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The concentration and pH value of the immobilized glucose oxidase (GOD) are critical parameters in a glucose sensor. In this study, we develop a glucose fiber sensor integrated with heterodyne interferometry to measure the phase difference arising from the chemical reaction between glucose and GOD. Our studies show that the response time and resolution of this sensor will be strongly affected by the pH and concentration properties of GOD. In addition, the results show good linearity between the calibration curve for the glucose solution and serum based sample. The response time of this sensor can be shorter than 3 seconds and best resolutions are 0.1 and 0.136 mg/dl for glucose solution and serum based sample, respectively.
©2010 Optical Society of America
1. Introduction
Fiber biosensors have become an important sensing technology. The properties of optical fibers such as excellent light delivery, less environmental interference, easy enhancement of the bio-reaction, low cost, and ability to capture of both excited and emitted light from the measured target with a single component, means that they are being widely used in many application areas. This has encouraged the proposal of various kinds of optical fiber sensors for glucose or chemical concentration measurement. Fiber biosensors can be divided into two categories depending on the measurement method, namely intensity related [1–4] or phase related [5,6]. Jiang et al. [1] used the lock-in technology to detect the fluorescent signal coming from the fluorescence indicator. The detection range and response time of their sensor were 50 to 500 mg/dl and 30 sec, respectively. Ganesh et al. [2] developed a fiber oxygen and glucose sensor which measured the fluorescence quenching activity. They obtained a 1.7 sec response time and a measurement range between 18 to 180 mg/dl in the glucose concentration. Rosenzweig et al. [3] developed a glucose fiber sensor analyzed the fluorescent signal by using a photomultiplier and photon counter installed in a fluorescent microscope. Their method obtained a rapid response time of about 1.5 sec and a measurement range between 18 to 180 mg/dl with high linearity. Portaccio et al. [4] designed a glucose biosensor which immobilized glucose oxidase entrapped by silica sol-gel matrix onto the optical fiber bundle. They analyzed the fluorescence spectra in the UV and visible regions of different glucose concentrations, obtaining a linear calibration curve from 3.6 to 180 mg/dl. Most of the afore-mentioned methods involve measuring the variation in intensity of the fluorescence or transmitted light. Hence, avoiding the influence of surrounding light was important and often requiring expensive photon detection equipments. In contrast, the phase-measured technique proposed by Chiu et al. [5] can ignore the influence of light intensity and the photomultiplier is no longer required. They demonstrated a fiber sensor with D-shape and SPR property which could detect variation in the alcohol concentration of 2%. Unfortunately, the sensitivity of the SPR type sensor is strongly influenced by the quality (thickness, roughness, and uniformity) of the deposited metal. They could not measure samples with complicated composition unless the SPR sensor was modified with specific enzyme. Wang et al. [6] described a glucose fiber sensing apparatus which combined the bifurcated fiber and sensing membrane. Based on analysis of the phase-shift of the fluorescent signal coming from the oxygen consumption in the membrane, the method can obtain the glucose concentration of the testing sample with detection limit of 50 mg/dl.
In this paper, we propose a fiber type measurement system consisting of a fiber sensor and heterodyne interferometry to achieve rapid glucose concentration measurement. An easy to fabricate optical biosensor was made with immobilized GOD at the core of a commercially available single mode fiber. Because of the true phase detection, the resolution of our method can reach 0.1 mg/dl for glucose solution and 0.136 mg/dl for a serum based sample. We also evaluated the pH and concentration properties (pH 7.5 and 10 μg/ml) of GOD, confirming that the sensor could be used for serum sample measurement. Our fiber biosensor has the advantages of easy fabrication, rapid response time, acceptable resolution, high linear calibration curve in a wide detection range, and suitability for glucose solution and serum sample.
