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High-sensitivity glycated hemoglobin (HbA1c) aptasensor in rapid-prototyping surface plasmon resonance
Sensors & Actuators: B. Chemical 279 (2019) 267–273
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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
High-sensitivity glycated hemoglobin (HbA1c) aptasensor in rapidprototyping surface plasmon resonance ⁎
Chen Guang Zhanga, Shwu Jen Changb, Kalpana Settuc, Ching Jung Chend, , Jen Tsai Liue,
T ⁎
a
School of Electronic, Electrical and Communication Engineering, University of Chinese Academy of Sciences, PR China Department of Biomedical Engineering, I-Shou University, Kaohsiung City, 82445, Taiwan c Department of Electrical Engineering, National Taipei University, Sanxia, Taiwan d School of Opto-Electronic Technology, University of Chinese Academy of Sciences, PR China e College of Materials Science and Opto-Electronic Technology, University of Chinese Academy of Sciences, PR China b
A R T I C LE I N FO
A B S T R A C T
Keywords: 3D printing Surface plasmon resonance Aptamer HbA1c Kinetics parameters
Diabetes mellitus is one of the most common noncommunicable diseases in the world. Fasting blood glucose level is one of the diabetes diagnostic criteria, but it is readily affected by stress, exercise and many other factors. Glycated hemoglobin (HbA1c) is an important clinical index for diabetes and related diseases. In this study, a cheaper, rapid-prototyping, high-sensitivity angle-scanning surface plasmon resonance (SPR) was developed by fused deposition modeling (FDM) 3D printing technology for HbA1c detection. The sensor chip is based on an aptamer that has high affinity and high specificity to HbA1c. The results showed that this HbA1c-specific aptamer had high specificity for HbA1c, and the calculated KD was 6.13 × 10−8 M. The linear detection response of HbA1c appeared in the range of 18–147 nM, with a detection limit of 1 nM. This SPR HbA1c detection system could be a promising platform for developing clinical point-of-care diagnostic applications.
1. Introduction Diabetes mellitus is one of the most common noncommunicable diseases in the world. Statistics in 2013 showed that there were approximately 382 million people suffering from diabetes, and the number of diabetes cases is increasing fast and is predicted to reach 592 million by 2035 [1]. Elevated glucose during fasting is one of the widely used diabetes diagnostic criteria, but it is readily affected by stress, exercise and many other factors. Compared with glucose level, the amount of glycated hemoglobin (HbA1c) reflects the average concentration of glucose over the past 100 to 200 days [2]. Glycated hemoglobin (HbA1c) is a glycated protein formed by a nonenzymatic reaction of glucose in the human body [3]. The HbA1c level, defined as the ratio between HbA1c concentration and total hemoglobin concentration, is an important clinical index for diabetes and related diseases [4]. Currently, HbA1c is measured by fluorescence sensing, cation-exchange chromatography, electrophoresis, and boronate affinity chromatography in clinics. However, these methods have some limitations, such as the need for labeling assays, drug interference or error due to wide detection ranges [5]. Aptamers are oligonucleotide or peptide molecules that bind to a specific target molecule. Aptamer-based assays hold great promise to
⁎
replace antibody-based assays, as aptamers are relatively cheap and easy to store, even at room temperature [6]. Li et al. [7] reported that the HbA1c-specific aptamer selected by SELEX can be chemically synthesized easily with high reproducibility at relatively low cost. The HbA1c aptamers showed high specificity and high affinities. Thus, it has great potential to replace the HbA1c antibody for the detection of HbA1c. Aptamer-based sensors (aptasensors) are mainly divided into two categories, electrochemical [8] and optical [9]. Electrochemical aptasensors are widely used due to their high sensitivity and low cost without any optical devices. However, they are usually affected by an unstable structure of aptamer on the electrode. Optical aptasensors are often based on labeling methods such as chemiluminescence or welldeveloped fluorescence probes. It is an indirect method to detect the target, it is difficult to collect all signals in time, and fluorescence decays gradually with time. Additionally, the aptamer-modified fluorescent probe might influence the DNA structure and affect the binding affinity. However, due to the advantages of DNA aptamers that have high affinity, high specificity and ease of storage, they have been commonly used in the existing biomolecule detection probes. For this reason, it is very important to develop a direct detection method without any labeling. Surface plasmon resonance (SPR) is a surface-sensitive optical
Corresponding authors at: No.19A Yuquan Road, Beijing, 100049, PR China. E-mail addresses: [email protected] (C. Jung Chen), [email protected] (J.T. Liu).
https://doi.org/10.1016/j.snb.2018.09.077 Received 10 February 2018; Received in revised form 11 September 2018; Accepted 18 September 2018 Available online 19 September 2018 0925-4005/ © 2018 Elsevier B.V. All rights reserved.
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Fig. 1. The scheme of the sensor modification process.
Tany TechnoGene Ltd. (Ness Ziona, Israel). 11-Mercaptoundecanoic acid (MUA) was purchased from Kuer Co. Ltd. (AnHui, China). NHydroxysulfosuccinimide sodium salt (sulfo-NHS), 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC) and 3-mercaptopropionic acid (MPA) were purchased from Aladdin Chemistry Co. Ltd. (Shanghai, China). Sodium dodecyl sulfate (SDS) was obtained from Beijing Biotopped Science & Technology Co. Ltd. (Beijing, China). Absolute ethyl alcohol and NaCl were purchased from Beijing Chemical Works (Beijing, China). Deionized water obtained from a CCT-3300 water purification system was used in all the experiments. Matching oil series E(1.520) was purchased from Cargill Dow LLC (USA). Polylactide (PLA) 3D printing filaments were purchased from 3D Systems (USA).
