Surface plasmon resonance based fiber optic glucose ...

Surface plasmon resonance based fiber optic glucose biosensor Sachin K. Srivastava, Roli Verma and Banshi D. Gupta* Indian Institute of Technology Delhi, New Delhi, India - 110016 ABSTRACT A surface plasmon resonance (SPR) based fiber optic biosensor has been fabricated and characterized for the detection of blood glucose. Optical fiber sensor was fabricated by first coating a 50 nm thick gold film on the bare core of optical fiber and then immobilizing glucose oxidase (GOx) over it. Aqueous glucose solutions of different concentrations were prepared. To mimic the blood glucose levels, the concentration of glucose solutions were kept equal to that in human blood. The refractive indices of these sample solutions were equal to that of water up to third decimal place. SPR spectra for the sensor were recorded for these glucose solutions. When the glucose comes in contact to glucose oxidase, chemical reactions take place and as a result, the refractive index of the immobilized GOx film changes, giving rise to a shift in the resonance wavelength. Unlike electrochemical sensors, the present sensor is based on optics and can be miniaturized because of optical fiber. The present study provides a different approach for blood glucose sensing and may be commercialized after optimization of certain parameters. Keywords: Surface plasmon resonance, biosensor, glucose, optical fiber, glucose oxidase

1. INTRODUCTION For past two decades, surface plasmons have gained wide attention due to their potential applications in the fields of chemical and biochemical sensing1, 2, wave guiding3, imaging4, metamaterials5 etc. Surface plasmons are basically charge density oscillations at a metal dielectric interface. Due to these oscillations, a wave, called surface plasma wave, gets generated at the interface and travels along it. The field associated to these waves decays exponentially in both the media. After a propagation of certain length, called propagation length, these waves die out due to non-radiative decay and the energy is dissipated in the form of heat. These waves are transverse magnetically (TM) polarized. When the wave vector of a beam of TM polarized light incident on the metal- dielectric interface matches with that of surface plasmons, they get excited. This is called surface plasmon resonance (SPR). In general, the wave vector of the light wave travelling in the dielectric medium is always greater than that of surface plasmons supported by the interface of that dielectric and a metal. Therefore the excitation of surface plasmons in this way is not possible due to mismatch in the wave vector. Special techniques, such as Kretschmann and Otto configuration are employed for surface plasmon excitation6. In a basic Kretschmann configuration scheme, the base of a prism of high refractive index is coated with a thin metal layer, which is further surrounded by the dielectric to be studied. A beam of TM polarized light incident on one side of the prism traverses it up to the base; gets total internally reflected at the prism-metal interface and comes back from the other face. A part of the incident power gets lost due to absorptions by metal because of its imaginary part of dielectric constant. This is termed as attenuated total internal reflection (ATR). At a particular value of the angle of incidence, the wave vector of the evanescent wave at prism-metal interface becomes equal to that of surface plasmon wave at metal dielectric interface. This corresponds to the excitation of surface plasmons and as a result, total incident power gets transferred to the surface plasmons giving rise to a sharp dip in the reflected power. This angle, called as resonance angle, is the characteristic of the metal-dielectric interface. This resonance angle is highly dependent on the refractive index of the dielectric medium and/or the metal film. After the metal film is coated on the base of the prism; the resonance angle changes with the refractive index of the dielectric medium around the metal. This is how surface plasmons are utilized in sensing. The configuration can also be used in spectral interrogation mode, in which light from a polychromatic source is used at a fixed angle of incidence. The wave vector of some particular wavelength matches with that of surface plasmons and a dip corresponding to resonance wavelength is observed in the reflected light. This resonance wavelength is also highly dependent on the refractive index of the surrounding medium. * [email protected]; phone+91-11-2659-1355; fax +91-11-2658-1114; iitd.ac.in

Third Asia Pacific Optical Sensors Conference, edited by John Canning, Gangding Peng, Proc. of SPIE Vol. 8351, 83511Z · © 2012 SPIE · CCC code: 0277-786X/12/$18 · doi: 10.1117/12.915978

