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Surface stress-based biosensor with stable conductive AuNPs network for biomolecules detection
Applied Surface Science 491 (2019) 443–450
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Surface stress-based biosensor with stable conductive AuNPs network for biomolecules detection
T
Dong Zhaoa,1, Yan Liua, Qiang Zhanga,1, Yixia Zhanga,b, Wendong Zhanga, Qianqian Duana, ⁎ Zhongyun Yuana, Riguang Zhangc, Shengbo Sanga, a
MicroNano System Research Center, Key Lab of Advanced Transducers and Intelligent Control System of the Ministry of Education & College of Information and Computer, Taiyuan University of Technology, Taiyuan 030024, China b Department of Biomedical Engineering, Shanxi Key Laboratory of Material Strength & Structural Impact, College of Mechanics, Taiyuan University of Technology, Taiyuan 030024, China c Key Laboratory of Coal Science and Technology of Ministry of Education and Shanxi Province, Taiyuan University of Technology, Taiyuan 030024, Shanxi, China
ARTICLE INFO
ABSTRACT
Keywords: Surface stress SC-AuNW Biosensor Glucose detection
The composite of gold and PDMS hold remarkable potential for surface stress-based biosensor applicaitons. However, the gold and PDMS in composite suffers from poor compatibility with each other and suppresses the performance of the biosensor. Here, we conduct two-step synthesis of the composite of gold and PDMS by in-situ reduction and glucose reduction of HAuCl4. The composite shows stable conductive gold nanoparticle network (SC-AuNW). In in-situ reduction, Au-seeds are embedded in PDMS surface. Furthermore, AuNPs are also generated inside the PDMS and enhance conductivity. The SC-AuNW plays a crucial role in tunneling conductivity of composite film and exhibits an excellent sensitivity to surface stress. We construct SC-AuNW-PDMS biosensors based on surface stress for glucose detection with a detection limit of 0.021 mM, high repeatability and specificity. Our research can promote the application of flexible conductive films in biosensors for disease diagnosis.
1. Introduction Biosensors have extensive applications in fields of food safety monitoring, clinical medicine, and biochemical detection. Especially, surface stress-based biosensors play an increasingly important role in detecting and analyzing the pathological state of cell [1], DNA [2], protein [3], chemical bond [4] and other biomolecules due to their rapid response time and label free. Most of the researches are focused on the surface stress-based micro-cantilever biosensors. However, despite the higher signal-to-noise ratio and sensitivity, little attention has been paid to the study of surface-based micro-membrane biosensors. [5]. Polydimethylsiloxane (PDMS) membranes have been selected as the transducer structure of surface stress-based biosensors to enhance the sensitivity in recent research due to its stability, good flexibility and biocompatibility [6,7]. Gold nanoparticles (AuNPs) was extensively employed as a sensitive element to detect biomolecules in recent years due to their desirable biosensing properties including no-toxicity to biomolecules, facile surface functionalization and good conductivity [8–10]. The AuNP-PDMS has drawn increasing attention in surface stress-based biosensor
applications [11–15]. The biosensor based on AuNP-PDMS composite was prepared by depositing AuNPs on PDMS surface and detected Creactive protein [12]. However, the AuNPs can fall off PDMS surface due to their different degrees of thermal expansion and contraction. Aptamer-based AuNP-PDMS composite was designed by silver enhancement and can visually detect biomolecules [16], which had a complicated preparation process and extended the biosensor production cycle. Moreover, the introduction of other nano metal particles has a negative impact on the sensitivity and specificity of the biosensors. Nowadays, thin film biosensors hold remarkable potential for glucose detection with glucose oxidase (GOx) and AuNPs. For example, chitosan bionanocomposite film was fabricated by electrochemical precipitation, and modified with GOx on AuNPs. The film biosensor successfully detected glucose [17,18]. Membrane cantilever biosensor also detected glucose by modifying GOx on AuNPs, and required chromium to stabilize the gold layer [19]. The chitosan and chromium layer have a negative effect on sensitivity of the film biosensor. Other work designed capacitive film biosensor. The biosensor directly modified the GOx and oxidized glucose on the gold electrode [20], leading to inaccurate test. Therefore, reliable film biosensors should be
Corresponding author. E-mail address: [email protected] (S. Sang). 1 These authors contributed equally to this work. ⁎
https://doi.org/10.1016/j.apsusc.2019.06.178 Received 9 April 2019; Received in revised form 11 June 2019; Accepted 17 June 2019 Available online 19 June 2019 0169-4332/ © 2019 Elsevier B.V. All rights reserved.
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Fig. 1. Photographic image of the PDMS with stable SC-AuNWs on both sides: (a) PDMS surface with AuNPs for glucose sensing; (b) PDMS back with AuNPs for surface stress sensing. (c) Schematic diagram the AuNP-PDMS with stable SC-AuNWs on both side. (d) The cross-section view of surface stress-based biosensor.
developed for glucose detection. Here, we prepared the gold-PDMS composite films with stable conductive gold nanoparticle network (SC-AuNW) by two-step reduction method. The morphology of SC-AuNW was characterized by the ultraviolet-visible (UV–Vis) absorption spectroscopy, scanning electron microscope (SEM) and transmission electron microscope (TEM). We have developed a surface stress-based biosensor using the SC-AuNWPDMS composite film. The biosensor was applied to detect glucose molecules based on the conductance change by modifying glucose oxidase (GOx) molecules. The biosensors capable of perceiving biomolecules has great potentials in the application of bio-detection.
solution was prepared by mixing 0.01 g/mL HAuCl4, 0.02 g/mL D-glucose, and 0.2 g/mL KHCO3 in a volume ratio of 2:1:1. The SC-AuNW was reduced and formed in the reducing solution for 4 h at room temperature [21]. 2.3. Characterization methods The SC-AuNW distribution was studied by scanning electron microscope (SEM, S-4800, Hitachi, Japan) carried out at 5 kV and transmission electron microscopy (TEM, HT7700, Hitachi, Japan) operated at an accelerating voltage of 200 kV. Absorption spectrums of SCAuNW-PDMS composite film with different incubation time were measured by UV–Visible spectrophotometer (Avantes, Beijing, China). Thickness of the films was tested by Stylus profiling system (Dektak XT, Bruke, Tucson, AZ, USA). Electrical characteristics were investigated by semiconductor analyzer (B1500A, Agilent, Santa Clara, CA, USA).