technique that is used to study a thin layer on a metal surface. SPR is a powerful analytical technology that can detect the thickness of the films absorbed onto the sensor surface and interactions between biomolecules, such as antigen-antibody or protein-DNA [10–14]. Compared with traditional detection methods, such as high-performance liquid chromatography (HPLC), enzyme-linked immunosorbent assays (ELISA), and fluorescence sensing methods, the major benefits of SPRbased sensors are the potential for label-free detection of analytes and the ability to overcome additive interferences. Moreover, SPR has numerous advantages, such as reliable instrumentation, automation, disposable sensor chips, and versatility due to the wide variety of surface chemistries and assay methods available for various biomolecules. SPR has also been applied in chemical, biology, agriculture, environment and food safety detection [15–21]. However, SPR systems usually require expensive equipment and complicated optics. The cost of commercial SPRs varies from $10,000 to $5,000,00, and the refractive index resolution ranges from 10−5 refractive index units (RIU) to 10−7 RIU. In previous studies [22–26], many low-cost SPR systems have been introduced, but the resolution and dynamic range are often sacrificed. Hence, developing an efficient SPR system with a modular design and use of advanced hardware resources is of great interest. 3D printing, also called additive manufacturing, has been a tool for developing rapid prototyping products since the 1980s. After the expiration of 3D printing technology patents, this field has witnessed great growth [27]. According to the open-source RepRap project, desktop FDM 3D printers become more popular among users, manufacturers and researchers due to their simplicity, cost-effectiveness and versatility. Application of FDM 3D printing technology is fast emerging in the development of prototypes, scientific tools and medical equipment [28–31]. Using 3D printing technology to design an SPR is an interesting and attractive option. Recently, some studies have demonstrated the use of 3D printing in developing SPR platforms. Hasan et al. [25] utilized the 3D printing technique to print the device holder for a smartphone-based SPR imaging platform. That system has a dynamic range less than 0.02 RIU, which limits the application for measuring large refractive index change. Hence, utilizing 3D printing technology to construct an SPR modularization for developing a high-sensitivity SPR system seems to be challenging. In this study, a low-cost, rapid-prototyping, high-sensitivity anglescanning SPR was developed by FDM 3D printing technology for HbA1c detection. To the best of our knowledge, this is the first study to develop an SPR system based on 3D printing SPR modularization for HbA1c detection. The SPR sensor chip was immobilized with an aptamer that had high affinity and high specificity to HbA1c. In addition, using highsensitivity SPR, we sought to detect HbA1c directly without any labeling and to arrive at a proper diagnosis in a clinical system.
2.2. Preparation of HbA1c aptamer sensor The gold sensor chip was first washed with SDS solution and exposed under UV-ozone for 20 min. The sensor chip was then immersed in 10 mM 3-mercaptopropionic acid and 11-mercaptoundecanoic acid (volume ratio 10:1) for 3 h. Then, the sensor chip was rinsed with absolute ethyl alcohol. Subsequently, the chip was incubated for 1 h in a solution that contained 1 μM DNA, 5 mM EDC and 3 mM NHS. Finally, the DNA aptamer–immobilized gold sensor chip was rinsed in deionized water and dried with a nitrogen gun. The preparation process of the aptamer sensor is schematically shown in Fig. 1 2.3. SPR system design The developed rapid-prototyping angle-scanning SPR system block diagram is shown in Fig. 2a, and the realized SPR device is depicted in Fig. 2b. Light from the laser source was first p-polarized by a polarizer, then projected onto the sensor surface via a triangle prism. The reflected light from the sensor surface was collected by a photodiode. The collected raw signal was amplified by a homemade amplifier circuit and sampled by the built-in ADC (analog to digital converter) unit of ATmega2560 microprocessor. The rotating platform could drive the system to scan from 40° to 72°. When performing the measurements, an angle range of 3° around the resonance angle was scanned at a speed of 0.3°/s instead of scanning a whole SPR spectrum. A peristaltic pump was used to transport the sample. 2.4. Electric circuit unit Fig. 2c shows the block diagram of the electric circuit associated with the proposed SPR system. Two microcontroller units (MCUs) were used to control the system hardware. The rotating platform comprised of the mechanical driving unit and an end-stop sensor. The end-stop sensor was used as the coordinate origin reference. All the abovementioned components, including the peristaltic pump and ATmega328, were controlled by ATmega2560 via different ports. The amplified signal from the photodiode was sampled at 1 kHz by the ATmega2560 built-in 10 bit Analog-to-Digital Converter (ADC) module. ATmega328 MCU was used to monitor the system temperature. Low temperature drift coefficient negative temperature coefficient (NTC) resistors (103AT-4 shape1) were employed as the temperature sensors. Signals from NTC resistors were quantified by a 16-bit high-precision ADC ADS1115. ATmega328 obtained the temperature information from
2. Materials and methods 2.1. Chemicals and reagents DNA sequences were purchased from Beijing Chief-East Tech Co., Ltd. (Beijing, China). The sequence of the HbA1c-specific DNA aptamer used in this work was 5′-GGC AGG AAG ACA AAC ACA TCG TCG CGG CCT TAG GAG GGG CGG ACG GGG GGG GGC GTT GGT CTG TGG TGC TGT-3′. Glycated hemoglobin (68 kDa) was purchased from ProSpec268
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Fig. 2. Schematic representation of the complete SPR system: (a) system block diagram; (b) the SPR biosensing platform designed by rapid prototyping 3D printing; (c) electric circuit block diagram; (d) flowcell diagram.
Fig. 3. NaCl measurement curves: (a) response for different concentrations of NaCl solution; (b) △RU vs refractive index.
2.5. SPR modular part design
ADS1115 via the Inter-Integrated Circuit (IIC) communication protocol and controlled the TEC drivers by IO ports. To maintain a stable temperature, the proportional–integral–derivative (PID) controller algorithm was incorporated with ATmega328.
The SPR rapid-prototyping system was constructed with the 3Dprinted components. The Cube3 desktop version 3D printer was used to print the components with PLA filaments. The component models were designed using SolidWorks 2012 software and saved in Standard Triangle Language (STL) file format. Then, these STL files were 269
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Fig. 4. Temperature curves: (a) temperature regulation curve; (b) system temperature versus time; (c) system response against temperature change; (d) system response versus temperature.
Fig. 5. Contact angle measurement. a) Bare Au film. b) Au film after 20 min UV exposure. c) Au/MUA/MPA SAM. d) Au/MUA/MPA/aptamer film.
accessories, 3D-printed components, and a rubber sample pretreatment channel. The arrangement of the whole flow cell module is shown in Fig. 2d. The rubber flow cell of volume 65 μL and dimension 5.3 mm × 23 mm × 0.5 mm was designed and constructed using desktop cutter Silhouette Portrait. A sample pretreatment channel was also made of rubber film and cut by Silhouette Portrait. To let the sample attain a stable experimental temperature before reaching the main flow cell, the pretreatment channel was designed to be as long as possible. These components were mounted onto a stainless block,
imported into the 3D printer’s software Cubify version 3.9.0 (3Dsystems, USA). The design of the optical light path is presented in Video S1 (Supplementary material). Two rotating platforms were employed, and a 3D printed optical platform was installed onto the rotating platforms. The laser source and photodiode could be easily mounted onto the optical platform with the 3D-printed holders. Furthermore, the alignment of optical parts could be achieved effortlessly. The flow cell module consisted of a rubber flow cell, stainless 270
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Table 1 Comparison of different HbA1c measurement methods. Method
Linear range (nmol/L)
Limit of detection (nmol/L)
Antibody
Reference
Aptamer-based sandwich Electrochemistry
104 μM–313 μM
–
No
Li et al. [7]
0.3 μM–2000 μM
0.2 μM
No
Chemiluminescence
30 nM–1200 nM
–
No
Microfluidic immunoassay Potentiometric method Surface plasmon resonance
0.75 nM–22.5nM
0.37 nM
Yes
89 μM–255 μM
–
Yes
18 nM–147 nM
1 nM
No
Jain et al. [39] Ahn et al. [38] Chen et al. [40] Liu et al. [37] This work
2.6. Resolution determination
Fig. 6. SPR response of aptasensor with BSA and HbA1c.
