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Abstract
Whispering gallery modes (WGM) have revolutionized the field of optical sensors. This paper presents a design and simulation of a novel structure called a “multi-core whispering gallery mode (WGM)” based on multiple evanescent waves coupling for detecting cancer cells and diabetes tear cells. This work is totally simulation based, and the simulation is done by a finite element method based simulation tool. From the simulation, it is expected that the proposed sensor exhibits a sensitivity of 650 nm/RIU, 666.67 nm/RIU, and 642.285 nm/RIU, respectively, for the detection of cancerous Basal, HeLa, and MDB-MB-231 cells. In addition, it is also capable of detecting affected diabetes tear cells from healthy tear cells with a sensitivity of 650 nm/RIU. To the best of our knowledge, the resultant sensitivity of the proposed sensor is probably the highest compared to other WGM based bio-sensors till now.
© 2021 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
1. Introduction
Optical biosensors have recently been considered transducers because of their capability to detect the label-free molecule's presence due to medium or analyte refractive index change. Any perturbation in the sensing medium's optical properties can be detected easily using these optical based sensors. For instance, Mach-Zehnder interferometers [1], microsphere resonators [2–4], surface plasmon resonance (SPR) based Photonic crystals fibers [5]. Among these optical sensors, the Whispering Gallery Mode (WGM) concept is well-known in bio-sensing applications. Lord Rayleigh first characterized the Principal of whispering gallery mode (WGM) in 1910 in St Paul's Cathedral for acoustic waves [6]. Later, the exact total internal reflection of any iso-intensity surface due to the contrast in the medium's refractive index is circumscribed for light waves. In addition, the WGM cavity gives a much higher quality factor than the traditional Fabry-Pérot resonators. When a light wave travels in a dielectric medium on circular shaped structure, the WGM occurs due to total internal reflections (TIR) at the curved surface boundary as the electromagnetic fields can close on itself by giving rise to the resonances [7]. In recent times the most WGM based resonators are spherical, cylindrical, rings, spheroidal or toroidal, and other circular shapes with various confinement principles [8,9]. In order to couple light in or out of the microspheres, it is very important to utilize an overlapping evanescent radiation field of the phased matched optical waveguide. The WGM light intensity spectrum curve has several sharp resonant peaks and has a very stable interval between two successive peaks. The resonance wavelength shifts due to the optical properties of the surrounding medium changes. As WGM is a total reflection based phenomenon, so the critical angle $({\mathrm{\theta }_\textrm{c}})$ is directly dependent on the surrounding medium refractive index $({{\textrm{n}_\textrm{s}}} )$ and the resonator core refractive index $({{n_m}} )$ relation shown in Eq. (1). As per the medium refractive index $({{n_s}} ) \textrm{increases}$ the critical angle increases; this results in the increment of the effective traveling distance of photon in resonance mode. Therefore, a longer wavelength is required to match the resonance condition to keep the same resonance mode [7].
Applying the phenomenon mentioned above of resonance wavelength increment due to the optical property’s variations of the surrounding medium, several applications and sensors based on the WGM phenomenon exist today. For instance, detection of bio-molecules, absorption of different molecules, filters, modulators, lasers [10], detection of the refractive index using backscattering light data, and the dielectric particles [11], the existence of a single virus, protein binding, and other molecular level ingredients can be located using the shift in resonance wavelength [4–13]. Again, Due to its small-sized cavity and high-quality factor, WGM based sensor opens a new dimension in optical biochemicals and bio aqueous sensing [14]. WGM-based sensors primarily operate on evanescent wave coupling. The structure is coupled using evanescent waves from the outside beam source, and for a particular frequency, the resonance condition is achieved. A resonance peak is created in the intensity of the scattering field in resonance wavelength, and by observing resonance wavelength shift for medium refractive index variation, the unknown component or solutions are being measured [7,15]. WGM sensing using refractive index change and evanescent wave coupling for the increased absorption and scattering field in an aqueous medium is explained in [16]. Besides some other applications regarding WGM sensing used in radio frequency (RF) communication, quantum optics and electric sensing are demonstrated in [8,17]. Moreover, even in free space, the WGM coupling is also possible using specific asymmetric structures where the mode matching of adjusted Gaussian beam allows