2. Methods 2.1. Materials PDMS (Sylgard 184 silicone elastomer kit) was purchased from Dow Corning (Midland, MI, USA). Chloroauric acid tetrahydrate (HAuCl4∙4H2O, Au ≥ 47.8%) was obtained from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, CHN). Potassium bicarbonate (KHCO3, ≥99.5%) was supplied by Fuchen Chemical Regents (Tianjin, China). Trimethylchlorosilane (TMCS,≥99%) was purchased from Damao Reagent (Tianjin, China). D-Glucose (C6H12O6∙H2O, ≥99.5%) and glucose oxidase (GOx, 10,000 units/g), phosphate buffered saline (PBS, pH = 7.4), 1-ethyl-3-carbodiimide (EDC), N-hydroxysulfosuccinimide (NHS), 11-Mercaptoundecanoic acid (MUA) were provided by SigmaAldrich (St. Louis, USA). Deionized water was purified using the Ulupure ultrapure water system (resistivity, 18.25 MΩ∙cm).
2.4. Density functional theory calculation method The highest occupied molecular orbital (HOMO) and the lowest unoccupied molecular orbital (LUMO) of PDMS and AuNP-PDMS were calculated by the density functional theory calculation. The energy levels were investigated as well. The double numerical plus polarization (DNP) basis sets was used in expanded electronic wave functional [22]. Generalized gradient approximation (GGA) Perdew-Burke-Ernzerhof (PBE) function was used to treat the electron exchange-correlation energy of interaction electrons [23]. All calculations were performed using Dmol3 package.
2.2. Preparation of SC-AuNW-PDMS composite films
2.5. Biosensors fabrication and functionalization
PDMS was prepared by mixing the base solution and the curing agent in a weight ratio of 10:1 (η = 0.1). After stirring uniformly, PDMS was removed air bubbles in vacuum environment. Glass substrate surfaces were hydrophobed in pure TMCS and kept for 10 min at room temperature. PDMS films were spin-coated (5000 rpm, 1 min) on glass by spin coater (SPIN-1200D, Daejeon, Korea) and cured at 70 °C for 4 h. In the first step of reduction, the films were incubated in 0.01 g/mL HAuCl4 alcohol solution for different times ranging from 3 to 48 h at room temperature. In the second step of reduction, the reducing
The native thick PDMS film (5 mm thickness) with a circular hole (Φ3mm) was used as the substrate. The hole provides space for film deformation in the process of glucose detection. The side of the composite film, reduced with SC-AuNWs, was fixed on the substrate using uncured PDMS, as the surface stress sensing layer for resistance measurement. On the other side of the film, another SC-AuNWs was reprepared by glucose reduction. Schematic diagram of the preparation process of the biosensor is shown in Fig. S1. The new SC-AuNWs provided a glucose sensing layer for GOx functionalization, as shown in 444
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Fig. 1. The 10 mM MUA alcohol solution was injected on glucose sensing layer and kept for 8 h at room temperature to form an Au-MUA self-assembly layer on biosensor surface. Then the biosensor was soaked and cleaned with ethanol and deionized water. The MUA-modified glucose sensing layer were activated with a mixture of 2 mM EDC and 5 mM NHS with a volume ratio of 1:1 for 1 h, and then, rinsed with the PBS. The activated glucose sensing layer were immersed in 400 μg/mL GOx for 2 h at room temperature. The biosensors were washed with PBS
shows uniform distribution of Au-seed, as shown in Fig. 2b. The selected area electron diffraction (SAED) indicates that the Au-seeds have been generated on PDMS surface (inset in Fig. 2b). In the second step of reduction, AuNPs are reduced by glucose (Eq. (2)). Then, gold nanoparticles can continually grow based on Au-seeds. Ultimately, SCAuNWs are formed on the PDMS surface, as shown in Fig. 2c. The SCAuNWs exhibited gold-intensive distribution. The AuNPs in SC-AuNWs had an average diameter of 124 ± 11.9 nm.
(1) buffer and dried under nitrogen. The glucose solutions in the PBS with different concentration varying from 0.05 mM to 60 mM were dripped on the biosensor and reacted for 15 min at 37 °C. The resistance changes of biosensors were monitored.
2HAuCl4 + 3HOCH2 (CHOH )4
2Au + 3HOCH2 (CHOH )4 COOH + 8
HCl + H2 O
(2)
3.2. Electrical investigation of SC-AuNW-PDMS biosensor
3. Results and discussion
A SC-AuNW-PDMS model was fabricated to understand the conductivity (Fig. 3a). In the model, conductive pathways of SC-AuNW are formed by AuNPs in PDMS. The AuNWs inside PDMS assists in enhancing conductivity. When the upper surface of the film is tensioned toward both sides, the film generates central convex deformation. As we all know, the pure PDMS is insulated; however, the SC-AuNW-PDMS exhibits electrical conductivity. Charge can transferred in the AuNWs to form current (Fig. 3b). On the other hand, the responses of SC-AuNWPDMS biosensors to central convex deformation ranging from 0 to 187.4 μm were investigated. The I-V curves in Fig. 3b shows that the current gradually weakens with the increase of deformation, indicating high-performance sensitivity of the biosensor to micro deformation.