Since the refractive index of NaCl solution could be obtained easily, different concentrations of NaCl solution (0.5%, 1%, 2%, 3% and 4%) were prepared and used as standard sample solutions to determine the resolution of the developed system. The sample delivery rate was set as 50 μL/min, and the system temperature was set as 20 °C. The system response for NaCl concentrations was recorded and used to compute the system resolution. 2.7. Temperature effect test Deionized water was used as analyte, and its resonance angle was monitored. The sample delivery speed was set as 50 μL/min. The range of temperatures during the experiment was 20 °C–28 °C. To avoid bubbles in the flow cell, the temperature of the experiment was increased and decreased gradually. 2.8. Contact angle measurement The water contact angle was measured by a Tianmin contact angle meter. The SPR sensor chip was mounted on the plane of the contact angle meter. Four microliters of deionized water was dropped onto the sensor chip surface. The droplet cross-section was recorded by the software provided by manufacturer. Contact angle was estimated from cross-sectional image using open source software ImageJ.
Fig. 7. Association curves of aptamer and HbA1c. The inset shows kapp versus concentration of HbA1c.
2.9. Kinetic parameters The kinetic parameter measurement capability of the developed system was validated using the HbA1c-specific aptamer. Affinity interactions between aptamer and HbA1c were characterized by the association rate constant ka, dissociation rate constant kd and equilibrium association constant KD. The experiment data were fitted with a 1:1 interaction model, A + B = AB, where A is the injected analyte, B is the immobilized ligand, and AB is the analyte-ligand complex. In principle, the interaction process can be described as shown in Eq. (1), where the response signal R of the SPR system is proportional to [AB], the maximum response signal Rmax is proportional to the initial [B], kapp is the apparent reaction rate constant, and t is the interaction time. From Eq. (1), we can obtain ka and kd from kapp. A series of concentrations of solution A was used to derive the relationship between kapp and [A], and a regression curve was obtained from Eq. (2). Thus ka is the slope of the regression, kd is the intercept of the regression curve, and KD is computed as ka / kd.
Fig. 8. System responses for different concentrations of HbA1c.
whose temperature was maintained stable. The assembling process of the flow cell module is described in Video S2 (Supplementary material). Refractive index matching oil was used to provide a gap between the Au-coated glass slide and the prism.
R=
271
ka [A] Rmax [1 − e−kapp t ], kapp = [A] ka + kd ([A] ka + kd )
(1)
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⎧ kapp1 = [A1 ] ka + kd ⎪k app2 = [A2 ] k a + k d ... ⎨ ⎪ kappn = [An ] ka + kd ⎩
Au film. When the MUA/MPA self-assembled monolayers were formed on Au film, the contact angle became 47.605°, which indicates the termination of the MUA/MPA self-assembled monolayer (SAM) with a hydrophilic domain, eCOOH [34]. Subsequently, after the immobilization of DNA onto the MUA/MPA SAM, the contact angle decreased to 33.030° from 46.720° of the MUA/MPA SAM. 21.911° The decrease in contact angle after DNA immobilization confirmed the successful immobilization of DNA onto the sensor surface [8].
(2)
3. Results and discussions 3.1. SPR refractive index resolution
3.4. Characterization of specificity
The strength and precision of the 3D-printed components can affect the performance of the developed system. Unlike the metal materials, the printed PLA components may deform greatly when a laser or photodiode is mounted on them. The tensile strength of 3D-printed parts was mainly influenced by the infill density. Thus, the infill density should be as high as possible. A 50% infill density is suitable for the trade-off between strength, printing time and material cost. Imprecision of the 3D-printed parts might have caused alignment problems on the optical path or assembly issues. If the mechanical adjustment of the printer was considered ideal, the precision of the printed parts would be affected by the nozzle diameters and printing layer thickness. The nozzle diameter affected the precision in the xy-plane, while the layer thickness affected the precision along the z-axis. These errors should be compensated for at the design stage. To evaluate the angular resolution of the developed system, the deionized water response was used as the background signal, and the signal noise level was 0.0008°, which can be seen as the angular resolution of our system. The system response for different concentrations of NaCl is shown in Fig. 3a. The regression equation obtained was RI = 0.00799 × [ΔRU] +1.333 (R2 = 0.994), as shown in Fig. 3b, where RU is the abbreviation of response unit (the resonance angle in degrees) and ΔRU is the angle shift of the resonance angle corresponding to the baseline signal; the sensitivity was 0.00799°/RIU. According to [25], the resolution of the refractive index can be obtained as the product of the noise level and the sensitivity. It can be inferred that the developed system exhibited a high refractive index resolution of 6.39 × 10−6 RIU.
The BSA protein contains hydrophobic, hydrophilic, cationic, and anionic regions. Therefore, the nonspecific adsorption of proteins may result from hydrogen bonding, electrostatic, charge-transfer, and/or hydrophobic interactions, depending on the surface properties. Therefore, BSA was suitable to check the specificity of the probe. To test the nonspecific binding ability of the HbA1c-specific aptamer–immobilized sensor, 149 nM BSA solution was employed as a control, and the measurement result is illustrated in Fig. 6. The resonance angle shifts of HbA1c (147 nM) and BSA (149 nM) were 0.2168 and 0.0005 RU, respectively. This result demonstrates that the HbA1c aptamer–functionalized sensor did not adsorb any BSA and hence was not specific to BSA. Thus, the aptamer-based sensors developed in this study were more specific to HbA1c. 3.5. Kinetics parameter measurement The interactions between aptamer and different concentrations of HbA1c solution are plotted in Fig. 7. A higher concentration of HbA1c showed a faster association rate. At lower concentrations of HbA1c, the association process occurred slowly. The relationship between kapp and the concentration of HbA1c is plotted as shown in the inset of Fig. 7. The linear regression equation was obtained as kapp = 5 × 103 × [C] + 0.00031 (R2 = 0.982), kapp = 0.00031 × [C ] + 0.00097 where C is the HbA1c concentration. Thus, the kinetic parameter kd was first obtained as 3.1 × 10―4 s―1 9.7 × 10−4s−1, ka was 5 × 103 M―1s―1, and Keq was 6.13 × 10―8 M, which was close to the KD value K eq ∼ 4 × 1011M−1 reported by Lin et al. [35].
