- Email: [email protected]
Biosensing with SiO2-covered SPR substrates in a commercial SPR-tool
Accepted Manuscript Title: Biosensing with SiO2 -covered SPR substrates in a commercial SPR-tool Author: Jef Ryken Jiaqi Li Tim Steylaerts Rita Vos Josine Loo Karolien Jans Willem Van Roy Tim Stakenborg Pol Van Dorpe Jeroen Lammertyn Liesbet Lagae PII: DOI: Reference:
S0925-4005(14)00471-7 http://dx.doi.org/doi:10.1016/j.snb.2014.04.060 SNB 16826
To appear in:
Sensors and Actuators B
Received date: Revised date: Accepted date:
20-2-2014 8-4-2014 17-4-2014
Please cite this article as: J. Ryken, J. Li, T. Steylaerts, R. Vos, J. Loo, K. Jans, W. Van Roy, T. Stakenborg, P. Van Dorpe, J. Lammertyn, L. Lagae, Biosensing with SiO2 covered SPR substrates in a commercial SPR-tool, Sensors and Actuators B: Chemical (2014), http://dx.doi.org/10.1016/j.snb.2014.04.060 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Biosensing with SiO2-covered SPR substrates in a commercial SPR-tool
1
Imec, Kapeldreef 75, B-3001 Leuven, Belgium
2
ip t
Jef Ryken1,2,*, Jiaqi Li1,3, Tim Steylaerts1, Rita Vos1, Josine Loo1, Karolien Jans1, Willem Van Roy1, Tim Stakenborg1, Pol Van Dorpe1,3, Jeroen Lammertyn2, Liesbet Lagae1,3
Department of Biosystems (MeBioS), KU Leuven, Willem de Croylaan 42, B-3001 Leuven, Belgium
3
cr
Department of Physics and Astronomy, KU Leuven, Celestijnenlaan 200d, B-3001 Leuven, Belgium
an
us
*Corresponding author. Tel: +32 16 28 78 37. E-mail address: [email protected]
M
Highlights:
Stable SPR substrates were fabricated with standard evaporation techniques.
•
Au substrates and SiO2-covered Au substrates were compared.
•
Minimal experimental bulk sensitivity change with additional layers was observed.
•
Experimental results were confirmed by theoretical simulations.
•
Similar bio-assay results for both substrate types.
Ac ce p
te
d
•
Abstract
Surface Plasmon Resonance (SPR) is a well-established technique for studying binding kinetics and is extensively used in assay development as well as in drug discovery. Many biosensors contain an oxide surface instead of the conventional Au surface typically used in SPR sensing, which may introduce additional variables when using established protocols. Therefore, SiOxcovered SPR substrates are of great interest as a benchmarking tool for silicon-based biosensors. Moreover, SiOx has multiple advantages over Au, for instance with respect to the thermal stability of commonly used surface coupling strategies. In this paper, the bulk sensitivity of SiO2-covered Au substrates was evaluated for use in SPR. Both theoretical simulations and experimental results showed that the presence of ~10 nm of SiO2 resulted in minimal loss of bulk sensitivity compared to Au substrates. This was proven for a prostate specific antigen (PSA) recognition immuno-assay. Thus we clearly demonstrated that SiO2-covered Au
Page 1 of 15
substrates can be used for biosensing applications and do not generate significant differences compared to the original Au substrates. Keywords: Surface Plasmon Resonance, Biacore, Silicon Dioxide, Bulk Sensitivity, Biosensing
an
us
cr
ip t
1 Introduction Surface Plasmon Resonance (SPR) is a spectroscopic technique that is commonly used to study biomolecular events at the sensor surface [1, 2, 3]. A beam of p-polarized light is reflected on a conducting layer (e.g. Au) at the interface of a high refractive index glass substrate and an external medium with low refractive index. In the Kretschmann configuration, the light illuminates the conducting film from the glass slide. At a specific incident angle (θSPR), the evanescent wave, that penetrates through the metal film, excites surface plasmons, resulting in a reduction of the intensity of the reflected light (intensity dip) [4, 5, 6]. The resonance angle varies with the refractive index of the medium outside the Au film. Possible causes of refractive index changes are buffer effects and binding events occurring at the Au sensor surface. Therefore, SPR is widely used to study DNA- and RNA-hybridization [7, 8, 9], antibody-antigen interactions [10, 11, 12], enzymatic reactions [13], etc.