3.1. Characterization of the SC-AuNW-PDMS biosensor The SC-AuNW-PDMS biosensors were successfully prepared by twostep reduction method. In the first step of reduction, HAuCl4 is reduced to AuNPs by silicon hydrogen groups (SieH) of PDMS. The reduction mechanism of AuNP is as follows in Eq. (1) [24]. The AuNPs, embedded on the surface of PDMS, serve as gold seed (Au-seeds), as shown in Fig. 2a. The Au-seeds play a vital role in forming the SC-AuNW-PDMS composite film. In addition, the HAuCl4 alcohol solution can penetrate into PDMS. Therefore, the AuNPs were also generated inside the PDMS, as shown in the section in Fig. 2a. TEM image of PDMS surface slice
Fig. 2. (a) SEM image showing the distribution of Au-seed on PDMS surface and section. The inset shows optical images of PDMS with Au-seed. (b) TEM image of PDMS surface slice. The inset shows the selected area electron diffraction (SAED) of the Au-seed. (c) SEM image and optical image (inset) of the SC-AuNW-PDMS. (d) Diameter distribution of AuNPs in PDMS surface after second step reduction, as measured from SEM images of samples of 200 particles using image processing using software Image J. 445
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Fig. 3. (a) The SC-AuNW-PDMS model with central convex deformation. Red line is the conductive path of AuNW in PDMS. Blue line is the conductive path of the AuNW on the surface of PDMS. (b) The response of the SC-AuNW-PDMS tunneling conductivity to different central convex deformation in the range of 0 to 187.4 μm.
Moreover, the current peak demonstrates that quantum tunneling occurs in SC-AuNW (the inset of Fig. 3b) [21], and the electronics of AuNPs can pass through a barrier between the close AuNPs to form tunneling conductive network. Hence, the biosensor can be applied to biomolecules detection based on surface stress.
the HOMO-LUMO gap of PDMS shows an energy gap of 5.967 eV (Table S1). In contract, the HOMO-LUMO gap was reduced to 4.779 eV, with the new induced gap states at −5.296 (new HOMO) and −0.417 eV (new LUMO), as shown in Fig. 4a and b. The narrowing of the gap is caused by the electron transfer from Au to the PDMS [25]. The HOMOLUMO gap of PDMS-AuNPs decreases rapidly with more AuNWs generated in PDMS (Fig. 4c and d, Table S1), leading to the interfacial interaction between PDMS and AuNPs [26,27]. This interesting result indicates that the electrons pass through the barrier in the AuNWs, and form the basis for extremely sensitive quantum tunneling conductivity
3.3. Density functional theory calculation The density functional theory calculation was conducted to understand the bandgap variations of the AuNW in PDMS. Without AuNW,
Fig. 4. DFT simulation results for (a) PDMS, (b) PDMS-Au unit, (c) PDMS-2Au and (d) PDMS-3Au. The Si, C, H, O, and Au atoms are shown in the yellow, grey, white, red, and light yellow, respectively. The PDMS-nAu indicates that PDMS chains with n number of Au atoms. 446
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Fig. 5. Scheme of the constructed SC-AuNW-PDMS biosensors based on surface stress.
with external energy. The AuNPs in the internal of PDMS indeed enhances the tunneling conductivity of SC-AuNW-PDMS films.
ranging from 3.0 to 9.0, and H2O2 concentrations from 0 to 30 mM, as shown in Fig. 6a and b. Rise in temperature causes linear increase in relative resistance change of SC-AuNW-PDMS biosensor glucose sensing layer (Fig. 6c). The result is also consistent with the sensing mechanism discussed above.
3.4. Specific modification and response strategy of SC-AuNW-PDMS biosensors The SC-AuNW-PDMS biosensors were chemically treated to form SAM on the glucose sensing layer by MUA. A carboxylated interface was generated for GOx immobilization [28], as showed in Fig. 5. The GOx is immobilized on the glucose sensing layer of the SC-AuNW-PDMS biosensors via activation of carboxyl groups by EDC/NHS. The GOx, as a protein, produces conformational changes in the enzyme reaction with glucose, causing a compressive surface stress on glucose sensing layer [29,30]. Therefore, the SC-AuNW-PDMS biosensor has a convex deformation during the enzyme reaction. The SC-AuNW is disconnected and causes the resistance increase of stress sensing layer. In addition, heat release in the enzyme reaction also generated a compressive surface stress due to the larger thermal expansion coefficient of the AuNPs than PDMS [19]. The GOx-modified SC-AuNW-PDMS biosensors can be applied to detect glucose molecules in the PBS buffer.