3.2. Temperature effect Temperature has a significant influence on SPR measurement [32]. A one-degree change in temperature alters the refractive index resolution by ∼10―4. Fig. 4a shows the temperature regulation curve. By carefully adjusting the PID parameters (P = 2.7, I = 0.0001, D = 0.0008), the temperature controller module reached a stable state within 3 min and maintained the system temperature variation within ± 0.05 °C. Fig. 4b shows the temperature change versus time, and the system responses for different temperature are shown in Fig. 4c. The results showed that the system response decreased when the temperature increased. The system response exhibited a linear relationship with temperature between 20 °C and 28 °C (Fig. 4d), and the linear equation was ΔRU = −0.015 × T + 0.300 (R2 = 0.997). Considering the refractive index resolution computed in the previous section (Section 3.1), we can easily obtain dRI / dT = 1.3 × 10−4 / °C , which is consistent with the previously reported study [33].
3.6. HbA1c detection To estimate the detectable range for clinical diagnosis application, we designed a concentration range of HbA1c between 18 nM (1.22 μg/ mL) and 147 nM (10.00 μg/mL), which was much lower than the physiologic range (3–13 mg/ml) [36]. A calibration curve was obtained with known amounts of HbA1c. Fig. 8 depicts the SPR system responses of the aptasensor with HbA1c concentrations ranging from 18 to 147 nM. The binding interaction between HbA1c and the immobilized aptamer on the sensor surface caused the system response to increase linearly with increasing HbA1c concentration, as seen from Fig. 8. According to the of signal:noise ratio criterion, the limit of detection (LOD) calculated was 1 nM (0.070 μg/mL). On comparing the analytical performance of the proposed SPR HbA1c aptamer sensor with some previously reported [37–40] biosensors for HbA1c detection (Table 1), we found that the present sensor exhibited a low LOD and a wide detection range at nanomolar concentration. In addition, the ability of the SPR system to make measurements on very small blood samples can reduce the substrate interference and enhance the signal/noise ratio and further improve the accuracy of HbA1c detection. Thus, the developed SPR HbA1c detection system can be considered a promising method to measure the concentration of HbA1c for clinical diagnosis. However, the dilution of the actual sample by one thousandfold may not be practical for now. But this study demonstrates the potential of an HbA1c aptasensor that can using 1 μL, or even less, of actual sample for detecting HbA1c in the future.
3.3. Contact angle Prior to the biosensing experiment with the SPR system and sensing chip, the effectiveness of the modified sensor surface needed to be verified. Contact angle measurement is a powerful tool to examine every step during the sensor functionalization. The results of contact angle measurement are shown in Fig. 5. The bare Au surface exhibited a contact angle value of 76.084°. After 20 min of UV exposure, the contact angle decreased to 65.985°. This decrease in contact angle indicated that the UV exposure effectively removed the impurities on the 272
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4. Conclusion
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In this investigation, a cheap, rapid-prototyping, high-sensitivity SPR was developed by FDM 3D printing technology, and the results demonstrated that the SPR HbA1c aptamer sensor could successfully detect and quantify the binding of HbA1c to an HbA1c-specific aptamer monolayer. Moreover, the HbA1c aptamer monolayer showed high specificity to HbA1c. The linear part of the calibration curve included the HbA1c range, from 18 to 147 nM, with an LOD of 1 nM. Taking advantage of SPR and the aptamer-based assay, this developed SPR system could be a promising prototype for providing a point-of-care device for diabetic patients. Acknowledgments Thiswork was supported by the State Key Project of Fundamental Research (Grant 2014CB931900), by the UCAS Young Teacher Research Fund (Grant Y55103NY00, Y55103EY00, Y65201FY00 and Y25102TN00), by the Beijing Natural Science Foundation (Z160002), and by the Chinese Academy of Sciences Key project foundation (KFZDSW-202). Appendix A. Supplementary data Supplementary material related to this article can be found, in the online version, at doi: https://doi.org/10.1016/j.snb.2018.09.077. References [1] R.F.J. Da, et al., IDF Diabetes Atlas estimates of 2014 global health expenditures on diabetes, Diabetes Res. Clin. Pract. 117 (2016) 48. [2] J.T. Liu, et al., The investigation of recognition interaction between phenylboronate monolayer and glycated hemoglobin using surface plasmon resonance, Anal. Biochem. 375 (1) (2008) 90. [3] J. Halámek, U. Wollenberger, Development of a biosensor for glycated hemoglobin, Electrochim. Acta 53 (3) (2008) 1127–1133. [4] A.T. Kharroubi, et al., Evaluation of glycated hemoglobin (HbA1c) for diagnosing type 2 diabetes and prediabetes among Palestinian Arab population, PLoS One 9 (2) (2014) e88123. [5] C. Weykamp, HbA1c: a review of analytical and clinical aspects, Ann. Lab. Med. 33 (6) (2013) 393–400. [6] S.D. Jayasena, Aptamers: an emerging class of molecules that rival antibodies in diagnostics, Clin. Chem. 45 (9) (1999) 1628. [7] J. Li, et al., On-chip, aptamer-based sandwich assay for detection of glycated hemoglobins via magnetic beads, Biosens. Bioelectron. 79 (2016) 887–893. [8] G. Wang, et al., Antifouling aptasensor for the detection of adenosine triphosphate in biological media based on mixed self-assembled aptamer and zwitterionic peptide, Biosens. Bioelectron. (2017). [9] A. Wochner, et al., A DNA aptamer with high affinity and specificity for therapeutic anthracyclines, Anal. Biochem. 373 (1) (2008) 34–42. [10] I. Stojanović, R.B.M. Schasfoort, L.W.M.M. Terstappen, Analysis of cell surface antigens by surface plasmon resonance imaging, Biosens. Bioelectron. 52 (2014) 36–43. [11] K.M. Byun, et al., Enhanced surface plasmon resonance detection of DNA hybridization based on ZnO nanorod arrays, Sens. Actuators B Chem. 155 (1) (2011) 375–379. [12] J.R. Wayment, J.M. Harris, Biotin-avidin binding kinetics measured by single-molecule imaging, Anal. Chem. 81 (1) (2009) 336–342. [13] T. Hiragun, et al., Surface plasmon resonance-biosensor detects the diversity of responses against epidermal growth factor in various carcinoma cell lines, Biosens. Bioelectron. 32 (1) (2012) 202–207. [14] A. Kausaite, et al., Surface plasmon resonance label-free monitoring of antibody antigen interactions in real time, Biochem. Mol. Biol. Educ. 35 (1) (2007) 57–63. [15] C. Wu, et al., Real-time evaluation of live cancer cells by an in situ surface plasmon resonance and electrochemical study, ACS Appl. Mater. Interfaces 7 (44) (2015) 24848–24854. [16] Z. Zhu, et al., An aptamer based surface plasmon resonance biosensor for the detection of ochratoxin A in wine and peanut oil, Biosens. Bioelectron. 65 (2015) 320–326. [17] H. Vaisocherová, et al., Rapid and sensitive detection of multiple microRNAs in cell lysate by low-fouling surface plasmon resonance biosensor, Biosens. Bioelectron. 70 (2015) 226–231. [18] F. Fernández, et al., A label-free and portable multichannel surface plasmon resonance immunosensor for on site analysis of antibiotics in milk samples, Biosens.
