The accuracy of the surface plasmon resonance (SPR) optical fiber sensor is affected by the change of ambient temperature. Therefore, we propose a simple dual channel SPR optical fiber sensor, which can measure both glucose concentration and ambient temperature. The proposed sensor is a two-channel structure based on a no-core optical fiber (NCF): one channel is coated with gold film and polydimethylsiloxane (PDMS) to sense the ambient temperature, and the other is coated with silver film to sense glucose concentration. The experimental results show that the sensor’s sensitivity for sensing glucose concentration is 2.882 nm / %, and for sensing temperature is -2.904 nm / °C. By monitoring the real-time temperature, the accuracy of glucose concentration detection was improved. The proposed sensor has a simple and compact structure, and it is suitable for sensing glucose solution or other analyte solutions that need temperature compensation.
© 2021 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
Optical fiber sensor based on surface plasmon resonance (SPR) has many advantages, such as miniaturization of probe, multi-channel sensing and no electrical contact between sensor and sample, and they are low in manufacturing cost, short in response time and can provide real-time testing [1–4]. Because of these excellent properties and the ability to adapt to the needs of different environments, optical fiber SPR sensors have been widely investigated and applied in many fields [5–10]. For optical fiber SPR sensors, the evanescent wave at the interface of the fiber and the dielectric resonates with surface plasmon wave (SPW), which would weaken the optical power received by the spectrometer, inducing a dip in the spectrum . The wavelength of the dip is called resonance wavelength, which depends on the RI change of the sensing medium. Optical fiber SPR sensors usually adopt the wavelength interrogation method to sense the change of surrounding medium, and because of their high sensitivity to environmental RI change, various parameters can be detected, for instance, temperature and humidity, magnetic field, pH value, liquid or gas concentration, etc [12–17].
The measurement of glucose concentration is of great significance and practical value in the fields of chemistry, physics, and especially biosensing, and a variety of sensors have been reported for this purpose [18–23]. However, the change of ambient temperature will affect the glucose detection [24,25]. In order to ensure the accuracy and reliability of the detection results, it is necessary to calibrate the ambient temperature at the measurement of glucose concentration. However, the addition of an external temperature sensor not only increases the complexity of the sensing system, for instance, it usually raises the demand for an additional light source and spectrometer, but also causes measurement errors due to the physical isolation between the temperature sensor and the glucose concentration sensor . How to better integrate with the temperature sensing is the problem that the development of glucose concentration sensor is facing.
In this paper, we propose a compact and highly sensitive dual channel no-core optical fiber (NCF) sensor based on SPR, which can measure both glucose concentration and ambient temperature. By monitoring the real-time change of ambient temperature, the accuracy of glucose concentration detection was improved and a calibrated sensing formula is obtained. Compared with the previously reported sensors, this proposed sensor exhibits a great improvement in sensitivity, accuracy and reliability at the measurement of glucose concentration.
2. Fabrication and principle
The fabrication method of the sensor proposed here is to insert a section of NCF between two sections of multimode optical fibers (MMF), as shown in Fig. 1(a). The MMF here is a commercial fiber with a core diameter of 62.5 µm and an outer diameter of 125 µm, and NCF is a self-made quartz fiber of the same outer diameter and with a length of 10 mm. A fusion splicer (Japan, Fitel, S179) was used to splice NCF between the MMFs as the sensing area.
The surface of NCF was carefully wiped with a piece of lens paper moistened with anhydrous ethanol, then dried with a nitrogen gun and placed in the evaporation chamber of magnetron sputtering. Then, the area outside NCF was covered with a mask, and the coating was carried out by DC magnetron sputtering equipment (HF-Kejing, VTC-16-SM), as shown in Fig. 1(b). The sputtering is completed in two steps (see Fig. 1(b)–1(f)). First, the gold film was sputtered on the surface of NCF, and it was necessary to flip NCF 180 ° and sputter again a gold film of the same thickness. Here, the magnetron sputtering current was controlled at 20 mA and the sputtering time was set to be 30 s. Then, the NCF with gold coating was partially embedded in polydimethylsiloxane (PDMS) which is a kind of hydrophobic organic polymer with high thermo-optic coefficient (∼ - 4.5 × 10−4 RIU / °C) (see Fig. 1(c)) . In the process of sensor production, PDMS coating was prepared by liquid casting and curing. A mold with a hole was made, and the diameter of the hole was similar to that of NCF. After the NCF with gold film was inserted into the mold through the small hole, PDMS was poured into the mold. The shape of PDMS was mainly determined by the mold, and the length of NCF embedded in PDMS can be adjusted by controlling the length of NCF embedded in the mold. The PDMS coating length was 5 mm.