an efficient excitation called “chaos-assisted dynamical tunneling;” details are explained in [18,19].A tunable laser source connected with a single mode fiber (SMF) has to be utilized to excite the resonator core shown in Fig. 1. Here we have to vary the wavelength of the laser source to find the resonance wavelength. At resonance, the constructive electromagnetic field in the resonator core builds up, which leads to increasing the scattering (radiation) through the resonator [7]. In the output, the scattered light from the resonator is filtered using a laser line filter and then imaged using a photomultiplier tube (PMT) [15] by employing a pair of lenses. PMT is connected to an optical spectra analyzer (OSA) to view the spectrum. A linear polarizer placed between the sphere and the front collecting lens is used for the identification of the observed mode, such as transverse electric (TE) or transverse magnetic (TM) [15]. The in and out ports are shown in the diagram in Fig. 1 used for maintaining the flow of preserve the analyte or biological solution in cladding which needs to be measured using our sensor [20,21]. When the cladding of the resonator is filled with biological solution, and as the resonator core is highly sensitive to the medium refractive index change, it can easily detect the specific biological component. By observing the shift of resonance wavelength due to biological solutions refractive index variation, the sensitivity of the proposed sensor is determined. Although our proposed model is totally 2D numerical based, a practical implementation may also be possible following the above process. A similar model having single-core point source and 2D simulation with nonstandard finite-difference time domain (NS-FDTD) method of WGM model also demonstrated in [22]. Some similar ways of solving 3D models in 2D methods are shown in [23]; moreover, the 3D whispering gallery mode resonator's modeling and efficiency in 2D using FEM are shown in [24,25].
Fig. 1. Basic probable setup of a practical sensing scheme for our proposed model of multicore sensor.
In this paper, we have designed and simulated a novel structure named “Multi-core Whispering gallery mode (WGM)” sensor based on multiple evanescent waves coupling using finite element method (FEM) based software “COMSOL Multiphysics” and measured output resonance wavelength shift due to refractive index variation of biological solutions. Later, we have plotted the graph from the simulation data and measured the wavelength sensitivity and response of our sensor for refractive index change. To the best of our knowledge, the resultant wavelength sensitivity is probably the highest compared to other WGM based bio-sensors. Furthermore, we have used our sensor to detect cancer cells from healthy cells and affected diabetes tear cells from healthy tear cells using a refractive index variation of the solutions due to resonance wavelength shift sensitivity.
1.1 Cancer cells
The generic term “cancer” is applied to define a large group of diseases due to the rapid creation of abnormal cells that grow beyond their usual boundaries, which spreads through other parts of the body and damages organs. According to the World Health Organization (WHO), cancer has caused nearly 10 million deaths worldwide in the year 2020 [26]. Usually, cancer arises from transforming a normal cell into tumor cells in a multi-stage process that generally progresses from a pre-cancerous lesion to malignant tumors. According to the UK cancer research group [27], spotting cancer in the early stage enhances the chance of survival. More than nine in ten bowel cancer patients survive the disease for five years or more if patients are diagnosed at the earliest stage in England. Among the most common cancer types, Basal, HeLa, and MDB-MB-231 cancers are dominating mostly. Basal cancer cells are formed in the outer layer of the epidermis (skin) due to intense sun exposer, and among 3 million non-melanoma cancers (NMSC), the basal cell carcinomas are the major among them [28]. Again, HeLa healthy cells are not very uncontrolled; in fact, they are human cervical cells. However, due to the Human papillomavirus 18 (HPV18) infection, these cells become uncontrolled compared to normal conditions [29]. Furthermore, human chest cells called MDA-MB-231 are separated from pleural effusion from patients having breast cancer. Pleural effusions are regarded as a reliable source of breast tumor cells because it generally provides a plentiful supply of tumors cells and fibroblasts were either rare or absent, and the fluid usually showed the viability of 70% or better [30]. The mentioned cells show refractive index deviation when cancer cells start growing in them. For example, healthy basal cells have a refractive index of 1.36, and again cancerous cells have a refractive index of 1.38. Similarly, HeLa healthy cells have a refractive index of 1.368, and HeLa cancerous cells have a refractive index of 1.392, and MDB-MB-231 healthy cells have a refractive index of 1.385, and cancerous have 1.399 [5,31].