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The prism based devices being quite bulky in size and having a number of mechanical moving parts are less favored. Optical fibers can be used in place of prism. They are advantageous in terms of miniaturization, cost, remote sensing, online measurements etc. In a fiber optic SPR sensor, the prism is replaced by the core. A small portion of the cladding of the optical fiber is removed and the metal film is coated on the bare core. This is further surrounded by the sensing medium. Light incident at the input face of the optical fiber gets guided along its length due to total internal reflections. The evanescent waves at the core-metal interface excite surface plasmons in the sensing region. Light transmitted off the output face is fed to a spectrometer, which is interfaced with a computer. Many studies utilizing fiber optic SPR based chemical and biochemical sensors have been reported in literature1, 2, 7. Several attempts have also been made to enhance the sensitivity of these sensors8. According to the American Diabetes Association, a post meal blood glucose level must be less than 180 mg/dl, while the recommended pre-meal glucose level should be between 90 to 130 mg/dl9. If the glucose level in human blood does not lie between the above ranges, the person is considered to be suffering from sugar diseases. Increase of blood sugar from reference values, termed as hyperglycemia, may damage organs. Diabetes mellitus is a major disease caused by increased glucose level in the blood. It is spread all over the globe and is a major health problem of developed societies. According to the reports of the World Health Organization (WHO), the number of diabetic patients throughout the world was approximately 171million in 2000 and this is expected to increase to 366 million by 203010. Decrease in blood glucose level is termed as hypoglycemia. It may cause many serious diseases due to inadequate supply of glucose to the brain. This may result in neuroglycopenia, mild dysphoria, unconsciousness or permanent brain damage. Therefore, the detection and monitoring of glucose level in blood at regular intervals is essential. Many studies have been performed for the detection of glucose11-13. Most of these use electrochemical methods12, 13. Absorption based optical fiber glucose sensors have also been reported in literature11. In the present study, the fabrication and characterization of a surface plasmon resonance based fiber optic biosensor for detection of blood glucose has been reported. The sensor was fabricated by coating the unclad portion of optical fiber with a thin film of gold and then immobilizing glucose oxidase over it. To characterize the sensor, aqueous glucose solutions were prepared. The concentrations of these samples were kept equal to that of glucose in human blood. SPR spectra for these solutions were recorded and the results were interpreted. Most of the commercially available glucose sensors are electrochemical in nature. The present study examines a new approach to fabricate optical glucose sensors. If all the parameters are carefully optimized, the present sensor may be advantageous in terms of sensitivity, miniaturization and cost.

2. MATERAILS AND METHODS 2.1 Materials Plastic clad silica (PCS) optical fiber of 600 µm core diameter and 0.37 numerical aperture was purchased from Fiberguide Industries (USA). Sodium dihydrogen phosphate dihydrate, di-sodium hydrogen phosphate dihydrate, and ethylenediol were obtained from Merck. Cystamine dihydrochloride was purchased from Acros organics. Sodium meta periodate and glucose oxidase (GOx) (from Aspergillus niger) were obtained from CDH lab reagents. These chemicals were used in the same form as received, without any further purification. 2.2 Fabrication of the sensor An optical fiber of 600 µm core diameter and 0.37 numerical aperture was taken and 1 cm length of its cladding was removed from the middle. A thin film of 50 nm thickness of gold (Au) was coated on the unclad portion of the optical fiber. This thickness of metal layer was coated because it has been reported as optimized thickness for the best performance. The coating procedure has already been described in our previous studies7, 8. A film of 1.6 nm thickness of chromium (Cr) was coated on the optical fiber core before the gold coating. Chromium acts as an adhesive between fiber core and gold layer. A vacuum coating unit with thermal evaporation scheme was used for the coatings. Before fixing in the coating unit, the optical fibers were cleaned with acetone to remove any impurities or dust etc. from the surface of the optical fiber core. High tension arc plasma in the vacuum chamber was used to further remove any dust from the core surface. The evaporation of metals in the chamber was performed by a low tension current. The thickness of the deposited film was continuously monitored by a quartz crystal thickness monitor fixed in the vacuum chamber. The vacuum in the chamber was kept at 5X10-6 mBar.