3.6. Glucose detection of the SC-AuNW-PDMS biosensors based on surface stress Before detecting glucose, we firstly proved that the detection signals came from the surface stress induced by glucose enzymatic hydrolysis. Hence, the control experiments were designed and conducted. The glucose (50 mM, 10 μL) and PBS (10 μL) were injected on GOx-modified and GOx-free SC-AuNW-PDMS biosensor glucose sensing layer of SCAuNW-PDMS biosensors, respectively. The dynamic responses of the biosensor are shown in Fig. 7. Due to no enzyme reaction, only gravity of PBS contributed to resistance changes of stress sensing layer, and the relative resistance changes are maintained at 10.5%. The same circumstance happens in response of the GOx-free biosensor to glucose. However, compressive surface stress is generated on glucose sensing layer in the enzymatic reaction. Gravity is gradually counteracted, leading that resistance change of stress sensing layer reduce. Until gravity is equal to surface stress and then exceeded, the resistance change of stress sensing layer increases gradually and finally reaches stability after 12 min. Hence, surface stress can be regarded as the main factor of biosensor resistance change. The responses of SC-AuNW-PDMS biosensors to glucose with different concentration were monitored by resistance change, as shown in Fig. 8. The relative resistance changes of the stress sensing layer linearly increases with the increasing glucose concentrations from 0 to 30 mM. The linear equation can be represented by ΔR/ R = 0.1032Cglucose+0.0388 (R2 = 0.9829). Ten parallel experiments indicates the high-performance stability and reproducibility of the SC-
3.5. The stability of the SC-AuNW-PDMS biosensors based on surface stress It is well known that gluconic acid and H2O2 are produced in the enzymatic hydrolysis of glucose, accompanied by heat release. Firstly, PBS with different pH and H2O2 with different concentrations were employed to demonstrate the stability of the SC-AuNW-PDMS biosensors. The PBS and H2O2 were dripped onto the glucose sensing layer of the biosensor and kept for 15 min. The resistances of stress sensing layer were measured once a minute by semiconductor analyzer. The stress sensing layer of SC-AuNW-PDMS biosensor has a relative resistance change of approximate 10.6% and remains unchanged. Therefore, the biosensors exhibits stability at different pH values 447
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Fig. 6. The relative resistance change of SC-AuNW-PDMS biosensors against pH (a), H2O2 (b) and temperatures (c).
Fig. 7. The dynamic response of the GOx-modified and GOx-free SC-AuNWPDMS biosensors to PBS and glucose with a concentration of 50 mM.
Fig. 8. The response of biosensor to the concentration of glucose solution. The inset shows biosensor micro-bending caused by surface stress after enzyme reaction.
AuNW-PDMS biosensors. On the other hand, the biosensors exhibits poor sensitivity to glucose solutions with concentration of > 30 mM. The limit of detection (LOD) of SC-AuNW-PDMS biosensors reaches 0.021 mM with 3σ/S calculation, where σ is the response standard deviation of Cglucose of the baseline, and S is the slope of linear curve [31]. We compared the performance of the biosensor with the latest reports on the detection of glucose (Table 1). Firstly, the detection limit of the proposed biosensors is higher than those reported previously for biosensors, and is just below CuO/TiO2 nanotube arrays biosensor [32,33]. In the linear range, the concentration of the glucose can be obtained directly by testing resistance. Hence, glucose detection of
Table 1 Comparison of different sensitive element materials for glucose determination.
448
Sensitive element materials
Linear range
LOD
Reference
CuO/TiO2 nanotube arrays SWCNTs Graphene patterned polyaniline Silicon nanochannels AuNP-PDMS
1–2.0 mM 1–15 mM 2.9–23 mM 0.5–8 mM 0.05–30 mM
1 μM 0.01 mM 2.9 mM 0.5 mM 0.021 mM
[32] [33] [34] [35] This work
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Author contributions The paper was written through contributions of all authors. Appendix A. Supplementary data Fig. S1 shows schematic diagram of the preparation process of the biosensor. Fig. S2 demonstrates increasing generation of Au-seeds following with longer reaction time. Fig. S3 shows the preparation process of SC-AuNW-PDMS. Fig. S4 and Fig. S5 reveals that gravity of PBS and H2O2 played the most important role in the change of biosensor resistance. Fig. S6 indicates the resistance change of the biosensor was mainly caused by surface stress and excludes the interference of temperature on resistance change. Fig. S7 shows strain of biosensor films. Table S1 shows that HOMO and LUMO for PDMS and PDMS-nAu unit. Supplementary data to this article can be found online at https://doi. org/10.1016/j.apsusc.2019.06.178. References
Fig. 9. The biosensor to other biomolecules with the concentration of 0.5 mM.
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proposed biosensors is more accurate and convenient. Especially, the proposed biosensor has both wider linear range and lower limitation of detection than reported previously for other glucose biosensors based on graphene patterned polyaniline and silicon nanochannels [34,35]. Hence, the surface stress-based membrane biosensors with tunneling conductive network an advantage in terms of linear range and LOD for biomolecule detection. 3.7. Biosensor specificity The specificity of SC-AuNW-PDMS biosensor was demonstrated by monitoring the responses to ructose, mannose and human serum albumin (HSA), each at a concentration of 0.5 mM. The biosensors exhibits a sensitive response to the glucose; however, they had an indistinctive response to other biomolecules due to their only nonspecific adsorption, as shown in Fig. 9. Hence, SC-AuNW-PDMS biosensors shows a high specificity to glucose, indicating that the biosensors can work properly in complex samples. 4. Conclusions In summary, we have reported a two-step reduction method for preparing SC-AuNW-PDMS composite film. In-situ reduction of gold has dual effects: Au-seeds and the enhancement of tunneling conductivity. The network provides tunneling conductive pathway of electrons, showing high sensitivity to micro deformation and surface stress. As a proof-of-demonstration, the SC-AuNW-PDMS composite films were employed to specifically detect glucose molecules in a linear range from 0.05 to 30 mM, with a detection limit of 0.021 mM and high-performance repeatability. In addition, the captured signals of biosensor were proven to come from surface stress generated by enzymatic reaction of glucose. The biosensor with surface stress sensitivity has a broad development potential of bioassay and can be further applied to disease diagnosis. Declaration of Competing Interest The authors declare no conflict of interest. Acknowledgements The authors are grateful for the support by the National Natural Science Foundation of China (No. 51622507, 61471255, 61474079, 61501316, 51505324), Doctoral Fund of MOE of China (No. 20131402110013), 863 Program (2015AA042601). 449
