensing area was then immersed for 30 minutes in piranha solution (H2O2:H2SO4 = 3:7). Finally, the probe was rinsed with deionized water and dried for 30 minutes in an oven set to 70°C. Salinization of etched optical fibers is accomplished by immersing the sensing region in a 5% ethanol solution of 3-(trimethoxymethilyl) propyl methacrylate (TMSPMA) for 30 minutes. The silane agent TMSPMA was used to immobilize the GO to the optical fiber structure. After that, the fiber was placed in a 70°C oven for 30 minutes to stabilize the fixed silane layer. After immersing the salinized fiber in a GO aqueous solution for 5 minutes, it was dried in a 70°C oven for 30 minutes. The procedure was repeated three times to ensure that the GO was distributed evenly across the fiber's surface.
The sensor probe with the GO layer was immersed in an ethanolic solution containing 1% MPTMS for 12 hours, and the sulfhydryl groups in the sensing region were modified. To remove the unbound MPTMS monomers, the optical fiber probe is rinsed with ethanol and then dried with nitrogen gas. After that, the sensing region was immersed in AuNPs solution for 48 hours. Because the MPTMS contained mercaptan groups, the AuNPs formed covalent bonds with it, that held the AuNPs to the surface of the etched fiber structure. To remove unfixed AuNPs from the surface of the fiber probe, it was rinsed with ethanol and then dried with nitrogen gas.
Thereafter, sensing area was immersed in MoS2-NPs solution for 20 seconds before being oxidized for 2 minutes at 50°C. This procedure is carried out eight times. Finally, to immobilize MoS2-NPs, the fiber was quenched in an oven at 50°C for 2 hours.
Before enzyme functionalization, the etched fiber surface was rinsed with DI water. After that, the fiber surface was carboxylated for 5 hours by immersing it in 11-Mercaptoundecanoic Acid (MUA) ethanolic solution (5 mL, 0.5 mM). To make the N-hydroxysuccinimide (NHS) ester, the carboxyl group was activated by adding fibers in ethyl (Dimethylaminopropyl) carbodiimide (EDC) (5 mL, 200 mM) and NHS (5 mL, 50 mM) solution for 30 minutes. Finally, the sensing region was immersed in CA enzyme solution (20 mL, 0.1 mM). The NHS ester formed a stable covalent bond with CA's primary amine, allowing the enzyme to functionalize the sensing probe. Figure 2 depicts the probe's functionalization process and its complete schematic.
Fig. 2. Schematic of nanomaterials-immobilization and enzyme-functionalization over optical fiber sensor probe.
2.7 Preparation of analytes solution
Creatinine is normally found in concentrations of less than 140 μM in the human body [3–5]. However, in patients with uremia, the creatinine level can exceed 2000 μM. To validate the biosensor's practical utility, solutions with concentrations ranging from 0 to 2000 μM were considered. There are eleven different concentrations were prepared using a 1×PBS buffer solution including 0 μM, 200 μM, 400 μM, 600 μM, 800 μM, 1000 μM, 1200 μM, 1400 μM, 1600 μM, 1800 μM, 2000 μM. 1×PBS buffer solution that dissolves protective reagents. The reason for this is that 1 ×PBS has the function of salt balance and adjustable suitable pH buffer.
2.8 Experimental setup
Figure 3 shows the experimental setup for estimating creatinine solution. The sensor probe was connected to the light source and the spectrometer, and the spectrometer sent the signal to the computer to be analyzed.
A tungsten-halogen light source was used to stimulate the LSPR effect of AuNPs. Creatinine binds specifically to CA enzyme on the probe's surface and decomposes into urea and sarcosine, resulting in a change in the surrounding RI. The absorption peak of the LSPR spectrum undergoes a red shift at this point in time, and the concentration of the creatinine sample can be determined further by the wavelength shift, that is designed to allow for creatinine concentration detection.