1.2 Diabetes
When the body's insulin production is not sufficient for beta pancreatic islet cells, the body cells cannot respond appropriately to blood sugar levels, causing increased blood glucose or blood sugar levels, leading to severe diseases. A statistical analysis by the World Health Organization (WHO) indicates that more than 220 million people have to live with diabetes, and in 2014, 8.5% of adults aged 18 years and older had diabetes. In 2019, diabetes was the direct cause of 1.5 million deaths [32]. So, continuous measurement of diabetes is essential. People with diabetes have to puncture their fingers multiple times to check the glucose level percentage in the blood, and a study shows it is almost 1800 times per year. However, a new technique has recently been developed to measure glucose levels through an alternative approach measuring body fluid accessibility like saliva, urine, and tears. Compared to other body fluids, tear cells are more accessible than blood. It is easily obtainable and less susceptible to dilution than urine and details optical measurement approaches given [33]. This [34] article confirms that a diabetic person can be separated from a regular healthy person measuring fluid tear cells. Employing Atago's refractometer by taking water as a base, the average healthy person's tear cell refractive index value is 1.35, and the diabetic person has a refractive index of tear cells 1.41 as explained in [5,31].
2. Design methodology
Figure 2 exhibits the structural design of the proposed multicore WGM with core diameter (${\textrm{d}_1}$) and cladding (${\textrm{d}_2})$, and Table 1 features the structural parameters. Core (1) base placed at the center coordinate(cos(0*(pi/180)))/5, (sin(0*(pi/180)))/5) and rotating core (1) 90 deg the second core was taken, and rotating core (1, 2) 180 deg core (3, 4) was taken. The size of the microsphere core represents the absorption or attachment of the molecule. The enlargement of microsphere radius size results in the increment of effective traveling distance of photons in the microsphere from the optical resonance principle. So, a longer wavelength is required to match the resonance condition under the same resonance mode [7]. So, we have used the following core diameter to get the resonance in all four cores and keep the wavelength within this range because a higher wavelength might affect the analyte refractive index and create many optical losses. A point source of diameter (a) is placed in the center with an electric field (${\textrm{E}_\textrm{z}}$ = 20V/m) have been used to excite the resonator for TM mode. The phase-matching layer (PML) width (${\textrm{d}_3}$) domain has been placed in an outer layer of the cladding to absorb undesired radiation and reflection.
The core material used here is a glass material named SUMITA-GIR (K-GIR79) [35,36], and the refractive index is denoted as (${\textrm{n}_1}$). The material dispersion properties for the applied laser wavelength ($\mathrm{\lambda }$) range (1.59 $\mathrm{\mu }\textrm{m}$ to 1.7 $\mathrm{\mu }\textrm{m) were}$ measured using Eq. (2) and have been plotted in Fig. 3.
We simulated 2D out of the plane transverse magnetic mode (TM) using the electromagnetic frequency domain and measured the output transmission of the average electric field (norm E). We have measured the total core and cladding average output electric field norm using a domain probe. We have observed the resonance wavelength shift due to biological solutions refractive index variation using frequency sweep ($f$) for a particular wavelength ($\mathrm{\lambda }$) range (1.6 $\mathrm{\mu }\textrm{m}$ to 1.7 $\mathrm{\mu }\textrm{m}$) with a 0.001 $\mathrm{\mu }\textrm{m}$ incremental wavelength change. The norm of the electric field generally represents the energy intensity in electromagnetism [7]. To investigate the influence of the refractive index of cladding or biological solutions, we have obtained the scattering spectra of the radiation energy which outflow from the resonator core surface with the varying excitation wavelength range (1.6 $\mathrm{\mu }\textrm{m}$ to 1.7 $\mathrm{\mu }\textrm{m}$) for different biological solutions refractive index optical properties [7]. When the analyte filled in the cladding region with a refractive index greater than the medium around the multiple resonators are adsorbed on the resonators, it pulls a part of the resonators’ optical field outward, increasing the optical path length and guiding to a redshift in the resonance mode. Finally, in the last section of this paper, we have explained the details of transverse electric (TE) mode analysis for the proposed sensor. Graphs are plotted using MATLAB software environment from the FEM simulated data, and we have focused on the norm of the electric field peak region. The extremely fine triangular mesh has been taken in core and cladding, and the free quadrilateral structured mesh has been taken in the PML region.
Table 1. Structural Parameters
3. Multi-core whispering gallery mode
The whispering gallery mode (WGM) sensor operates on the principle of evanescent waves coupling from the external light beam. The radiative coupling of microspheres with the light beam is described using generalized Lorentz Mie scattering theory, including the coupling of high-quality factors and controlling the coupling factors of the WGM resonator's detailed theoretical explanation given in [9]. In a multicore optical fiber, the same phenomenon occurs because the source evanescent waves couple's multiple cores. Hence, it also has multiple resonator frequencies compared to single-core fiber. Coupling multiple microspheres using evanescent waves, several techniques and applications are explained in [37].