After the fiber was coated with gold film, glucose oxidase was immobilized over it. Before immobilization on gold surface, the peripheral carbohydrate groups of glucose oxidase were modified to aldehydes. The reaction was performed as follows. A 20 μM glucose oxidase solution was prepared in 0.1 M 5 ml sodium phosphate buffer and maintained at pH 6.8. 30 mg of sodium metaperiodate was added to this solution, and then allowed to react in dark for 1 hour with continuous stirring at 0 oC in an ultra cold immersion cooler. The reaction was stopped by adding 6.97 μl ethylenediol. The solution was kept at room temperature for 30 minutes with continuous stirring. To functionalize the gold coated optical fibers with glucose oxidase, the sensing probes were first incubated in 1 mM aqueous solution of cystamine dihydrochloride for 1 hour. This provides -NH2 groups over the Au surface. The modified GOx gets attached to these -NH2 groups. For the binding of GOx on Au surface, the Au coated optical fiber probes were incubated in GOx solution for 12 hours. The fiber probes were then washed with Millipore® water to remove any unbound species, dried at room temperature and again incubated in the GOx solution for next 12 hours. The process was repeated for effective binding of GOx over the surface of Au surface. Finally, the sensor were removed from the GOx solution, washed with Millipore® water and stored at 4 oC before their characterization. A schematic of complete layer by layer self assembled monolayer (SAM) formation on Au coated optical fiber is shown in figure 1.

Figure 1. Schematic of layer by layer self assembled monolayer formation

2.3 Preparation of sample solutions Aqueous solutions of varying concentrations of glucose were prepared in Millipore® water. The concentrations of these sample solutions were kept in range of 0-250 mg/dl to mimic the blood glucose concentration. The refractive indices of the samples were measured in white light by an Abbe’s refractometer having resolution of 0.001. The refractive indices of these solutions were found to be the same, equal to that of water, up to third decimal place. 2.4 Experimental Setup A schematic of the experimental setup used for the present study is shown in figure 2. The sensor was fixed in a glass flow cell with inlet and outlet for sample solutions to be kept around the probe. Polychromatic light from a tungsten halogen lamp (AvaLight-HAL) was focused on the input end of the optical fiber probe with the help of a microscope objective and a three dimensional (3D) translational stage. The evanescent field of the light incident on the input end of


fiber excites surface plasmons in the sensing region. The light coming out of the other end of the fiber was then fed to a spectrometer (AvaSpec-3648), interfaced with a computer. The sample solutions of glucose were filled in the flow cell and the corresponding SPR spectra were recorded.

Figure 2. A schematic of the experimental setup

Figure 3. SPR spectra for varying glucose concentrations


3. RESULTS AND DISCUSSIONS Figure 3 shows the experimental SPR spectra of the fabricated sensor for varying concentrations of glucose in water. It is observed that for each sample of glucose solution, the SPR spectrum shows a dip in the transmittance. The wavelength corresponding to the dip is termed as resonance wavelength. This resonance wavelength is a characteristic of the sample solution kept around the sensing region and is bound to change with change in the refractive index of the solution. We have already mentioned that the refractive indices of the sample solutions are the same up to the third decimal place. This means, that resonance wavelength should remain the same for all the samples. But, as observed from the figure, it changes. The reason can be given as follows. When glucose comes in contact with the glucose oxidase, the following reaction occurs in the vicinity of the sensing region. Glucose Oxidase Glucose + O 2 ⎯⎯⎯⎯⎯→ GluconicAcid + H 2 O 2 Due to the above reaction, the local refractive index slightly changes, giving rise to a change in the SPR resonance wavelength. As the concentration of the glucose sample around the sensor increases, the SPR resonance wavelength shows a red shift. This can be seen more clearly in figure 4, where we have plotted the variation of resonance wavelength with change in glucose concentration. The symbols are experimental data points obtained from the SPR spectra and the line passing through them is the best polynomial fit. It is observed that the resonance wavelength increases with an increase in the concentration of glucose solution. The sensitivity of the sensor is defined as the change in resonance wavelength per unit change in glucose concentration. The response curve can be approximated as linear, and hence the sensitivity of the present sensor, as calculated from the graph is 0.0366 nm/(mg/dl). The accuracy and resolution of the sensor depend on the spectral resolution and quality of the spectrometer.

Figure 4. Variation of SPR resonance wavelength with glucose concentration

The experiments were repeated with three different sensing probes, thrice with each one, over a period of twenty days. Fairly good repeatability and reproducibility of results was obtained. The sensor can be reused after washing it with distilled water. The selectivity and specificity of glucose oxidase to glucose have already been well studied and established14. It has also been reported that hydroquinone, bilirubin, uric acid and aspartic acid act as inhibitors and for the glucose- glucose oxidase reaction15. Therefore, the performance of the sensor may get affected by these compounds if they are present in the sample. However, these compounds are mainly the constituents of urine and not of blood. Therefore, the present sensor cannot be used for detection of glucose in urine samples. The present scheme of fabrication of the sensor seems better than that utilizes gel entrapment method2, in a way that all the active sites of the immobilized enzyme are open to the analyte.