2. Principles
The fiber-type sensing system for rapid glucose concentration measurement is shown in Fig. (1a) . The heterodyne light source was coupled into a single mode fiber by a collimating lens CL1 with suitable numerical aperture (NA). The guide light passes through the sensing part which the fiber cladding was removed and immobilized glucose oxidase (GOD) was applied on the fiber core. The sensing structure is shown in the enlarged diagram inset to Fig. (1a). The high selectivity of the GOD on the sensing part can be used to detect the glucose concentration of the test sample dripped on it. As the heterodyne light source enters the sensing part, the light beam undergoes total internal reflection (TIR) as shown in the enlarged diagram in Fig. (1a). Based on the Jones calculation and Fresnel equation [7], the phase difference between the p- and s- polarization states can be written as
where , n 1 and n 2 are the refractive indices of the immobilized GOD and the testing solution. θt and m are the incident angle and the number of TIRs that occur at the interface between the GOD and the testing solution. The number m can be represented aswhere L and h are the length of the sensing part and diameter of the fiber core, respectively.
Fig. 1 Schematic diagram of the measurement system and preliminary test of the glucose fiber sensor; (a) optical configuration of the system; (b) preliminary test with POD method.
After passing through the sensing part, the light beam will be collimated by CL2 and guided through the analyzer AN1, to be detected by the photodetector D1. The test signal detected by D1 can be written as
The reference signal Ir comes from the function generator and can be written as . Theoretically, before the testing solution is dripping onto the sensor, the phase difference between Ir and It can be made equal to zero by adjusting the initial phase of Ir. After dripping the testing sample onto the sensor, the phase will vary as the glucose reacts with the GOD to be converted into gluconic acid and hydrogen peroxide. The chemical reaction can be formulated as follows:It means that the refractive index (n 2) will change and consequently the phase will change. Besides, the phase difference ϕt is a function of refractive index n 2. Furthermore, it is also a function of the concentration of the testing sample. Theoretically, the phase difference is proportional to the change of the refractive index and also to the concentration variation of testing sample [5,8,9]. Based on the Eqs. (1) and (2), we can simulate the phase variation by the change of the refractive index from 1.330 to 1.341 as shown in Fig. 2 . In practice, the phase difference ϕt can be obtained immediately upon sending the testing and reference signals into the lock-in amplifier.3. Experiment and results
To demonstrate our method, a heterodyne light source was constructed by a linearly polarized He-Ne laser at 632.8 nm with frequency stabilization control, and an electro-optic modulator EO with the fast axis located at 45° to the x-axis was driven by a function generator with 1 kHz sawtooth signal, as shown in Fig. 1(a). The fiber sensor is made by the removal of the fiber cladding by immersion in hydrofluoric acid (HF acid) for about 10 min and the fiber core was then cleaned in 3:1 mixture of sulfuric acid (H2SO4) and hydrogen peroxide (H2O2) for 10 min. After that, the surface of the fiber core surface was modified by treatment with the amino linkage aldehyde group with 1% (v/v) 3-(trimethoxysilyl)propyl aldehyde in absolute ethanol for 30 min at room temperature which was utilized to immobilize GOD. The modified fibers were immersed into three solutions with different GOD concentrations (10 μg/ml, 5 μg/ml, and 1 μg/ml) (pH 7.5) for 1 hour. Then, the unreacted aldehyde groups were quenched by 15 mM of Tris buffer (pH 7.5) for 10 min at room temperature [10]. The dimensions of the fiber sensor are h = 8 μm and L = 15 mm. The NA of the collimators CL1 and CL2 are 0.5 purchased by Thorlabs Inc. The azimuth angle of the analyzer AN1 is 45° with respect to the x-axis. The lock-in amplifier (model: SR 850, Stanford Research System) with the phase resolution of 0.001° was used to measure the phase difference. It is connected to a PC by RS-232 interface for data acquisition. To avoid the influence of the ambient temperature fluctuation, the room temperature was controlled to remain at 25 °C.