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Surface stress-based biosensor with stable conductive AuNPs network for biomolecules detection
T
Dong Zhaoa,1, Yan Liua, Qiang Zhanga,1, Yixia Zhanga,b, Wendong Zhanga, Qianqian Duana, ⁎ Zhongyun Yuana, Riguang Zhangc, Shengbo Sanga, a
MicroNano System Research Center, Key Lab of Advanced Transducers and Intelligent Control System of the Ministry of Education & College of Information and Computer, Taiyuan University of Technology, Taiyuan 030024, China b Department of Biomedical Engineering, Shanxi Key Laboratory of Material Strength & Structural Impact, College of Mechanics, Taiyuan University of Technology, Taiyuan 030024, China c Key Laboratory of Coal Science and Technology of Ministry of Education and Shanxi Province, Taiyuan University of Technology, Taiyuan 030024, Shanxi, China
ARTICLE INFO
ABSTRACT
Keywords: Surface stress SC-AuNW Biosensor Glucose detection
The composite of gold and PDMS hold remarkable potential for surface stress-based biosensor applicaitons. However, the gold and PDMS in composite suffers from poor compatibility with each other and suppresses the performance of the biosensor. Here, we conduct two-step synthesis of the composite of gold and PDMS by in-situ reduction and glucose reduction of HAuCl4. The composite shows stable conductive gold nanoparticle network (SC-AuNW). In in-situ reduction, Au-seeds are embedded in PDMS surface. Furthermore, AuNPs are also generated inside the PDMS and enhance conductivity. The SC-AuNW plays a crucial role in tunneling conductivity of composite film and exhibits an excellent sensitivity to surface stress. We construct SC-AuNW-PDMS biosensors based on surface stress for glucose detection with a detection limit of 0.021 mM, high repeatability and specificity. Our research can promote the application of flexible conductive films in biosensors for disease diagnosis.
1. Introduction Biosensors have extensive applications in fields of food safety monitoring, clinical medicine, and biochemical detection. Especially, surface stress-based biosensors play an increasingly important role in detecting and analyzing the pathological state of cell [1], DNA [2], protein [3], chemical bond [4] and other biomolecules due to their rapid response time and label free. Most of the researches are focused on the surface stress-based micro-cantilever biosensors. However, despite the higher signal-to-noise ratio and sensitivity, little attention has been paid to the study of surface-based micro-membrane biosensors. [5]. Polydimethylsiloxane (PDMS) membranes have been selected as the transducer structure of surface stress-based biosensors to enhance the sensitivity in recent research due to its stability, good flexibility and biocompatibility [6,7]. Gold nanoparticles (AuNPs) was extensively employed as a sensitive element to detect biomolecules in recent years due to their desirable biosensing properties including no-toxicity to biomolecules, facile surface functionalization and good conductivity [8–10]. The AuNP-PDMS has drawn increasing attention in surface stress-based biosensor
applications [11–15]. The biosensor based on AuNP-PDMS composite was prepared by depositing AuNPs on PDMS surface and detected Creactive protein [12]. However, the AuNPs can fall off PDMS surface due to their different degrees of thermal expansion and contraction. Aptamer-based AuNP-PDMS composite was designed by silver enhancement and can visually detect biomolecules [16], which had a complicated preparation process and extended the biosensor production cycle. Moreover, the introduction of other nano metal particles has a negative impact on the sensitivity and specificity of the biosensors. Nowadays, thin film biosensors hold remarkable potential for glucose detection with glucose oxidase (GOx) and AuNPs. For example, chitosan bionanocomposite film was fabricated by electrochemical precipitation, and modified with GOx on AuNPs. The film biosensor successfully detected glucose [17,18]. Membrane cantilever biosensor also detected glucose by modifying GOx on AuNPs, and required chromium to stabilize the gold layer [19]. The chitosan and chromium layer have a negative effect on sensitivity of the film biosensor. Other work designed capacitive film biosensor. The biosensor directly modified the GOx and oxidized glucose on the gold electrode [20], leading to inaccurate test. Therefore, reliable film biosensors should be
Corresponding author. E-mail address: [email protected] (S. Sang). 1 These authors contributed equally to this work. ⁎
https://doi.org/10.1016/j.apsusc.2019.06.178 Received 9 April 2019; Received in revised form 11 June 2019; Accepted 17 June 2019 Available online 19 June 2019 0169-4332/ © 2019 Elsevier B.V. All rights reserved.
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Fig. 1. Photographic image of the PDMS with stable SC-AuNWs on both sides: (a) PDMS surface with AuNPs for glucose sensing; (b) PDMS back with AuNPs for surface stress sensing. (c) Schematic diagram the AuNP-PDMS with stable SC-AuNWs on both side. (d) The cross-section view of surface stress-based biosensor.
developed for glucose detection. Here, we prepared the gold-PDMS composite films with stable conductive gold nanoparticle network (SC-AuNW) by two-step reduction method. The morphology of SC-AuNW was characterized by the ultraviolet-visible (UV–Vis) absorption spectroscopy, scanning electron microscope (SEM) and transmission electron microscope (TEM). We have developed a surface stress-based biosensor using the SC-AuNWPDMS composite film. The biosensor was applied to detect glucose molecules based on the conductance change by modifying glucose oxidase (GOx) molecules. The biosensors capable of perceiving biomolecules has great potentials in the application of bio-detection.
solution was prepared by mixing 0.01 g/mL HAuCl4, 0.02 g/mL D-glucose, and 0.2 g/mL KHCO3 in a volume ratio of 2:1:1. The SC-AuNW was reduced and formed in the reducing solution for 4 h at room temperature [21]. 2.3. Characterization methods The SC-AuNW distribution was studied by scanning electron microscope (SEM, S-4800, Hitachi, Japan) carried out at 5 kV and transmission electron microscopy (TEM, HT7700, Hitachi, Japan) operated at an accelerating voltage of 200 kV. Absorption spectrums of SCAuNW-PDMS composite film with different incubation time were measured by UV–Visible spectrophotometer (Avantes, Beijing, China). Thickness of the films was tested by Stylus profiling system (Dektak XT, Bruke, Tucson, AZ, USA). Electrical characteristics were investigated by semiconductor analyzer (B1500A, Agilent, Santa Clara, CA, USA).