In this paper, PDMS was prepared by mixing liquid silicone oil (Dow Corning SYLGARD, 184A) and curing agent (Dow Corning SYLGARD, 184B) with a configuration ratio of 10:1. The mixed solution stood for 5 h to make bubbles float to the surface and break. The sensor coated with PDMS was put into the test chamber for heating, and the heating time and temperature were set to be 20 minutes and 110 °C, respectively. After that, the sensor was cooled to room temperature. Finally, the gold coating on NCF surface which was not covered by PDMS was wiped off by the lens paper soaked in anhydrous ethanol, and silver coating was sputtered in this area with the same magnetron sputtering current and time (see Fig. 1(d)–1(e)). So far, the NCF sensor has been manufactured, as illustrated in Fig. 1(f). The sensor consists of two channels: glucose concentration measurement area which refers to the NCF section coated with silver (Channel A), and temperature measurement area which refers to the NCF section coated with gold and further embedded in PDMS (Channel B).
After a long-term experimental exploration, it is found that the coating time of 30 s can ensure a smoothness of the metal film. Moreover, the sensor with such a metal film exhibits a smooth spectral curve with low noise. Due to the limitation of magnetron sputtering equipment, the thickness of metal film cannot be accurately obtained in the experiment. However, it can be approximately evaluated using probe surface profilometer by measuring the thickness difference between uniform film and non-film area. The thickness of silver and gold film was about 43 nm and 57 nm, respectively.
During the light transmission, light will leak from the MMF into NCF, inducing a strong evanescent field at the fiber surface. If a thin metal layer is deposited on the NCF surface, SPR will be excited and can be used to sense the surrounding medium. TM mode can excite SPR along the direction perpendicular to the interface, but TE mode cannot . The coupling condition of SPR excitation is:29]. Then the coupling conditions of ns and λ can be distinguished to obtain the sensitivity, as shown by the following equation:
Before the experiment, we use the finite element method to analyze the two channels of the sensor. The thickness of the two metal coatings is set to be 50 nm and the size of NCF is the same. Figure 2 shows respectively the simulated transmission spectra with the RI change and temperature variation. We can see that the increase of RI causes the dip of Channel A to move to the long wavelength, and the increase of temperature causes the dip of Channel B to move to the short wavelength. The results show that the former is sensitive to the change of RI and the latter is sensitive to the change of ambient temperature. This dual channel sensing ability of the proposed sensor enables the simultaneous measurement of the glucose concentration and the ambient temperature. In addition, because the thermo-optic coefficient of silica is much smaller than that of water, although Channel A doesn’t respond to temperature, a temperature variation will still induce a change in the liquid RI, which would in turn cause a resonance wavelength shift of Channel A. When Channel A is not placed in the liquid, SPR will not be excited and there will be no dip in the spectrum.
3. Experimental results and discussion
3.1 Device characterization
The experimental setup is shown in Fig. 3. The light from a halogen light source (Ocean Insight, HP-2000LL, 400-2500 nm) was coupled to the prepared sensor, and the output light was detected by a spectrum analyzer (Ocean Insight, USB2000+VIS-NIR, 350-1000 nm). In the experiment, an environmental test chamber (DOAHO TEST, - 20 ∼ 180 °C) was used to control the temperature, and there were air holes sealed with rubber plug on its both sides, which was convenient to extend the jumper of the sensor. The sensor was immersed in the solution to be tested.
Comparative experiments were carried out to characterize the sensor design with its counterpart with only metal coating but no PDMS. The NCF region of the sensor was immersed in deionized water. Figure 4 shows the transmission spectrum of the proposed sensor (red dotted line) and its counterpart without PDMS coating (red solid line). The spectral characteristics of silver and gold coating areas to the surrounding media when they were immersed separately are indicated by the spectral curves of black solid line and black dotted line. It is easy to distinguish that the respective resonance wavelength of SPR produced by the silver and gold film is different, the former corresponding to 492 nm while the latter corresponding to 609 nm. In the mixed spectrum indicated by red solid line, due to the close distance between the dips of gold and silver coating, there is only one obvious resonant dip with lower transmittance, which would lead to a difficulty in the dual channel spectrum demodulation. However, after coating PDMS on the gold film, the resonance dip of Channel B shifted to 900 nm. The wavelength spacing between the dips of the two channels was increased, which not only made it possible to distinguish easily the resonance dip of two channels in the spectrum, realizing the measurement of the liquid concentration and temperature at the same time, but also provided a larger detection range for the sensor.