The preliminary testing followed the method [11,12] presented by the World Health Organization, (WHO). Glucose present in the testing solution will be oxidized by GOD to form gluconic acid and hydrogen peroxide. The hydrogen peroxide will be converted into water and oxygen by the peroxidase (POD). An oxygen acceptor 4-4-aminophenazone takes up the oxygen and together with phenol forms a pink coloured chromogen. Thus, when the fiber sensor is placed into the solution containing the glucose, POD, and 4-4-aminophenazone, the pink color will appear about 30 minutes later, as shown in Fig. 1(b).
The glucose solutions of seven concentrations were prepared by dissolving glucose anhydrous in a phosphate buffer solution (PBS) (10, 40, 80, 100, 200, 300, and 400 mg/dl). Their pH values were controlled at 7.5 and 9 by mixture of HCl and NaOH. The glucose concentrations were certified using GC-MS by Medical Metrology Department, Center for Measurement Standards in Taiwan. Besides, the serum based sample (SRM 965a) was purchased from National Institute of Standards and Technology (NIST) and consisted of eight flame sealed ampoules of frozen human serum, two ampoules at each of four different glucose concentration levels. The four different glucose concentration levels were 34.56, 78.5, 122.1, and 292.6 mg/dl, as certified by NIST following the standard procedure. The pH values of each level were equal to 7.5.
Blank control can be another preliminary test for the fiber sensor. The tendency for phase variation indicates the efficacy of the fiber sensor. We prepared 5 testing samples consisting of 2 glucose solutions, 2 serum based samples, and one glucose free sample (PBS solution). Small amounts were dripped them onto the fiber sensor (10 μl in volume). The results are shown in Fig. 3 . The phase variation tendencies of the sample with glucose are much greater and sharper than the sample without glucose. In addition, we can confirm that the phase variation comes from the reaction between the GOD and the glucose.
Fig. 3 Blank control of the fiber sensor using two glucose solutions, 400 and 10 mg/dl; two serum based samples SRM 965a level 1 and level 4; and one glucose free sample (PBS) for the control experiment.
The response time and slope of the calibration curve will be affected by the pH values and concentration of the GOD immobilized on fiber. In Fig. 4(a) shows the response time given at different pH values for 10 μl of glucose solution, a glucose concentration of 80 mg/dl, and GOD concentrations fixed at 10 μg/ml with pH 7.5. Each pH value was measured 10 times and the variation shown here was one standard deviation. It is clear that there was no significant difference in the response time for glucose solutions with pH 7 and pH 8. However, the response time of the sample where pH values between the GOD and glucose solution were identical was better than those for samples where the pH values were different. Variation of the response time for each level of the SRM 965a is shown in Fig. 4(b). By contrast, the response time was 0.1 sec longer for level 2 (78.5 mg/dl in glucose concentration) of SRM 965a than for the glucose solution. This difference might be coming from the interference of the serum based sample.
Fig. 4 pH dependence of the response time of testing sample; (a) glucose solution with 7 different pH values; (b) serum based sample (SRM 965a) with 4 concentration levels
Figure 5 shows the response time of two kinds of testing samples measured by the sensor controlled pH at 7.5 with three different GOD concentrations. Figure 5(a) shows the results of glucose solution measurement which the glucose concentration was 100 mg/dl with pH 7.5.
Fig. 5 GOD concentration dependence of the response time of the fiber sensor; (a) glucose solution with 100 mg/dl in a glucose concentration with pH 7.5; (b) serum based sample (SRM 965a level 1).
There was no significant difference in the response time (marked A in the figure) between the sample with GOD concentrations of 10 and 5 μg/ml, but the response time of GOD concentration 1 μg/ml was about 0.4 sec longer (marked B in the figure). Figure 5(b) shows the results of serum based sample (SRM 965a level 1). The response times were 1.2, 1.4, and 1.7 sec, marked A, B, and C, respectively. It is clear that for the serum based sample, the response time is strongly related to the GOD concentration of the fiber sensor. Based on the findings shown in Figs. 4 and 5, we can conclude that both the pH and concentration properties of the GOD will affect the response time of the fiber sensor. As the pH values of the sample being tested and the GOD are similar, the response time of the fiber sensor will be faster. Meanwhile, a higher concentration of GOD will lead to the faster response time for both glucose solution and serum based sample.