2. Methods 2.1. Materials PDMS (Sylgard 184 silicone elastomer kit) was purchased from Dow Corning (Midland, MI, USA). Chloroauric acid tetrahydrate (HAuCl4∙4H2O, Au ≥ 47.8%) was obtained from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, CHN). Potassium bicarbonate (KHCO3, ≥99.5%) was supplied by Fuchen Chemical Regents (Tianjin, China). Trimethylchlorosilane (TMCS,≥99%) was purchased from Damao Reagent (Tianjin, China). D-Glucose (C6H12O6∙H2O, ≥99.5%) and glucose oxidase (GOx, 10,000 units/g), phosphate buffered saline (PBS, pH = 7.4), 1-ethyl-3-carbodiimide (EDC), N-hydroxysulfosuccinimide (NHS), 11-Mercaptoundecanoic acid (MUA) were provided by SigmaAldrich (St. Louis, USA). Deionized water was purified using the Ulupure ultrapure water system (resistivity, 18.25 MΩ∙cm).
2.4. Density functional theory calculation method The highest occupied molecular orbital (HOMO) and the lowest unoccupied molecular orbital (LUMO) of PDMS and AuNP-PDMS were calculated by the density functional theory calculation. The energy levels were investigated as well. The double numerical plus polarization (DNP) basis sets was used in expanded electronic wave functional [22]. Generalized gradient approximation (GGA) Perdew-Burke-Ernzerhof (PBE) function was used to treat the electron exchange-correlation energy of interaction electrons [23]. All calculations were performed using Dmol3 package.
2.2. Preparation of SC-AuNW-PDMS composite films
2.5. Biosensors fabrication and functionalization
PDMS was prepared by mixing the base solution and the curing agent in a weight ratio of 10:1 (η = 0.1). After stirring uniformly, PDMS was removed air bubbles in vacuum environment. Glass substrate surfaces were hydrophobed in pure TMCS and kept for 10 min at room temperature. PDMS films were spin-coated (5000 rpm, 1 min) on glass by spin coater (SPIN-1200D, Daejeon, Korea) and cured at 70 °C for 4 h. In the first step of reduction, the films were incubated in 0.01 g/mL HAuCl4 alcohol solution for different times ranging from 3 to 48 h at room temperature. In the second step of reduction, the reducing
The native thick PDMS film (5 mm thickness) with a circular hole (Φ3mm) was used as the substrate. The hole provides space for film deformation in the process of glucose detection. The side of the composite film, reduced with SC-AuNWs, was fixed on the substrate using uncured PDMS, as the surface stress sensing layer for resistance measurement. On the other side of the film, another SC-AuNWs was reprepared by glucose reduction. Schematic diagram of the preparation process of the biosensor is shown in Fig. S1. The new SC-AuNWs provided a glucose sensing layer for GOx functionalization, as shown in 444
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Fig. 1. The 10 mM MUA alcohol solution was injected on glucose sensing layer and kept for 8 h at room temperature to form an Au-MUA self-assembly layer on biosensor surface. Then the biosensor was soaked and cleaned with ethanol and deionized water. The MUA-modified glucose sensing layer were activated with a mixture of 2 mM EDC and 5 mM NHS with a volume ratio of 1:1 for 1 h, and then, rinsed with the PBS. The activated glucose sensing layer were immersed in 400 μg/mL GOx for 2 h at room temperature. The biosensors were washed with PBS
shows uniform distribution of Au-seed, as shown in Fig. 2b. The selected area electron diffraction (SAED) indicates that the Au-seeds have been generated on PDMS surface (inset in Fig. 2b). In the second step of reduction, AuNPs are reduced by glucose (Eq. (2)). Then, gold nanoparticles can continually grow based on Au-seeds. Ultimately, SCAuNWs are formed on the PDMS surface, as shown in Fig. 2c. The SCAuNWs exhibited gold-intensive distribution. The AuNPs in SC-AuNWs had an average diameter of 124 ± 11.9 nm.
(1) buffer and dried under nitrogen. The glucose solutions in the PBS with different concentration varying from 0.05 mM to 60 mM were dripped on the biosensor and reacted for 15 min at 37 °C. The resistance changes of biosensors were monitored.
2HAuCl4 + 3HOCH2 (CHOH )4
2Au + 3HOCH2 (CHOH )4 COOH + 8
HCl + H2 O
(2)
3.2. Electrical investigation of SC-AuNW-PDMS biosensor
3. Results and discussion
A SC-AuNW-PDMS model was fabricated to understand the conductivity (Fig. 3a). In the model, conductive pathways of SC-AuNW are formed by AuNPs in PDMS. The AuNWs inside PDMS assists in enhancing conductivity. When the upper surface of the film is tensioned toward both sides, the film generates central convex deformation. As we all know, the pure PDMS is insulated; however, the SC-AuNW-PDMS exhibits electrical conductivity. Charge can transferred in the AuNWs to form current (Fig. 3b). On the other hand, the responses of SC-AuNWPDMS biosensors to central convex deformation ranging from 0 to 187.4 μm were investigated. The I-V curves in Fig. 3b shows that the current gradually weakens with the increase of deformation, indicating high-performance sensitivity of the biosensor to micro deformation.