The quality of WGM depends significantly on the contrast between core and cladding refractive indexes. Since the core refractive index is almost fixed, the cladding index varies due to bio-sensing and bio-fluid measurement, and the refractive index of the cladding (biological analyte) is in the range of (1.33 to 1.41). Therefore, as the refractive index of the analyte increments, the contrast between core and cladding refractive index decreases. It results in mode leakage, and decreasing the overall quality of WGM has been shown by comparing Fig. 4 the cladding refractive index 1 with Fig. 5 cladding refractive index 1.33.
However, it is still susceptible to refractive index change explained in the next section of this paper. The works in [7,38,39] described that due to variation (increment) in cladding or medium refractive index, output resonance wavelength shifts right side, which means the wavelength increases (redshift) for obtaining the same resonance conditions. Focusing on the resonance wavelength in Fig. 4 and Fig. 5 for similar mode, we have justified that the same phenomenon occurs, and the wavelength shifts 1.4925 $\mathrm{\mu }\textrm{m}$ to 1.627 $\mathrm{\mu }\textrm{m}$ due to cladding refractive index change from 1 (air) to 1.33 (solutions or liquid). Again, Fig. 6 (a) shows the none resonance condition for the cladding refractive index 1.33, where minimal transmission of electric field norm is seen in the core region. Besides, from Fig. 4(b) and Fig. 5(b), the electric field norm height plot shows the comparison of leakage in the core due to medium refractive index change, and Fig. 6(b) confers at the none resonance condition no cores are excited, so no spike in the core region, which results in a reduction in output-transmission spectrum at none resonance conditions.
Fig. 4. WGM in multicore optical fiber with cladding refractive index 1 and wavelength 1.4925 $(\mathrm{\mu }\textrm{m)}$ of electric field norm with respect to (a) surface plot (b) height plot.
Fig. 5. WGM in multicore optical fiber with cladding refractive index 1.330 and wavelength 1.627 $(\mathrm{\mu }\textrm{m)}$ of electric field norm with respect to (a) surface plot (b) height plot.
Fig. 6. No WGM in multicore optical fiber with cladding refractive index 1.33 and the none resonance condition in the core of electric field norm with respect to (a) surface plot (b) height plot.
4. Performance analysis (TM mode)
4.1 Wavelength sensitivity
Wavelength shifting is a significant sensor sensitivity evaluating parameters that define the ratio of resonance wavelength shift with respect to the change in refractive index. The wavelength sensitivity can be expressed as:
In Eq. (3), the resonance wavelength shift is $\mathrm{\Delta }{\mathrm{\lambda }_{\textrm{peak}}}$ between two consecutive resonant due to refractive index change, and the refractive index unit (RIU) change is $\mathrm{\Delta n}$ between two neighboring. Using electric field norms peak for respected resonance wavelength shift has been taken from Fig. 7, and the resonance wavelength sensitivity has been calculated using Eq. (3). This calculated resonance wavelength shift sensitivity for each refractive index change is shown in Table 2 respectively. Equation (4) explains the relationship between wavelength and frequency where ($f$) is the frequency and ($\lambda $) is the corresponding wavelength.Fig. 7. Norms of electric field with respect to wavelength due to refractive index variation (TM mode).
Figure 7 shows the following resonance wavelength shift due to the refractive index change 0.01 from the refractive index range (1.33 to 1.40). Using an optical spectrum analyzer (OSA), the output transmission spectrum in the terahertz range has been measured. From Table 2 and Fig. 7, the 6 nm resonance wavelength shift for each 0.01 refractive index increment has been obtained, and the sensitivity calculated using Eq. (3) is 600 nm/RIU. The resonance wavelength shift sensitivity of our proposed sensor is quite large compared to other WGM based sensors’ sensitivity.
Table 2. Sensing Parameters (refractive index 1.33-1.40)
Moreover, Fig. 8 shows a relational graph of resonance wavelength shift taking refractive index 1.33 as a base and adding concerning resonance wavelength change due to refractive index change for 1.34, 1.35, 1.36, 1.37, 1.38, 1.39, 1.40. The refractive index has been taken in a range (1.33 to 1.40) because most biological solutions refractive index changes are in this range, and our multicore WGM sensor can give an output of resonance wavelength shift for each 0.01 refractive index change. Later we have shown the detection of cancer cells and diabetes tear cells using our multicore WGM sensor.