ChenGuang Zhang is a graduate student at University of Chinese Academy of Sciences. He received his BS degree in 2014 from Guangdong University of Technology. His current research is the design and implementation of surface plasmon resonance systems. Kalpana Settu is assistant professor at National Taipei University. She received his MS degree in 2012 and her Ph.D. in 2016 in Electrical Engineering, National Central University, her current research work has been mainly focused on biosensors, bioelectrics and biomaterials. Shwu Jen Chang is professor at I-Shou University. She received her MS degree in 1993 from Chung Yuan Christian University, and her Ph.D. in 1999 in National Yang-Ming University, Taiwan. Her current research work has been mainly focused on biosensors and biomaterials. Ching-Jung Chen is associate professor at University of Chinese Academy of Sciences. She received his BS degree in 2002 and MS degree in 2004 from I-Shou University, and her Ph.D. in 2010 in Electrical Engineering, National Central University, Taiwan. Her current research work has been mainly focused on biosensors, bioelectrics and biomaterials. Jen-Tsai Liu is associate professor at University of Chinese Academy of Sciences. He received his BS degree in 2002 and MS degree in 2004 from I-Shou University, and his Ph.D. in 2010 in Chemical and Materials Engineering, National Central University, Taiwan. His current research work has been mainly focused on biosensors and bioelectrics.
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Sensors and Actuators B: Chemical journal homepage: www.elsevier.com/locate/snb
High-sensitivity glycated hemoglobin (HbA1c) aptasensor in rapidprototyping surface plasmon resonance ⁎
Chen Guang Zhanga, Shwu Jen Changb, Kalpana Settuc, Ching Jung Chend, , Jen Tsai Liue,
T ⁎
a
School of Electronic, Electrical and Communication Engineering, University of Chinese Academy of Sciences, PR China Department of Biomedical Engineering, I-Shou University, Kaohsiung City, 82445, Taiwan c Department of Electrical Engineering, National Taipei University, Sanxia, Taiwan d School of Opto-Electronic Technology, University of Chinese Academy of Sciences, PR China e College of Materials Science and Opto-Electronic Technology, University of Chinese Academy of Sciences, PR China b
A R T I C LE I N FO
A B S T R A C T
Keywords: 3D printing Surface plasmon resonance Aptamer HbA1c Kinetics parameters
Diabetes mellitus is one of the most common noncommunicable diseases in the world. Fasting blood glucose level is one of the diabetes diagnostic criteria, but it is readily affected by stress, exercise and many other factors. Glycated hemoglobin (HbA1c) is an important clinical index for diabetes and related diseases. In this study, a cheaper, rapid-prototyping, high-sensitivity angle-scanning surface plasmon resonance (SPR) was developed by fused deposition modeling (FDM) 3D printing technology for HbA1c detection. The sensor chip is based on an aptamer that has high affinity and high specificity to HbA1c. The results showed that this HbA1c-specific aptamer had high specificity for HbA1c, and the calculated KD was 6.13 × 10−8 M. The linear detection response of HbA1c appeared in the range of 18–147 nM, with a detection limit of 1 nM. This SPR HbA1c detection system could be a promising platform for developing clinical point-of-care diagnostic applications.
1. Introduction Diabetes mellitus is one of the most common noncommunicable diseases in the world. Statistics in 2013 showed that there were approximately 382 million people suffering from diabetes, and the number of diabetes cases is increasing fast and is predicted to reach 592 million by 2035 [1]. Elevated glucose during fasting is one of the widely used diabetes diagnostic criteria, but it is readily affected by stress, exercise and many other factors. Compared with glucose level, the amount of glycated hemoglobin (HbA1c) reflects the average concentration of glucose over the past 100 to 200 days [2]. Glycated hemoglobin (HbA1c) is a glycated protein formed by a nonenzymatic reaction of glucose in the human body [3]. The HbA1c level, defined as the ratio between HbA1c concentration and total hemoglobin concentration, is an important clinical index for diabetes and related diseases [4]. Currently, HbA1c is measured by fluorescence sensing, cation-exchange chromatography, electrophoresis, and boronate affinity chromatography in clinics. However, these methods have some limitations, such as the need for labeling assays, drug interference or error due to wide detection ranges [5]. Aptamers are oligonucleotide or peptide molecules that bind to a specific target molecule. Aptamer-based assays hold great promise to
⁎
replace antibody-based assays, as aptamers are relatively cheap and easy to store, even at room temperature [6]. Li et al. [7] reported that the HbA1c-specific aptamer selected by SELEX can be chemically synthesized easily with high reproducibility at relatively low cost. The HbA1c aptamers showed high specificity and high affinities. Thus, it has great potential to replace the HbA1c antibody for the detection of HbA1c. Aptamer-based sensors (aptasensors) are mainly divided into two categories, electrochemical [8] and optical [9]. Electrochemical aptasensors are widely used due to their high sensitivity and low cost without any optical devices. However, they are usually affected by an unstable structure of aptamer on the electrode. Optical aptasensors are often based on labeling methods such as chemiluminescence or welldeveloped fluorescence probes. It is an indirect method to detect the target, it is difficult to collect all signals in time, and fluorescence decays gradually with time. Additionally, the aptamer-modified fluorescent probe might influence the DNA structure and affect the binding affinity. However, due to the advantages of DNA aptamers that have high affinity, high specificity and ease of storage, they have been commonly used in the existing biomolecule detection probes. For this reason, it is very important to develop a direct detection method without any labeling. Surface plasmon resonance (SPR) is a surface-sensitive optical
Corresponding authors at: No.19A Yuquan Road, Beijing, 100049, PR China. E-mail addresses: [email protected] (C. Jung Chen), [email protected] (J.T. Liu).
https://doi.org/10.1016/j.snb.2018.09.077 Received 10 February 2018; Received in revised form 11 September 2018; Accepted 18 September 2018 Available online 19 September 2018 0925-4005/ © 2018 Elsevier B.V. All rights reserved.