3.1. Characterization of the SC-AuNW-PDMS biosensor The SC-AuNW-PDMS biosensors were successfully prepared by twostep reduction method. In the first step of reduction, HAuCl4 is reduced to AuNPs by silicon hydrogen groups (SieH) of PDMS. The reduction mechanism of AuNP is as follows in Eq. (1) [24]. The AuNPs, embedded on the surface of PDMS, serve as gold seed (Au-seeds), as shown in Fig. 2a. The Au-seeds play a vital role in forming the SC-AuNW-PDMS composite film. In addition, the HAuCl4 alcohol solution can penetrate into PDMS. Therefore, the AuNPs were also generated inside the PDMS, as shown in the section in Fig. 2a. TEM image of PDMS surface slice
Fig. 2. (a) SEM image showing the distribution of Au-seed on PDMS surface and section. The inset shows optical images of PDMS with Au-seed. (b) TEM image of PDMS surface slice. The inset shows the selected area electron diffraction (SAED) of the Au-seed. (c) SEM image and optical image (inset) of the SC-AuNW-PDMS. (d) Diameter distribution of AuNPs in PDMS surface after second step reduction, as measured from SEM images of samples of 200 particles using image processing using software Image J. 445
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Fig. 3. (a) The SC-AuNW-PDMS model with central convex deformation. Red line is the conductive path of AuNW in PDMS. Blue line is the conductive path of the AuNW on the surface of PDMS. (b) The response of the SC-AuNW-PDMS tunneling conductivity to different central convex deformation in the range of 0 to 187.4 μm.
Moreover, the current peak demonstrates that quantum tunneling occurs in SC-AuNW (the inset of Fig. 3b) [21], and the electronics of AuNPs can pass through a barrier between the close AuNPs to form tunneling conductive network. Hence, the biosensor can be applied to biomolecules detection based on surface stress.
the HOMO-LUMO gap of PDMS shows an energy gap of 5.967 eV (Table S1). In contract, the HOMO-LUMO gap was reduced to 4.779 eV, with the new induced gap states at −5.296 (new HOMO) and −0.417 eV (new LUMO), as shown in Fig. 4a and b. The narrowing of the gap is caused by the electron transfer from Au to the PDMS [25]. The HOMOLUMO gap of PDMS-AuNPs decreases rapidly with more AuNWs generated in PDMS (Fig. 4c and d, Table S1), leading to the interfacial interaction between PDMS and AuNPs [26,27]. This interesting result indicates that the electrons pass through the barrier in the AuNWs, and form the basis for extremely sensitive quantum tunneling conductivity
3.3. Density functional theory calculation The density functional theory calculation was conducted to understand the bandgap variations of the AuNW in PDMS. Without AuNW,
Fig. 4. DFT simulation results for (a) PDMS, (b) PDMS-Au unit, (c) PDMS-2Au and (d) PDMS-3Au. The Si, C, H, O, and Au atoms are shown in the yellow, grey, white, red, and light yellow, respectively. The PDMS-nAu indicates that PDMS chains with n number of Au atoms. 446
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Fig. 5. Scheme of the constructed SC-AuNW-PDMS biosensors based on surface stress.
with external energy. The AuNPs in the internal of PDMS indeed enhances the tunneling conductivity of SC-AuNW-PDMS films.
ranging from 3.0 to 9.0, and H2O2 concentrations from 0 to 30 mM, as shown in Fig. 6a and b. Rise in temperature causes linear increase in relative resistance change of SC-AuNW-PDMS biosensor glucose sensing layer (Fig. 6c). The result is also consistent with the sensing mechanism discussed above.
3.4. Specific modification and response strategy of SC-AuNW-PDMS biosensors The SC-AuNW-PDMS biosensors were chemically treated to form SAM on the glucose sensing layer by MUA. A carboxylated interface was generated for GOx immobilization [28], as showed in Fig. 5. The GOx is immobilized on the glucose sensing layer of the SC-AuNW-PDMS biosensors via activation of carboxyl groups by EDC/NHS. The GOx, as a protein, produces conformational changes in the enzyme reaction with glucose, causing a compressive surface stress on glucose sensing layer [29,30]. Therefore, the SC-AuNW-PDMS biosensor has a convex deformation during the enzyme reaction. The SC-AuNW is disconnected and causes the resistance increase of stress sensing layer. In addition, heat release in the enzyme reaction also generated a compressive surface stress due to the larger thermal expansion coefficient of the AuNPs than PDMS [19]. The GOx-modified SC-AuNW-PDMS biosensors can be applied to detect glucose molecules in the PBS buffer.