Ac ce p
te
d
M
SPR is commonly used as a benchmark to evaluate assays for biosensing applications, but as its use is currently limited to Au substrates, SiO2-covered Au substrates for SPR would be of great interest in the field [14, 15]. Recently, there has been a growing interest in using complementary metal-oxide semiconductor (CMOS) compatible materials for biosensor fabrication [16, 17]. SiO2 is a good candidate because of the ease and cost effectiveness of its processing [18, 19]. Moreover, it has several advantages over Au as a biosensor surface material. An important advantage is the thermal stability of the most commonly used surface biocoupling agents employed for immobilization of capture molecules. Silanes show larger thermal stability than thiols, which are typically used for Au-biofunctionalization, because of the stronger bonds with the SiOx compared to Au, respectively [20]. Different strategies, to deposit SiOx thin films on top of Au SPR-sensor chips in a simple and cost-effective manner, have been reported. The different deposition strategies applied include multitarget magnetron sputtering of Au/SiO2 composite multilayer films [21], sol-gel techniques [22] or vapor deposition approaches [23, 24, 25]. All these studies investigated the effect of the additional SiOx-layer on the SPR dip and its stability using experimental SPR set-ups. None of the above studies show the functionality of the substrates in real bio-assays. Different studies have been performed in order to increase the sensitivity of SPR tools by means of dielectric overcoatings onto the metal films. This principle uses low density dielectric coatings with thicknesses ranging from 100 nm to 750 nm and are referred to as coupled plasmon-waveguide resonance [26, 27]. These studies however cannot be performed with a commercial Biacore tool but need specific equipment. In this paper, standard evaporation and sputtering techniques were used for the preparation of the Au substrates and SiO2-covered Au substrates. The bulk sensitivities of these substrates are compared using a commercial Biacore 3000 tool. Experimentally obtained sensitivities are
Page 2 of 15
2
ip t
compared to theoretically simulated sensitivities and for the first time to our knowledge, applicability of the substrates for biosensing is demonstrated with a bio-assay to detect prostate specific antigen (PSA). Hereby, the use of a standardized SPR system, currently still one of the best and most widely used tools for bio-assay optimization and kinetics studies on Au, was extended towards dielectric coatings, strongly increasing the field of use.
Materials and Methods
d
M
an
us
cr
2.1 Materials All materials were used as received. Glycerol was obtained from Fluka analytical (St.-Gallen, Switzerland). Sodium L-ascorbate, toluene, N, N-diisopropylethylamine (DIPEA), sodium chloride (NaCl), glycine, sodium phosphate dibasic (Na2HPO4) and potassium phosphate monobasic (KH2PO4) were purchased from Sigma Aldrich (Missouri, USA). Acetone, ethanol (EtOH), hydrogen peroxide (H2O2) and sulfuric acid (H2SO4) were obtained from Honeywell (Minneapolis, USA). 11-azidoundecane-1-thiol was bought from Prochimia (Sopot, Poland). Hydrogenchloride (HCl) was purchased from OM Group (Ohio, USA). 11azidoundecyltriethoxysilane was obtained from ABCR (Karlsruhe, Germany). S1828 photoresist was purchased from micro resist technology (Berlin, Germany). Dibenzylcyclooctyne polyethylene glycol N-hydroxysuccimide ester (DBCO-PEG4-NHS) was purchased from Jena Bioscience (Jena, Germany). Sodium acetate was bought from Roth (Karlsruhe, Germany). The PSA (MW 30kDa) protein was bought from Scipac (Kent, England). The monoclonal mouse antibody against PSA (Anti-PSA) was obtained from Fujirebio Diagnostics (Pennsylvania, USA)
2.3
Ac ce p
te
2.2 Substrate preparation Both Ti and Au were deposited on a 5x5 cm quartz substrate (Schott, Germany) using thermal evaporation (CIT-Alcatel model SCM600) under vacuum (4x10-6 Pa). Ti (2 nm) was deposited at a rate of 0.1-0.2 nm/s followed by 46 nm ± 3 nm at a rate of 0.4-0.6 nm/s. For the SiO2substrates, the same Ti and Au deposition was used and additionally, 2 nm of Ti was sputtered (1x10-8 Pa, 300 W) (Pfeiffer Spider 630 tool) followed by sputtering of 10.5 nm of SiO2 (1x10-8 Pa, 250 W). Afterwards, photoresist (S1828) was deposited by spin-coating, as a protection layer, before dicing the substrates to 1x1 cm pieces. The SiO2-covered Au substrates were characterized with Transmission Electron Microscope (TEM) (Tecnai G2 F30). Multiple thickness measurements were conducted on different positions of the substrate for the different layers. Substrate cleaning and surface functionalization
2.3.1 Au substrates Prior to self-assembled monolayer (SAM) deposition, the substrates were rinsed with acetone and ethanol to remove the photoresist, followed by UV-ozone treatment for 15 min to remove all organic contaminants. Immediately after cleaning, the substrates were immersed in the 11azidoundecane-1-thiol solution (1mM thiol in EtOH). After overnight incubation, the substrates were rinsed with EtOH and dried with N2. 2.3.2 SiO2-covered Au substrates Substrate cleaning was performed as above. Afterwards, the SiO2-covered Au substrates were transferred to a controlled N2 atmosphere (glovebox) and immersed in a 11-
Page 3 of 15
azidoundecyltriethoxysilane solution (2% in toluene/DIPEA mixture). After overnight incubation, the substrates were rinsed with toluene and dried with N2. Finally, the substrates were baked for 1 h at 110 °C.
us
cr
ip t
2.4 SPR measurements SPR sensitivity measurements were performed on a Biacore 3000 system (GE Healthcare, United Kingdom). SPR reflectance dip positions are automatically converted into arbitrary response units (RU). The following approximate correspondences are often found in literature: 1 RU ≈ 0.0001° shift in SPR angle and 1 RU ≈ 10-6 RIU (refractive index units) [28]. These are however no absolute values but merely guidelines. The actual sensitivity depends on the detail of the SPR substrate, and a calibration for our specific substrates is given in the body of the paper.