The performance of the proposed sensor at single parameter measurement needs to be carefully evaluated before conducting the dual parameter measurement. First, its performance at detecting glucose concentration is evaluated. Glucose powder (purchased from Merck, Germany) was dissolved in deionized water by electronic balance, and glucose solution with concentration range of 5 ∼ 25% (interval of 5%) was prepared for testing. Figure 5(a) and 5(b) show the corresponding transmission spectrum recorded by the sensor at 30 °C when measuring the glucose solution with an increasing or a decreasing concentration. It can be seen that the resonance dip of channel A shifts from 508.542 to 565.528 nm with the increase of concentration, and from 565.173 nm to 508.513 nm with the decrease of concentration, which is consistent with the change trend obtained from the previous theoretical analysis. The average sensitivity of the sensor is 2.882 nm / % and 2.832 nm / %, respectively. The slight disturbance of resonance wavelength may come from the measurement error caused by the insufficient resolution of spectrometer (full width at half maxima, 1.5nm). Figure 5(c) fits the resonance wavelength of the corresponding concentration and shows the high linearity of the sensor. After analysis, the regression coefficient R-square value with the rise or fall of the glucose concentration was 0.99766 and 0.99833, respectively, which was of great significance in practical application. As can be seen from Fig. 6, although the intensity of transmitted light changes, the resonance dip of Channel B shows no shift, which indicates that Channel B is insensitive to glucose concentration. This is because its surface is wrapped by PDMS, which separates it from the measured liquid. As a result, the signal interference of Channel B to Channel A is avoided successfully.
Then the performance of the sensor in temperature measurement is evaluated. The sensor was placed in the environmental test chamber, and the temperature interval was set to be 10 °C. The heating and cooling processes from 20 °C to 60 °C were tested, as shown in Fig. 6(a) and 6(b). Due to the high thermo-optic coefficient of PDMS, the change of temperature will lead to a significant change of RI, and ultimately affects the resonance wavelength of Channel B. In this case, the RI of PDMS decreases with the increase of temperature. Therefore, it can be seen from the figure that when the temperature increases from 20 °C to 60 °C, the dip of Channel B moves to the short wavelength direction, and an opposite trend occurs when the temperature decreases from 60 °C to 20 °C, consistent with the change trend obtained from the previous theoretical analysis. Because the sensor was not placed in the solution, Channel A did not excite SPR, thus the spectrum showed only the dip of Channel B and the temperature sensing performance was unaffected. In the process of temperature rising, the resonance wavelength shifted from 783.814 nm (20 °C) to 900.024 nm (60 °C), and in the decreasing state, the resonance wavelength shifted from 900.329 nm (60 °C) to 783.489 nm (20 °C). The corresponding average temperature sensitivity was calculated to be -2.882 nm / °C for the former and -2.904 nm / °C for the latter. Figure 6(c) fits the measured data with temperature, which exhibits good linearity and stability at the temperature measurement. The respective regression coefficients were as high as 0.9992 and 0.9982. In addition, a new tiny dip was found at 474 nm, and this slight interference was most likely caused by the instability of the reference spectrum stored in the spectrometer, whose effect on the proposed sensor can be ignored.
Finally, the dual parameter sensing characteristics of the sensor were evaluated experimentally. For the first step, the sensor was placed in the environmental test chamber (temperature range: 20 ∼ 60 °C, interval: 10 °C) after immersing the NCF area in 10% glucose solution. Considering that it will take some time for the solution to reach the same temperature and keep stable in the test chamber, we recorded the measurement data every 10 minutes after reaching the preset temperature, and the experimental transmission spectra was shown in Fig. 7(a). It can be seen that as the temperature increased, the dips of both channels blue shifted, and that of Channel A moved from 525.141 nm to 513.603 nm. Because the glucose solution has the similar thermo-optic coefficient with the deionized water (∼ - 1.12 × 10−4 RIU / °C) , the change of temperature would produce the same effect upon them: when measuring the same glucose concentration, different temperature would lead to a shift of the resonance wavelength of Channel A, whose moving trend is contrary to that induced by the increase of glucose concentration. Consequently, the combined working of different glucose concentration and different temperature may lead to a coincidence of the resonance wavelengths of Channel A, resulting in an error in the measurement results of the sensor. To guarantee the accuracy of real-time measurement of glucose concentration, the solution temperature should be taken into consideration.