3. Results and discussions
3.1 Optimization of an optical fiber sensor probe
Numerous modes of cladding can be excited by various types of fiber coupling [47,48]. The output power is determined by the relative phase difference between each cladding mode excited by the coupling fiber, that is highly dependent on its length. This experiment makes use of the optical fiber structure SMF-MCF-MMF-SMF-based structure. Since the length of the coupling fiber has a variable effect on the sensor probe's transmission intensity and the intensity of the EWs in the cladding. We fabricate MCF-MMFs with varying length ratios and measure the transmission intensity. Three fibers with the same length ratio are prepared for the purpose of determining their average transmission intensity to minimize the error caused by fiber splicing loss. The results of the measurement are depicted in Fig. 4. When the length ratio of MCF-MMF is between 15 and 20 mm, the transmission intensity of the optical fiber is at its lowest. In theory, there will be more EWs in the optical fiber cladding at this time, that means that the fiber probe with this ratio will be more advantageous for achieving sensing performance. As a result, the following procedure will use a fiber optic probe of this length.
When the MCF-MMF structure is immersed in HF acid solution to etch away some of the cladding, the ratio of core to cladding increases. It enhances EWs while facilitating modal coupling. The MCF-MMF fiber was immersed in the HF solution in the petri dish for 5 minutes, 10 minutes, 15 minutes, 20 minutes, and 25 minutes, respectively, before being cleaned with ethanol. Figure 5(a) and (b) show the maximum transmission intensity and diameter scanning images of the etched SMF-MCF-MMF-SMF structure, respectively. Figure 5 (c) shows CCD images of the MCF cross-section at different etching time. The cladding of the MCF structure was dissolved in HF acid, and the etching time of 20 minutes resulted in the lowest transmission intensity and an intact MCF core. Figure 5(d) shows the diameter uniformity (about 90 μm) of the three fibers etched by HF and the reproducibility of the fiber structure. MCF structures that have been etched are more appropriate to modal coupling. SMF modes are coupled into the MCF structure's seven cores and then into the MMF. As a result, in all experiments, the SMF-MCF-MMF-SMF structure was etched for 20 minutes.
Fig. 5. (a) Transmitted intensity through etched fibers, (b) scanning diameters of etched fibers, (c) cross-sectional view of etched MCF’s section taken with a CCD, and (d) uniformity of etched fiber diameter.
3.2 Characterization of nanoparticles
UV-vis spectrophotometer and TEM were used to characterize the GO aqueous dispersion, AuNPs solution, and MoS2-NPs solution. GO, AuNPs, and MoS2-NPs had absorption peaks at 230 nm/310 nm, 519 nm, and 330 nm, respectively. Figures 6(a), 7(a), and 8(a) show the absorption spectra of the respective NMs. This ensures that the sensor probe's transmission spectrum's peak value is in the visible region. TEM was used to examine the distribution and shape of NMs in aqueous solution, and TEM images of GO, AuNPs, and MoS2-NPs are shown in Figs. 6(b), 7(b), and 8(b), respectively.
3.3 Characterization of nanomaterial-immobilized structure
SEM was used to observe the arrangement of GO, AuNPs, and MoS2-NPs on the surface of optical fiber structure. Multiple resolutions were used to observe the various morphologies. The sensor probe was observed in its entirety at a ratio of 29× and the probe's diameter was determined to be 90 ± 1 µm, as shown in Fig. 9(a). At a 5 K× ratio, the thin film region of GO was observed. As illustrated in Fig. 9(b), GO was fixed to the optical fiber's surface. The strong absorption peak at 230 nm of GO corresponds to the conjugated conversion of the aromatic C = C bond π-π*, and the weak peak at 310 nm corresponds to the n-π* conversion of the C = O bond. The 2D lamellar structure of GO transverse extension is beneficial to the immobilization of AuNPs [49]. As illustrated in Fig. 9(c), AuNPs are uniformly attached to the GO surface, and uniform and dense AuNPs are more favorable to LSPR excitation. As illustrated in Fig. 9(d), MoS2-NPs are immobilized on the surface of AuNPs. In the lattice structure of MoS2-NPs, a layer of molybdenum (Mo) atoms is covalently bonded between two layers of sulfur (S) atoms, that makes the atomic structure more stable, the electron mobility increases and the surface volume ratio is larger [50]. This is equivalent to increasing a spatial structure with a large specific region to increase the number of interaction sites of target biomolecules and improve the sensitivity of the probe.
Fig. 9. SEM images (a) sensor probe, (b) GO-coated, (c) GO/AuNPs-coated, (d) GO/AuNPs/MoS2-NPs-immobilized sensor structure.