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Fig. 1. The scheme of the sensor modification process.
Tany TechnoGene Ltd. (Ness Ziona, Israel). 11-Mercaptoundecanoic acid (MUA) was purchased from Kuer Co. Ltd. (AnHui, China). NHydroxysulfosuccinimide sodium salt (sulfo-NHS), 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC) and 3-mercaptopropionic acid (MPA) were purchased from Aladdin Chemistry Co. Ltd. (Shanghai, China). Sodium dodecyl sulfate (SDS) was obtained from Beijing Biotopped Science & Technology Co. Ltd. (Beijing, China). Absolute ethyl alcohol and NaCl were purchased from Beijing Chemical Works (Beijing, China). Deionized water obtained from a CCT-3300 water purification system was used in all the experiments. Matching oil series E(1.520) was purchased from Cargill Dow LLC (USA). Polylactide (PLA) 3D printing filaments were purchased from 3D Systems (USA).
technique that is used to study a thin layer on a metal surface. SPR is a powerful analytical technology that can detect the thickness of the films absorbed onto the sensor surface and interactions between biomolecules, such as antigen-antibody or protein-DNA [10–14]. Compared with traditional detection methods, such as high-performance liquid chromatography (HPLC), enzyme-linked immunosorbent assays (ELISA), and fluorescence sensing methods, the major benefits of SPRbased sensors are the potential for label-free detection of analytes and the ability to overcome additive interferences. Moreover, SPR has numerous advantages, such as reliable instrumentation, automation, disposable sensor chips, and versatility due to the wide variety of surface chemistries and assay methods available for various biomolecules. SPR has also been applied in chemical, biology, agriculture, environment and food safety detection [15–21]. However, SPR systems usually require expensive equipment and complicated optics. The cost of commercial SPRs varies from $10,000 to $5,000,00, and the refractive index resolution ranges from 10−5 refractive index units (RIU) to 10−7 RIU. In previous studies [22–26], many low-cost SPR systems have been introduced, but the resolution and dynamic range are often sacrificed. Hence, developing an efficient SPR system with a modular design and use of advanced hardware resources is of great interest. 3D printing, also called additive manufacturing, has been a tool for developing rapid prototyping products since the 1980s. After the expiration of 3D printing technology patents, this field has witnessed great growth [27]. According to the open-source RepRap project, desktop FDM 3D printers become more popular among users, manufacturers and researchers due to their simplicity, cost-effectiveness and versatility. Application of FDM 3D printing technology is fast emerging in the development of prototypes, scientific tools and medical equipment [28–31]. Using 3D printing technology to design an SPR is an interesting and attractive option. Recently, some studies have demonstrated the use of 3D printing in developing SPR platforms. Hasan et al. [25] utilized the 3D printing technique to print the device holder for a smartphone-based SPR imaging platform. That system has a dynamic range less than 0.02 RIU, which limits the application for measuring large refractive index change. Hence, utilizing 3D printing technology to construct an SPR modularization for developing a high-sensitivity SPR system seems to be challenging. In this study, a low-cost, rapid-prototyping, high-sensitivity anglescanning SPR was developed by FDM 3D printing technology for HbA1c detection. To the best of our knowledge, this is the first study to develop an SPR system based on 3D printing SPR modularization for HbA1c detection. The SPR sensor chip was immobilized with an aptamer that had high affinity and high specificity to HbA1c. In addition, using highsensitivity SPR, we sought to detect HbA1c directly without any labeling and to arrive at a proper diagnosis in a clinical system.
2.2. Preparation of HbA1c aptamer sensor The gold sensor chip was first washed with SDS solution and exposed under UV-ozone for 20 min. The sensor chip was then immersed in 10 mM 3-mercaptopropionic acid and 11-mercaptoundecanoic acid (volume ratio 10:1) for 3 h. Then, the sensor chip was rinsed with absolute ethyl alcohol. Subsequently, the chip was incubated for 1 h in a solution that contained 1 μM DNA, 5 mM EDC and 3 mM NHS. Finally, the DNA aptamer–immobilized gold sensor chip was rinsed in deionized water and dried with a nitrogen gun. The preparation process of the aptamer sensor is schematically shown in Fig. 1 2.3. SPR system design The developed rapid-prototyping angle-scanning SPR system block diagram is shown in Fig. 2a, and the realized SPR device is depicted in Fig. 2b. Light from the laser source was first p-polarized by a polarizer, then projected onto the sensor surface via a triangle prism. The reflected light from the sensor surface was collected by a photodiode. The collected raw signal was amplified by a homemade amplifier circuit and sampled by the built-in ADC (analog to digital converter) unit of ATmega2560 microprocessor. The rotating platform could drive the system to scan from 40° to 72°. When performing the measurements, an angle range of 3° around the resonance angle was scanned at a speed of 0.3°/s instead of scanning a whole SPR spectrum. A peristaltic pump was used to transport the sample. 2.4. Electric circuit unit Fig. 2c shows the block diagram of the electric circuit associated with the proposed SPR system. Two microcontroller units (MCUs) were used to control the system hardware. The rotating platform comprised of the mechanical driving unit and an end-stop sensor. The end-stop sensor was used as the coordinate origin reference. All the abovementioned components, including the peristaltic pump and ATmega328, were controlled by ATmega2560 via different ports. The amplified signal from the photodiode was sampled at 1 kHz by the ATmega2560 built-in 10 bit Analog-to-Digital Converter (ADC) module. ATmega328 MCU was used to monitor the system temperature. Low temperature drift coefficient negative temperature coefficient (NTC) resistors (103AT-4 shape1) were employed as the temperature sensors. Signals from NTC resistors were quantified by a 16-bit high-precision ADC ADS1115. ATmega328 obtained the temperature information from
2. Materials and methods 2.1. Chemicals and reagents DNA sequences were purchased from Beijing Chief-East Tech Co., Ltd. (Beijing, China). The sequence of the HbA1c-specific DNA aptamer used in this work was 5′-GGC AGG AAG ACA AAC ACA TCG TCG CGG CCT TAG GAG GGG CGG ACG GGG GGG GGC GTT GGT CTG TGG TGC TGT-3′. Glycated hemoglobin (68 kDa) was purchased from ProSpec268
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Fig. 2. Schematic representation of the complete SPR system: (a) system block diagram; (b) the SPR biosensing platform designed by rapid prototyping 3D printing; (c) electric circuit block diagram; (d) flowcell diagram.