Then, the temperature variation range was set to be 20 ∼ 60 °C and the concentration of the glucose solution 5 ∼ 25%. Due to long time exposure to the air, the container was sealed to prevent the change of solution concentration caused by water evaporation. The corresponding resonance wavelength of Channel A was recorded, and the fitted straight line was shown in Fig. 7(b). It can be seen that with the increase of temperature, the resonance wavelength of all glucose solutions with different concentrations shows blue shifts. By transforming the x-axis into the concentration variation range, the relationship of solution concentration, temperature, and the corresponding resonance wavelength were demonstrated more clearly (see Fig. 7(c)). It was to be noted that since Channel B was not affected by solution concentration, the resonance wavelength corresponding to temperature change should be consistent with that in Fig. 6.
Ignoring the data deviation caused by the insufficient resolution of the spectrometer, Figs. 7(b) and 7(c) were used for mathematical deduction, through which the concentration and temperature were calibrated corresponding to each measurement result obtained by the sensor. We obtained the intercept of the five fitting lines in Fig. 7(c), and gave the relationship between the intercept and the corresponding temperature (see Fig. 7(d)). It can be seen that the intercept has a high linearity with temperature, and the relationship between the temperature, the glucose solution concentration and the detected resonance wavelength can be expressed as follows:7(c). It can be seen from the formula that Y is affected by C and T. At the operation of this proposed sensor, T and Y can be experimentally obtained, and based on Eq. (3), the corresponding glucose concentration at this temperature can be obtained.
Repeatability and stability are internal quality evaluation indicators of sensors, and usually sensors with poor repeatability have problems in their internal mechanism, leading to a poor stability. In this work, in order to verify the stability and repeatability of the proposed sensor, two additional tests were conducted. First, we put the sensor in 5%, 10% and 15% glucose solution (30 °C), recorded the measurement data of Channel A, took the sensor out, cleaned and dried it, and repeated the whole process 5 times. For the five repeated measurements, the maximum fluctuation △λ of resonance wavelength was less than 0.334 nm, 0.715 nm and 0.309 nm respectively (see Fig. 8(a)). Next, it was placed in the environmental test chamber, and the measured data of Channel B were recorded respectively at 20 °C, 30 °C and 40 °C for 5 times. The corresponding maximum fluctuation △λ was ≤ 0.611 nm, 0.621 nm and 0.652 nm (see Fig. 8(b)). The test results show that the performance of our sensor is relatively stable and repeatable, confirming its application significance.
Finally, a comparison with the representative glucose concentration and temperature sensors of different optical configurations is listed in Table 1, which highlights the advantages of our proposed sensor. However, our sensor can not eliminate the interference of non-glucose factors in complex environment, thus it is necessary to functionalize the optical fiber surface in order to meet the needs of biosensing. In addition, the oxidation of silver can affect the performance of the sensor. Once silver is oxidized, the corresponding resonance wavelength will shift, which will eventually lead to an instability of the sensing performance. However, compared with the high temperature environment, the working range of the sensor is 20 ∼ 60 °C, which makes the oxidation speed very slow. It is worth noting that sulfur-containing substances in the air are easy to react with silver to form silver sulfide and other substances attached to the surface of silver film. Therefore, the proposed sensor cannot work in such an environment for a long time, and needs to be calibrated regularly to ensure the best performance.
In this work, we experimentally demonstrated a compact dual channel SPR NCF sensor which can measure both glucose concentration and temperature. The results showed that the glucose concentration sensitivity and temperature sensitivity were 2.882 nm / % and - 2.904 nm / °C, respectively. Moreover, by carefully calibrating the ambient temperature, accurate measurement of glucose concentration was achieved. The proposed sensor has high stability, repeatability, simple structure and easy fabrication, which makes it very suitable for sensing glucose solution or other analyte solutions requiring temperature compensation.
Higher Education Discipline Innovation Project (B16009); Japan Society for the Promotion of Science (KAKENHI Grant 17K18891, KAKENHI Grant 18H01504); National Key Research and Development Program of China (2017YFA0701201, 2019YFB2204001); National Natural Science Foundation of China (11604042, 61775032); Fundamental Research Funds for the Central Universities (N180406002, N180408018, N2004021); JSPS and CERN under the JSPS-CERN joint research program.
The authors thank Liao Ning Revitalization Talents Program.
The authors declare no conflicts of interest.
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