SEM-EDX was used to examine the probe region at various stages. Figure 9(b) was analyzed using an energy spectrum, and the results are shown in Fig. 10(a). The presence of GO is evidenced by the relatively high carbon (C) and oxygen (O) levels. The presence of manganese (Mn) is due to the KMnO4 residue used as an oxidant in the synthesis of GO. The presence of Pt is the result of spraying to increase electrical conductivity during SEM scanning. The energy spectrum analysis results for the area depicted in Fig. 9(c) are represented in Fig. 10 (b). The sheer volume of Au, C, and O elements demonstrates that AuNPs are adhered to the GO surface. The energy spectrum results for the region depicted in Fig. 9(d) are depicted in Fig. 10(c). The presence of S and Mo elements demonstrates that MoS2-NPs are bonded to the surface of GO/AuNPs.
Fig. 10. SEM-EDX images (a) GO-coated, (b) GO/AuNPs-coated, (c) GO/ AuNPs/ MoS2-NPs-immobilized sensor structure.
3.4 Measurement of analytes
The LSPR spectrum was determined by performing a response test on creatinine samples of various concentrations using the experimental setup illustrated in Fig. 3. Three distinct sensing probes were used in the experiment to determine responsiveness in order to minimise experimental error and avoid the occurrence of contingency.
The LSPR spectrum obtained by averaging three independent sensor probes yielded the LSPR spectrum shown in Fig. 11(a). As the concentration of creatinine in the sample increases, the peak wavelength of the sensing spectrum shifts to the red. This indicates that the combination of the creatinine solution sample and the CA enzyme on the sensor probe's surface alters the RI of the surrounding medium, resulting in the sensing spectrum's wavelength drift. Additionally, by analyzing the test results from the three sensor probes, the probe sensitivity, linearity, and other related performance parameters were determined. Figure 11(b) illustrates the sensor probe's linear fitting results. The linear regression equation for it is as follows:
where C is the concentration of sample solution, sensitivity is 0.0025 nm/μM, correlation fitting factor R2=0.9738. This demonstrates that the developed sensor probe has a good linear fit.Fig. 11. LSPR sensing of creatinine solutions (a) transmission spectra, (b) linearity plot of proposed sensor.
3.5 Reproducibility and reusability test
Reproducibility and reusability of sensor probes are critical criteria for their practical application and performance indices. Two sensors from the same batch were used to measure creatinine samples at a concentration of 2000 μM to ensure reproducibility. Figure 12(a) illustrates the LSPR spectra obtained from the two sensing probes. The wavelength positions of the peak are 663.904 nm and 661.859 nm, respectively, with only a 0.3% error. After being washed with 1×PBS solution, the sensor probe can perform several low errors or even no error measurements for the reusability test. The same sensing probe was used to measure creatinine samples of 1000 μM and 2000 μM. The samples above were remeasured after being rinsed with a 1×PBS solution. Figure 12(b) illustrates its measurement spectra in support of reusability test. The spectrogram demonstrates that when the sample solution is measured at 1000 μM and 2000 μM, a wavelength shift occurs. Additionally, the spectra of the same concentration of sample solution were measured using the same peak wavelength. It is demonstrated that the sensor probe can achieve high levels of reusability and reproducibility.
3.6 Stability and pH test
Stability is a critical measurement property of sensor probes developed for measurement. This value is the foundation for ensuring the measurement value's accuracy. The sensor probe was used to measure 1×PBS solution ten times to determine its initial peak wavelength and to investigate the sensor probe's stability. Figure 13(a) illustrates the peak wavelength results of multiple-time PBS measurement. As illustrated in Fig. 13(a), there is only few deviations in ten-times measurements. The standard deviation (SD) and limit of detection (LoD) were determined to be 0.107 and 128.4 μM, respectively, indicating that the sensor probe was stable. LoD is the sensor's ability to detect minimal changes in analyte concentration. The calculation formula is: LoD = (3 × SD)/ sensitivity [26].