4. CONCLUSIONS We have fabricated and characterized a surface plasmon resonance based fiber optic biosensor for the detection of blood glucose. The sensor shows a linear response in the 0 to 250 mg/dl concentration of glucose. The sensitivity of the sensor is found to be 0.0366 nm/(mg/dl). Various steps involved in the fabrication and characterization have been present in detail and response curves and advantageous features of the sensor have been discussed. The current study presents a new approach for blood glucose sensing and may be advantageous in terms of increased sensitivity, miniaturization and cost.

ACKNOWLEDGEMENTS Sachin K. Srivastava and Roli Verma are thankful to the Council of Scientific and Industrial Research (CSIR), India for research fellowships. This work was partially supported by the Department of Science and Technology (DST), India.

REFERENCES [1 ] [2 ] [3 ] [4 ] [5 ] [6 ] [7 ] [8 ] [9 ] [10 ] [11 ] [12 ] [13 ] [14 ] [15 ]

Jorgenson, R. C. and Yee, S. S., "A fiber-optic chemical sensor based on surface plasmon resonance," Sens. Act. B 12, 213-220 (1993). Rajan, Chand, S., and Gupta, B. D., "Fabrication and characterization of a surface plasmon resonance based fiber-optic sensor for bittering component--Naringin," Sens. Act. B 115, 344-348 (2006). Lee, B., Lee, I.-M., Kim, S., Oh, D.-H. and Hesselink, L., “Review on subwavelength confinement of light with plasmonics,” J. Mod. Opt. 57, 1479-1497 (2010). Berguiga, L., Zhang, S., Argoul, F. and Elezgaray, J., “High-resolution surface-plasmon imaging in air and in water: V(z) curve and operating conditions,” Opt. Lett. 32, 509-511 (2007). Kabashin, A. V., Evans, P., Pastkovsky, S., Hendren, W., Wurtz, G. A., Atkinson, R., Pollard, R., Podolskiy, V. A. and Zayats, A. V., “Plasmonic nanorod metamaterials for biosensing,” Nat. Mater. 8, 867-871 (2009). Kretschmann, E., "The angular dependence and the polarisation of light emitted by surface plasmons on metals due to roughness," Opt. Comm. 5, 331–336 (1972). Srivastava, S. K., Verma, R. and Gupta, B. D., “Surface plasmon resonance based fiber optic sensor for the detection of low water content in ethanol,” Sens. Act. B 153, 194-198 (2011). Srivastava, S. K. and Gupta, B. D., “A multitapered fiber-optic SPR sensor with enhanced sensitivity,” Photon. Technol. Lett. 23, 923-925 (2011). American Diabetes Association, “Standards of medical care-table 6 and table 7, correlation between A1C level and mean plasma glucose levels on multiple testing over 2–3 months,” Diabetes Care 29, 51–580 (2006). Wild, S.; Roglic, G.; Green, A.; Sicree, R.; King, H., “Global prevalence of diabetes: estimates for the year 2000 and projections for 2030,” Diabetes Care 27, 1047-1053 (2004). Abdel-latif, M. S., Suleiman, A. A. and Guilbault, G. G., “Enzyme-based fiber optic sensor for glucose determination,” Anal. Lett. 21, 943-951 (1988). Zhang, S., Wang, N., Yu, H., Niu, Y. and Sun, C., “Covalent attachment of glucose oxidase to an Au electrode modified with gold nanoparticles for use as glucose biosensor,” Bioelectrochemistry. 67, 15-22 (2005). Hsieh, H. V., Pfeiffer, Z. A., Amiss, T. J., Sherman, D. B. and Pitner, J. B., “Direct detection of glucose by surface plasmon resonance with bacterial glucose/galactose-binding protein,” Biosens. Bioelectron. 19, 653-660 (2004). Yoo, E.-H. and Lee, S.-Y., “Glucose biosensors: an overview of use in clinical practice,” Sensors 10, 45584576 (2010). Naganna, B., Rajamma, M. and Rao, K. V., “On the failure of enzyme paper strips to detect glucose in certain abnormal urines,” Clin. Chim. Act. 17, 219-221 (1967).