Fig. 3. NaCl measurement curves: (a) response for different concentrations of NaCl solution; (b) △RU vs refractive index.
2.5. SPR modular part design
ADS1115 via the Inter-Integrated Circuit (IIC) communication protocol and controlled the TEC drivers by IO ports. To maintain a stable temperature, the proportional–integral–derivative (PID) controller algorithm was incorporated with ATmega328.
The SPR rapid-prototyping system was constructed with the 3Dprinted components. The Cube3 desktop version 3D printer was used to print the components with PLA filaments. The component models were designed using SolidWorks 2012 software and saved in Standard Triangle Language (STL) file format. Then, these STL files were 269
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Fig. 4. Temperature curves: (a) temperature regulation curve; (b) system temperature versus time; (c) system response against temperature change; (d) system response versus temperature.
Fig. 5. Contact angle measurement. a) Bare Au film. b) Au film after 20 min UV exposure. c) Au/MUA/MPA SAM. d) Au/MUA/MPA/aptamer film.
accessories, 3D-printed components, and a rubber sample pretreatment channel. The arrangement of the whole flow cell module is shown in Fig. 2d. The rubber flow cell of volume 65 μL and dimension 5.3 mm × 23 mm × 0.5 mm was designed and constructed using desktop cutter Silhouette Portrait. A sample pretreatment channel was also made of rubber film and cut by Silhouette Portrait. To let the sample attain a stable experimental temperature before reaching the main flow cell, the pretreatment channel was designed to be as long as possible. These components were mounted onto a stainless block,
imported into the 3D printer’s software Cubify version 3.9.0 (3Dsystems, USA). The design of the optical light path is presented in Video S1 (Supplementary material). Two rotating platforms were employed, and a 3D printed optical platform was installed onto the rotating platforms. The laser source and photodiode could be easily mounted onto the optical platform with the 3D-printed holders. Furthermore, the alignment of optical parts could be achieved effortlessly. The flow cell module consisted of a rubber flow cell, stainless 270
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Table 1 Comparison of different HbA1c measurement methods. Method
Linear range (nmol/L)
Limit of detection (nmol/L)
Antibody
Reference
Aptamer-based sandwich Electrochemistry
104 μM–313 μM
–
No
Li et al. [7]
0.3 μM–2000 μM
0.2 μM
No
Chemiluminescence
30 nM–1200 nM
–
No
Microfluidic immunoassay Potentiometric method Surface plasmon resonance
0.75 nM–22.5nM
0.37 nM
Yes
89 μM–255 μM
–
Yes
18 nM–147 nM
1 nM
No
Jain et al. [39] Ahn et al. [38] Chen et al. [40] Liu et al. [37] This work
2.6. Resolution determination
Fig. 6. SPR response of aptasensor with BSA and HbA1c.
Since the refractive index of NaCl solution could be obtained easily, different concentrations of NaCl solution (0.5%, 1%, 2%, 3% and 4%) were prepared and used as standard sample solutions to determine the resolution of the developed system. The sample delivery rate was set as 50 μL/min, and the system temperature was set as 20 °C. The system response for NaCl concentrations was recorded and used to compute the system resolution. 2.7. Temperature effect test Deionized water was used as analyte, and its resonance angle was monitored. The sample delivery speed was set as 50 μL/min. The range of temperatures during the experiment was 20 °C–28 °C. To avoid bubbles in the flow cell, the temperature of the experiment was increased and decreased gradually. 2.8. Contact angle measurement The water contact angle was measured by a Tianmin contact angle meter. The SPR sensor chip was mounted on the plane of the contact angle meter. Four microliters of deionized water was dropped onto the sensor chip surface. The droplet cross-section was recorded by the software provided by manufacturer. Contact angle was estimated from cross-sectional image using open source software ImageJ.
Fig. 7. Association curves of aptamer and HbA1c. The inset shows kapp versus concentration of HbA1c.
2.9. Kinetic parameters The kinetic parameter measurement capability of the developed system was validated using the HbA1c-specific aptamer. Affinity interactions between aptamer and HbA1c were characterized by the association rate constant ka, dissociation rate constant kd and equilibrium association constant KD. The experiment data were fitted with a 1:1 interaction model, A + B = AB, where A is the injected analyte, B is the immobilized ligand, and AB is the analyte-ligand complex. In principle, the interaction process can be described as shown in Eq. (1), where the response signal R of the SPR system is proportional to [AB], the maximum response signal Rmax is proportional to the initial [B], kapp is the apparent reaction rate constant, and t is the interaction time. From Eq. (1), we can obtain ka and kd from kapp. A series of concentrations of solution A was used to derive the relationship between kapp and [A], and a regression curve was obtained from Eq. (2). Thus ka is the slope of the regression, kd is the intercept of the regression curve, and KD is computed as ka / kd.
Fig. 8. System responses for different concentrations of HbA1c.
whose temperature was maintained stable. The assembling process of the flow cell module is described in Video S2 (Supplementary material). Refractive index matching oil was used to provide a gap between the Au-coated glass slide and the prism.
R=
271
ka [A] Rmax [1 − e−kapp t ], kapp = [A] ka + kd ([A] ka + kd )
(1)
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⎧ kapp1 = [A1 ] ka + kd ⎪k app2 = [A2 ] k a + k d ... ⎨ ⎪ kappn = [An ] ka + kd ⎩
Au film. When the MUA/MPA self-assembled monolayers were formed on Au film, the contact angle became 47.605°, which indicates the termination of the MUA/MPA self-assembled monolayer (SAM) with a hydrophilic domain, eCOOH [34]. Subsequently, after the immobilization of DNA onto the MUA/MPA SAM, the contact angle decreased to 33.030° from 46.720° of the MUA/MPA SAM. 21.911° The decrease in contact angle after DNA immobilization confirmed the successful immobilization of DNA onto the sensor surface [8].