A 1×PBS solution was used as the solvent in the experiment to recreate the pH (7.4) environment found in the normal human body. 1×PBS solution not only has the same pH environment as human body, but also can protect the reagent from being damaged. To ascertain the applicability of 1×PBS, various pH environments as a reference were used. Acetic acid or potassium hydroxide solutions with pH values of 4, 6, 8, and 14 (measure by pH test indicator litmus paper) were used as solvents to prepare 2000 μM creatinine sample solutions and to determine the wavelength shift of their LSPR spectra. Figure 13(b) illustrates the pH test results. Obviously, 1×PBS solution with a pH of 7.4 exhibits the greatest wavelength shift, demonstrating the rationale for using 1×PBS solution as the solvent.
3.7 Selectivity test
Creatinine is mainly found in human serum, but serum also contains uric acid, ascorbic acid, creatine, sarcosine, and pyruvate. Considering the presence of these interfering substances in practical applications, selectivity tests were performed to verify the specificity of the sensor probe. The selectivity test results are shown in Fig. 14. When the concentration difference is 2000 μM, the wavelength shift of creatinine sample solution is more than 5 nm, while other interfering substances are only about 1 nm, or even less than 1 nm. The absolute specificity of the enzyme determines that CA primarily acts on creatinine. The combination of these produces the following chemical reaction as follow [51]:
Fig. 14. Sensor probe selectivity test in the presence of different biomolecules found in human body.
Because of the sensor's specificity, it is nearly impossible for CA enzyme to bind to substances other than creatinine and cause decomposition. This also explains the developed sensor probe's specific selectivity.
3.8 Evaluation of sensing performance
It was found that the use of AuNPs could stimulate the LSPR phenomenon, while the use of GO/MoS2-NPs could increase the binding site for CA enzyme. This comparison is made between the developed biosensor and commercially available sensors in the following areas: experimental mechanism, material composition, linear range, sensitivity, and detection limit. The results are shown in Table 1. When compared to earlier traditional biosensors, the proposed optical fiber biosensor has the advantages of being simple to operate, having a short response time, having a high sensitivity, and being inexpensive. As a result, optical sensors are becoming increasingly popular among researchers.
Table 1. Comparison of the proposed sensor's performance to that of existing sensors.
4. Conclusion
This experiment proposes a sensor model based on the SMF-MCF-MMF-SMF structure. To increase the evanescent waves generated by the LSPR effect, this type of fiber structure was etched with HF acid to achieve a thickness of 90 μm. To maximize the efficiency of LSPR, GO, AuNPs, and MoS2-NPs were sequentially immobilized on the sensing surface of the probe. Finally, a layer of CA enzyme was functionalized onto the probe's surface and combined with the characteristic form of creatinine present in the sample solution to determine the concentration of creatinine. The synthesized nanomaterials were characterized using HR-TEM, and their absorption spectrum was determined using UV-Vis Spectrophotometer to ensure that their peak value was in the visible light region. Additionally, SEM was used to characterize the NMs that were successfully fixed to the surface of the fiber sensing probe. Various creatinine solution concentrations were measured during the testing phase, as well as the sensor probe's reusability, reproducibility, stability, selectivity, and pH test, that all yielded favourable experimental results. When the creatinine sample solution concentration was between 0 and 2000 μM, the probe had a sensitivity of 0.0025 nm/μM, a SD of 0.107, and a LoD of 128.4 μM. The developed creatinine biosensor is capable of measuring creatinine concentrations comparable to those found in patients with severe uremia. It enables the rapid detection of creatinine levels in the human body, enables early detection and treatment, and decreases the likelihood of chronic kidney disease deterioration.
Funding
Double-Hundred Talent Plan of Shandong Province; Natural Science Foundation of Shandong Province (ZR2020QC061); Special Construction Project Fund for Shandong Province Taishan Mountain Scholars; Fundação para a Ciência e a Tecnologia (UIDB/50025/2020, UIDP/50025/2020); Fundação para a Ciência e a Tecnologia (CEECIND/00034/2018); Liaocheng University (31805180301, 31805180326, 318051901).
Acknowledgments
This work was supported by Special Construction Project Fund for Shandong Province Taishan Mountain Scholars, China. C. Marques acknowledges Fundação para a Ciência e a Tecnologia (FCT) through the CEECIND/00034/2018 (iFish project), and this work was developed within the scope of the project i3N, UIDB/50025/2020 & UIDP/50025/2020, financed by national funds through the FCT/MEC.
Disclosures
The authors declare no conflicts of interest.
Data availability
Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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