Finally, we show the calibration curves for the glucose solution samples and the serum based sample. The measurement conditions include controlling the pH values of GOD at 7.5 and fixing the concentration at 10 and 1 μg/ml. In Fig. 6(a) , we have the calibration curves for the glucose solution samples with pH 7.5 and 9, concentration in the range of 10 ~400 mg/dl, using the fiber sensor of 10 μg/ml in GOD concentration. The slopes of the calibration curves between pH 7.5 and 9 have a small difference. Those are similarly exhibited in Fig. 6(b), where the sensor with 1 μg/ml in GOD concentration is used. The slopes of calibration curves of the same pH value of glucose solution measured by the sensors with different GOD concentrations have no significant difference, as shown in Figs. 6(a) and 6(b). The results of the serum based sample are similar, as shown in Fig. 6(c), where the glucose concentration is in the range of 34 ~290 mg/dl and the sensors with two different GOD concentrations are used. These findings indicate that the slope of calibration curve was not related to the GOD concentration of fiber sensor but highly related to the pH values of the tested sample and GOD.
Fig. 6 The calibration curves of the glucose fiber sensors: (a) glucose solution samples with GOD two pH values, 7.5 and 9 with sensor’s GOD concentration at 10 μg/ml; (b) glucose solution samples with GOD two pH values, 7.5 and 9 with sensor’s GOD at 1 μg/ml; (c) serum based samples (SRM 965a) with two different sensor (GOD concentrations at 10 and 1 μg/ml).
4. Discussion
The resolution Δc of our method can be obtained by calculating the ratio of the phase error |Δϕ| of our method to the slope s of the calibration curve which is represented by
The phase error |Δϕ| in our method may come from angular resolution of the lock-in amplifier, second harmonic error, polarization-mixing error, and incident angle error due to the misalignment between collimator and the fiber. Based on the analysis method discussed in previous work [13] the phase error in our method can be calculated as equaled to 0.02°.According to the results of Fig. 6, the resolution of our method can be determined and summarized in Table 1 . It is clear that the resolution will be affected by the pH properties of the testing sample rather than the concentration of the GOD. However, the response time will be affected by the pH of the testing sample, and the pH and concentration of GOD. Hence, the pH property of the sensor is critical for rapid measurement. In this case, we find that the rapid response time can be achieved by controlling the pH and concentration of GOD at 7.5 and 10 μg/ml, respectively, during measurement of both glucose solution and serum based sample.
Table 1. Resolution of glucose fiber sensor of two types of testing sample
5. Conclusion
In this paper, we fabricated the glucose fiber sensor with immobilized GOD on the core surface, integrated with the heterodyne interferometry. Given the selectivity of GOD and true phase measurement, the best resolutions were 0.1 and 0.136 mg/dl for glucose solution and serum based sample, respectively. Furthermore, the response times were shorter than 1 sec as the pH values of testing sample and GOD were identical. Our results show that the pH property of the GOD strongly affected the response time and measurement resolution of the sensor. Besides, the concentration of GOD influenced the response time of the fiber sensor only, it did not significantly influence the measurement resolution. Therefore, we can conclude that the suitable pH value and concentration of GOD will allow us to achieve rapid measurement of both glucose solution and serum based sample. In this study, the concentration and pH of GOD were 10 μg/ml and 7.5 for achieving rapid measurement of the glucose solution samples and serum based samples.
Acknowledgement
The authors would like to thank the National Science Council of the Republic of China for financially supporting this research under Contract No NSC 98-2221-E-155-001. We would like to thank Prof. Der-Chin Su for support with essential experiment equipments and useful discussion related to this work. We also would like to acknowledge of the contributions of Mr. Daniel Irwin King, Mrs. Debbie Nester, and Prof. Nien-Po Chen in proofreading this article.
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