(2)
3. Results and discussions 3.1. SPR refractive index resolution
3.4. Characterization of specificity
The strength and precision of the 3D-printed components can affect the performance of the developed system. Unlike the metal materials, the printed PLA components may deform greatly when a laser or photodiode is mounted on them. The tensile strength of 3D-printed parts was mainly influenced by the infill density. Thus, the infill density should be as high as possible. A 50% infill density is suitable for the trade-off between strength, printing time and material cost. Imprecision of the 3D-printed parts might have caused alignment problems on the optical path or assembly issues. If the mechanical adjustment of the printer was considered ideal, the precision of the printed parts would be affected by the nozzle diameters and printing layer thickness. The nozzle diameter affected the precision in the xy-plane, while the layer thickness affected the precision along the z-axis. These errors should be compensated for at the design stage. To evaluate the angular resolution of the developed system, the deionized water response was used as the background signal, and the signal noise level was 0.0008°, which can be seen as the angular resolution of our system. The system response for different concentrations of NaCl is shown in Fig. 3a. The regression equation obtained was RI = 0.00799 × [ΔRU] +1.333 (R2 = 0.994), as shown in Fig. 3b, where RU is the abbreviation of response unit (the resonance angle in degrees) and ΔRU is the angle shift of the resonance angle corresponding to the baseline signal; the sensitivity was 0.00799°/RIU. According to [25], the resolution of the refractive index can be obtained as the product of the noise level and the sensitivity. It can be inferred that the developed system exhibited a high refractive index resolution of 6.39 × 10−6 RIU.
The BSA protein contains hydrophobic, hydrophilic, cationic, and anionic regions. Therefore, the nonspecific adsorption of proteins may result from hydrogen bonding, electrostatic, charge-transfer, and/or hydrophobic interactions, depending on the surface properties. Therefore, BSA was suitable to check the specificity of the probe. To test the nonspecific binding ability of the HbA1c-specific aptamer–immobilized sensor, 149 nM BSA solution was employed as a control, and the measurement result is illustrated in Fig. 6. The resonance angle shifts of HbA1c (147 nM) and BSA (149 nM) were 0.2168 and 0.0005 RU, respectively. This result demonstrates that the HbA1c aptamer–functionalized sensor did not adsorb any BSA and hence was not specific to BSA. Thus, the aptamer-based sensors developed in this study were more specific to HbA1c. 3.5. Kinetics parameter measurement The interactions between aptamer and different concentrations of HbA1c solution are plotted in Fig. 7. A higher concentration of HbA1c showed a faster association rate. At lower concentrations of HbA1c, the association process occurred slowly. The relationship between kapp and the concentration of HbA1c is plotted as shown in the inset of Fig. 7. The linear regression equation was obtained as kapp = 5 × 103 × [C] + 0.00031 (R2 = 0.982), kapp = 0.00031 × [C ] + 0.00097 where C is the HbA1c concentration. Thus, the kinetic parameter kd was first obtained as 3.1 × 10―4 s―1 9.7 × 10−4s−1, ka was 5 × 103 M―1s―1, and Keq was 6.13 × 10―8 M, which was close to the KD value K eq ∼ 4 × 1011M−1 reported by Lin et al. [35].
3.2. Temperature effect Temperature has a significant influence on SPR measurement [32]. A one-degree change in temperature alters the refractive index resolution by ∼10―4. Fig. 4a shows the temperature regulation curve. By carefully adjusting the PID parameters (P = 2.7, I = 0.0001, D = 0.0008), the temperature controller module reached a stable state within 3 min and maintained the system temperature variation within ± 0.05 °C. Fig. 4b shows the temperature change versus time, and the system responses for different temperature are shown in Fig. 4c. The results showed that the system response decreased when the temperature increased. The system response exhibited a linear relationship with temperature between 20 °C and 28 °C (Fig. 4d), and the linear equation was ΔRU = −0.015 × T + 0.300 (R2 = 0.997). Considering the refractive index resolution computed in the previous section (Section 3.1), we can easily obtain dRI / dT = 1.3 × 10−4 / °C , which is consistent with the previously reported study [33].
3.6. HbA1c detection To estimate the detectable range for clinical diagnosis application, we designed a concentration range of HbA1c between 18 nM (1.22 μg/ mL) and 147 nM (10.00 μg/mL), which was much lower than the physiologic range (3–13 mg/ml) [36]. A calibration curve was obtained with known amounts of HbA1c. Fig. 8 depicts the SPR system responses of the aptasensor with HbA1c concentrations ranging from 18 to 147 nM. The binding interaction between HbA1c and the immobilized aptamer on the sensor surface caused the system response to increase linearly with increasing HbA1c concentration, as seen from Fig. 8. According to the of signal:noise ratio criterion, the limit of detection (LOD) calculated was 1 nM (0.070 μg/mL). On comparing the analytical performance of the proposed SPR HbA1c aptamer sensor with some previously reported [37–40] biosensors for HbA1c detection (Table 1), we found that the present sensor exhibited a low LOD and a wide detection range at nanomolar concentration. In addition, the ability of the SPR system to make measurements on very small blood samples can reduce the substrate interference and enhance the signal/noise ratio and further improve the accuracy of HbA1c detection. Thus, the developed SPR HbA1c detection system can be considered a promising method to measure the concentration of HbA1c for clinical diagnosis. However, the dilution of the actual sample by one thousandfold may not be practical for now. But this study demonstrates the potential of an HbA1c aptasensor that can using 1 μL, or even less, of actual sample for detecting HbA1c in the future.
3.3. Contact angle Prior to the biosensing experiment with the SPR system and sensing chip, the effectiveness of the modified sensor surface needed to be verified. Contact angle measurement is a powerful tool to examine every step during the sensor functionalization. The results of contact angle measurement are shown in Fig. 5. The bare Au surface exhibited a contact angle value of 76.084°. After 20 min of UV exposure, the contact angle decreased to 65.985°. This decrease in contact angle indicated that the UV exposure effectively removed the impurities on the 272
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4. Conclusion
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ChenGuang Zhang is a graduate student at University of Chinese Academy of Sciences. He received his BS degree in 2014 from Guangdong University of Technology. His current research is the design and implementation of surface plasmon resonance systems. Kalpana Settu is assistant professor at National Taipei University. She received his MS degree in 2012 and her Ph.D. in 2016 in Electrical Engineering, National Central University, her current research work has been mainly focused on biosensors, bioelectrics and biomaterials. Shwu Jen Chang is professor at I-Shou University. She received her MS degree in 1993 from Chung Yuan Christian University, and her Ph.D. in 1999 in National Yang-Ming University, Taiwan. Her current research work has been mainly focused on biosensors and biomaterials. Ching-Jung Chen is associate professor at University of Chinese Academy of Sciences. She received his BS degree in 2002 and MS degree in 2004 from I-Shou University, and her Ph.D. in 2010 in Electrical Engineering, National Central University, Taiwan. Her current research work has been mainly focused on biosensors, bioelectrics and biomaterials. Jen-Tsai Liu is associate professor at University of Chinese Academy of Sciences. He received his BS degree in 2002 and MS degree in 2004 from I-Shou University, and his Ph.D. in 2010 in Chemical and Materials Engineering, National Central University, Taiwan. His current research work has been mainly focused on biosensors and bioelectrics.
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