Fig. 8. Resonance wavelength shift with respect to refractive index variation for our proposed biosensor.
4.2 Sensor resolution
Sensor resolution indicates the smallest number of changes in the refractive index, which a sensor can detect precisely. As mentioned earlier, the multicore sensor can easily detect 0.01 refractive index changes with a 6 nm wavelength shift in the output electric field norms. However, if the output optical measurement instrument can detect a 0.9 nm wavelength shift, then the sensor can give resonance wavelength shift for a 0.001 refractive index change shown in Fig. 9. From Fig. 9 graph, the measured norm of electric field peak for refractive index changes from 1.330 to 1.331, and 0.9 nm resonance wavelength of redshift occurs. So, the resonance wavelength sensitivity becomes 900 nm/RIU. We have obtained this result by varying the source wavelength range from 1.61 $\mathrm{\mu }\textrm{m}$ to 1.645 $\mathrm{\mu }\textrm{m}$ with wavelength increment taken 0.0001 $\mathrm{\mu }\textrm{m}$ and by detecting the change in wavelength shift. So, the resolution of this sensor can be enormously higher depending on the detector device at the output.
Fig. 9. The resonance wavelength shift respected to the average norm with respect to refractive index change.
Table 3 shows some previously proposed WGM based sensors with previously proposed WGM based. Thus, from Table 3, we can see that our proposed sensor exhibits better sensitivity than these previously proposed WGM based sensors. Its sensitivity may not be as high as SPR based photonic crystal sensors. However, considering the low complexity in fabrication and simplicity, our multicore WGM based sensor may be used for a wide range of applications.
Table 3. Sensitivity comparison with existing published models
5. Detection of cancer and diabetes cells
The bio-components are basically biological solutions that we have to introduce into the cladding as the cladding is unfilled. The physical size of the cladding layer is not altered just by pumping the analyte solutions into the cladding details mentioned earlier section in (Fig. 1) and [20,21]. As the resonator's cores are highly sensitive to medium or resonator core surface optical property or refractive indexes, so for different analytes filled in the cladding, it gives resonance wavelength shifts. Due to the increment in the biological component refractive index, the resonance wavelength shifts right, which is also called redshift, indicates the increment of resonance wavelength due to the refractive index increment. Here as the refractive index of biological solutions incremented from 1.33 to 1.40, as a result, the resonance wavelength shifts right for each biological solution refractive index change shown in Fig. 7 and Fig. 8. Again, for healthy basal cells, the analyte solution refractive index is 1.36, which has to be pumped into the cladding and then using the tunable laser source varying the wavelength and observing the output of scattering field intensity, the resonance wavelength 1.645 $\mathrm{\mu }\textrm{m}$ is measured for healthy cells.
Similarly, basal cancer cells have a refractive index of 1.38, so when the analyte of the basal cancer cell is pumped into the cladding, the refractive index of the cladding changes accordingly. So, by using a tunable laser source, varying the input wavelength, and observing the output of the scattering field intensity, we get a resonance wavelength of 1.658 $\mathrm{\mu }\textrm{m}$. Thus, this resonance wavelength shifts 13 nm due to cancer affected from healthy basal cells. Therefore, following this process, we detect the cancer present in the following cells or the analyte solutions. Similarly, for HeLa and MDA-MB-231 cells and diabetes tear cells, we see variations in the refractive index as well, so they are also detected using a similar process using our multicore sensor.
As stated earlier, cancer-affected cells and diabetes tear cells could easily be separated by observing the resonance wavelength change for the refractive index variation between healthy and affected cells. Figure 10 and Table 4 show the detection of Basal cancerous cells, HeLa cancerous cells, and MDB-MB-231 cancerous cells by observing the wavelength shifts due to the refractive index difference between healthy and affected cells. Here we can see that for Basal healthy cells and cancerous cells refractive index difference is 0.02 for corresponding healthy and affected cells. As the refractive index of the solution varies from the reference healthy cells solution, the resonance wavelength shifts right 13 nm, which results in wavelength shift sensitivity calculated from Eq. (3) is 650 nm/RIU. Similarly, for HeLa and MDB-MB-231 cells, resonance wavelength shifts 16 nm, and 9 nm. Again, the wavelength sensitivity calculated from Eq. (3) is 666.667 nm/RIU and 642.857 nm/RIU, respectively, for the refractive index variation of 0.024 and 0.014 healthy and affected HeLa and MDB-MB-231 cells. So, using this resonance wavelength shift due to the healthy and affected cell's refractive index variation, cancer cells can be detected using our proposed multicore WGM sensor.