3.6. Glucose detection of the SC-AuNW-PDMS biosensors based on surface stress Before detecting glucose, we firstly proved that the detection signals came from the surface stress induced by glucose enzymatic hydrolysis. Hence, the control experiments were designed and conducted. The glucose (50 mM, 10 μL) and PBS (10 μL) were injected on GOx-modified and GOx-free SC-AuNW-PDMS biosensor glucose sensing layer of SCAuNW-PDMS biosensors, respectively. The dynamic responses of the biosensor are shown in Fig. 7. Due to no enzyme reaction, only gravity of PBS contributed to resistance changes of stress sensing layer, and the relative resistance changes are maintained at 10.5%. The same circumstance happens in response of the GOx-free biosensor to glucose. However, compressive surface stress is generated on glucose sensing layer in the enzymatic reaction. Gravity is gradually counteracted, leading that resistance change of stress sensing layer reduce. Until gravity is equal to surface stress and then exceeded, the resistance change of stress sensing layer increases gradually and finally reaches stability after 12 min. Hence, surface stress can be regarded as the main factor of biosensor resistance change. The responses of SC-AuNW-PDMS biosensors to glucose with different concentration were monitored by resistance change, as shown in Fig. 8. The relative resistance changes of the stress sensing layer linearly increases with the increasing glucose concentrations from 0 to 30 mM. The linear equation can be represented by ΔR/ R = 0.1032Cglucose+0.0388 (R2 = 0.9829). Ten parallel experiments indicates the high-performance stability and reproducibility of the SC-
3.5. The stability of the SC-AuNW-PDMS biosensors based on surface stress It is well known that gluconic acid and H2O2 are produced in the enzymatic hydrolysis of glucose, accompanied by heat release. Firstly, PBS with different pH and H2O2 with different concentrations were employed to demonstrate the stability of the SC-AuNW-PDMS biosensors. The PBS and H2O2 were dripped onto the glucose sensing layer of the biosensor and kept for 15 min. The resistances of stress sensing layer were measured once a minute by semiconductor analyzer. The stress sensing layer of SC-AuNW-PDMS biosensor has a relative resistance change of approximate 10.6% and remains unchanged. Therefore, the biosensors exhibits stability at different pH values 447
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Fig. 6. The relative resistance change of SC-AuNW-PDMS biosensors against pH (a), H2O2 (b) and temperatures (c).
Fig. 7. The dynamic response of the GOx-modified and GOx-free SC-AuNWPDMS biosensors to PBS and glucose with a concentration of 50 mM.
Fig. 8. The response of biosensor to the concentration of glucose solution. The inset shows biosensor micro-bending caused by surface stress after enzyme reaction.
AuNW-PDMS biosensors. On the other hand, the biosensors exhibits poor sensitivity to glucose solutions with concentration of > 30 mM. The limit of detection (LOD) of SC-AuNW-PDMS biosensors reaches 0.021 mM with 3σ/S calculation, where σ is the response standard deviation of Cglucose of the baseline, and S is the slope of linear curve [31]. We compared the performance of the biosensor with the latest reports on the detection of glucose (Table 1). Firstly, the detection limit of the proposed biosensors is higher than those reported previously for biosensors, and is just below CuO/TiO2 nanotube arrays biosensor [32,33]. In the linear range, the concentration of the glucose can be obtained directly by testing resistance. Hence, glucose detection of
Table 1 Comparison of different sensitive element materials for glucose determination.
448
Sensitive element materials
Linear range
LOD
Reference
CuO/TiO2 nanotube arrays SWCNTs Graphene patterned polyaniline Silicon nanochannels AuNP-PDMS
1–2.0 mM 1–15 mM 2.9–23 mM 0.5–8 mM 0.05–30 mM
1 μM 0.01 mM 2.9 mM 0.5 mM 0.021 mM
[32] [33] [34] [35] This work
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Author contributions The paper was written through contributions of all authors. Appendix A. Supplementary data Fig. S1 shows schematic diagram of the preparation process of the biosensor. Fig. S2 demonstrates increasing generation of Au-seeds following with longer reaction time. Fig. S3 shows the preparation process of SC-AuNW-PDMS. Fig. S4 and Fig. S5 reveals that gravity of PBS and H2O2 played the most important role in the change of biosensor resistance. Fig. S6 indicates the resistance change of the biosensor was mainly caused by surface stress and excludes the interference of temperature on resistance change. Fig. S7 shows strain of biosensor films. Table S1 shows that HOMO and LUMO for PDMS and PDMS-nAu unit. Supplementary data to this article can be found online at https://doi. org/10.1016/j.apsusc.2019.06.178. References
Fig. 9. The biosensor to other biomolecules with the concentration of 0.5 mM.
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proposed biosensors is more accurate and convenient. Especially, the proposed biosensor has both wider linear range and lower limitation of detection than reported previously for other glucose biosensors based on graphene patterned polyaniline and silicon nanochannels [34,35]. Hence, the surface stress-based membrane biosensors with tunneling conductive network an advantage in terms of linear range and LOD for biomolecule detection. 3.7. Biosensor specificity The specificity of SC-AuNW-PDMS biosensor was demonstrated by monitoring the responses to ructose, mannose and human serum albumin (HSA), each at a concentration of 0.5 mM. The biosensors exhibits a sensitive response to the glucose; however, they had an indistinctive response to other biomolecules due to their only nonspecific adsorption, as shown in Fig. 9. Hence, SC-AuNW-PDMS biosensors shows a high specificity to glucose, indicating that the biosensors can work properly in complex samples. 4. Conclusions In summary, we have reported a two-step reduction method for preparing SC-AuNW-PDMS composite film. In-situ reduction of gold has dual effects: Au-seeds and the enhancement of tunneling conductivity. The network provides tunneling conductive pathway of electrons, showing high sensitivity to micro deformation and surface stress. As a proof-of-demonstration, the SC-AuNW-PDMS composite films were employed to specifically detect glucose molecules in a linear range from 0.05 to 30 mM, with a detection limit of 0.021 mM and high-performance repeatability. In addition, the captured signals of biosensor were proven to come from surface stress generated by enzymatic reaction of glucose. The biosensor with surface stress sensitivity has a broad development potential of bioassay and can be further applied to disease diagnosis. Declaration of Competing Interest The authors declare no conflict of interest. Acknowledgements The authors are grateful for the support by the National Natural Science Foundation of China (No. 51622507, 61471255, 61474079, 61501316, 51505324), Doctoral Fund of MOE of China (No. 20131402110013), 863 Program (2015AA042601). 449
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