Fig. 10. Basal, Hela, and MDB-MB-231 healthy and cancerous cells refractive index with respect to the norm of electric field (V/m) respected to wavelength ($\mathrm{\mu }\textrm{m}$).
Table 4. Cancer cell detection
Again, for diabetes measurement using the tear cells refractive index difference of 0.06 between healthy and affected tear cells. In Fig. 11 and Table 5, we can see the resonance wavelength redshift for the affected cells, almost 39 nm, from the healthy cells. Again, the sensitivity is calculated from Eq. (3) is 650 nm/RIU. So, by observing the output resonance wavelength shift due to the variation between the healthy and affected cells, our proposed sensor can easily distinguish cancer cells and affected diabetes tear cells by using the optical properties of refractive index difference corresponding to healthy and affected cells.
Fig. 11. Healthy diabetes tear cells and affected tear cells refractive index with respect to the norm of electric field (V/m) with respect to wavelength ($\mathrm{\mu }\textrm{m}$).
Table 5. Diabetes tear cells
6. Transverse electric (TE) mode analysis
To calculate transverse electric (TE) mode, we have simulated the proposed sensor for in plane electric field component having ${\textrm{E}_\textrm{x}}$ and ${\textrm{E}_\textrm{y}}$ of the point source shown in Fig. 2. For investigation, we have calculated the average electric field (norm E) for scattering field and varied wavelength range (1.42 $\mathrm{\mu }\textrm{m}$ to 1.5 $\mathrm{\mu }\textrm{m}$) to observe the scattering resonance wavelength shift due to medium refractive index variation. Here, we focused only on the resonance conditions of electric field norm peak due to resonance wavelength shits. Similarly, as earlier mentioned, process biological solutions having a refractive index range (1.33 to 1.40) are pumped into the cladding.
In TE mode, resonance wavelength shift is calculated. From Fig. 12, we see for biological components 1.33, 1.34, 1.35, 1.36, 1.37, 1.38, 1.39, 1.40, the resonance wavelength shifts 5 nm, 5nm. 5 nm, 6 nm, 5 nm, 5 nm, 6 nm, respectively. In addition, calculating wavelength sensitivity from Eq. (3), we get 500 nm/RIU, 500 nm/RIU, 500 nm/RIU, 600 nm/RIU, 500 nm/RIU, 500 nm/RIU, 600 nm/RIU for each 0.01 refractive index change range from 1.33 to 1.40. In Fig. 13, we have plotted the wavelength change concerning the refractive index for the proposed sensor for TE and TM mode. In Fig. 13, we have noticed that the resonance wavelength shift is higher in TM mode than TE mode and resultant wavelength sensitivity is also higher in TM mode.
Fig. 12. Norm of electric field respected to wavelength due to refractive index variation (TE mode).
Fig. 13. Comparison between TM and TE mode resonance wavelength shift with respect to refractive index increment.
It is because the electric field direction is drawn outward by the analytes, which are adsorbed on the resonators, and the electric field distribution inside the resonators are mutually perpendicular to each other in a case when there is a 90° shift. A similar comparison from Fig. 13 of TM and TE mode wavelength comparisons can also be found in [44].
7. Conclusion
In this paper, a novel WGM based multicore optical fiber sensor structure has been proposed and determined the resonance wavelength shift sensitivity for a refractive index range (1.33 to 1.41) with sensitivity around 600 nm/RIU for 0.01 refractive index change. However, the resolution of our sensor refractive index change can be 0.001 with a resonance wavelength shift sensitivity of 900 nm/RIU for 0.9 nm resonance wavelength shift. In addition, our multicore optical fiber WGM sensor can detect cancer cells and affect diabetes tear cells accurately with wavelength shift sensitivity around 650 nm/RIU. As our proposed model is a simple and less complex structure, it may become of great use for numerous bio-sensing and medical applications.
Disclosures
The authors declare no conflicts of interest
Data availability
No data was generated or analyzed in the presented research.
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