M
an
2.5 Glycerol measurements Glycerol measurements were performed on a Biacore 3000 system at a temperature of 20 °C. The substrates were docked and primed twice with de-ionized water (DIW) as the running buffer. Afterwards, the freshly prepared glycerol-DIW solutions (0 %, 1 %, 2 %, 3 %, 4 %, 5 %, 10 % and 15 %) were injected at a flow rate of 20 µl/min. During the injection phase, the dips were monitored. The reported responses were monitored during injection relative to the response before injection. All sensitivity measurements were performed in triplex for both the Au substrates and the SiO2-covered Au substrates.
Ac ce p
te
d
2.6 Bio-assay The immobilization of the capture antibodies was performed in the Biacore 3000 tool at a temperature of 20 °C. After docking the substrate, a priming step with phosphate buffered saline (PBS) running buffer (150 mM NaCl, 50 mM potassium phosphate, pH 7,4) was conducted. Prior to immobilization, the thiol-functionalized Au substrates and silane-functionalized SiO2covered Au substrates were activated with an injection of 50 µl DBCO-PEG4-NHS in EtOH/DIW (1/1) at a flow rate of 5 µl/min. Afterwards, 150 µl of the anti-PSA antibody solution (250 µg/ml in 15 mM acetate buffer, pH 5.5) was injected at the same flow rate. Two short pulses of glycine/HCl (5 µl, 10 mM, pH 2.2) were used to remove non-covalently coupled antibodies from the surface. For PSA detection, the flow rate was set to 20 µl/min. Different two-fold dilutions of PSA, ranging from 0 µg/ml to 20 µg/ml (5.88x10-7 M), in 200 µl running buffer were injected. Injection of 5 µl Glycine/HCl was used to regenerate the antibody surface in between the PSA injections. The PSA binding signals were calculated relative to the baseline before each injection. The average responses and standard deviations were calculated from triplex measurements on all four flowpaths. The obtained responses were fitted using a sigmoidal fit [29] in Origin 8.1 (OriginLab, USA). The dissociation constants (Kd), the concentration where the signal reaches 50% of its saturation value, were determined to calculate the affinity constants (Ka) = 1/Kd. The limit of detection (LOD) was determined as the antigen concentration that gave a response equal to three times the standard deviation of the blank signal (0 µg/ml PSA) after baseline correction [30].
Page 4 of 15
3
an
us
cr
ip t
2.7 Simulations Two-dimensional finite-difference time-domain (2D FDTD)-based simulations were performed with FDTD Solutions (Lumerical Solutions, Vancouver, Canada). A sweep function was used, with the incidence angle of a plane wave source varying from 70° to 85°. The refractive index parameters for Au and SiO2, preloaded in the software, were used: Au from the Johnson and Christy handbook [31] and SiO2 from Palik [32]. The optical parameters, at a wavelength of 760 nm, of both Ti layers were obtained with ellipsometry (nTi1 = 2.97, kTi1 = 2.84, nTi2 = 3.19 and kTi2 = 3.15). The mesh size of the simulations was 0.5 nm. Two different layer structures for the SiO2-covered Au films were simulated. The first one corresponds to the nominal structure 2.0 nm Ti, 46.0 nm Au, 2 nm Ti, 10.5 nm SiO2, where we took the Au and SiO2 thickness as observed by TEM (see section 3.2.1), and where the Ti thicknesses were the nominal values as deposited. However, TEM suggested that the central Ti layer was ~3.5 nm thick (see section 3.2.1). Therefore, also the following structure was simulated: 2.0 nm Ti, 46 nm Au, 3.5 nm TiO2, 10.5 nm SiO2.
Results and Discussion
Ac ce p
Figure 1
te
d
M
3.1 Theoretical study Previous studies investigated the optimal thickness of the Au layer with results ranging from 42 nm [33] to 50 nm [34]. To synthesize stable SiO2-covered Au substrates, a Ti adhesion layer is required. The thickness of the Ti layer was chosen to be approximately 2 nm to minimize the decay of the SPR-signal. The SiO2 layer also needed to have a sufficient thickness to cover the complete surface but also needed to be thin enough to maintain original sensitivity, 10 nm was selected.
To understand the impact of the SiO2-surface layer on top of Au, 2D-FDTD simulations of the expected reflectance spectra and refractive index sensitivities were performed. The simulated reflection dips for the bare Au and SiO2-covered Au substrates in contact with DIW are shown in Figure 1. As expected, a significant shift in resonance angle by the addition of the SiO2 layer was observed. The shift does not only depend on the SiO2 layer, but also on the adhesion layer, with a larger shift (3.8°) for the 3.5 nm TiO2 layer than for the 2 nm Ti layer (3.0°). The width and the depth were also strongly affected by overlayer, with a much broader and less deep dip for the metallic (hence lossy) Ti adhesion layer as compared to the insulating TiO2 layer. However, the influence of the overlayers on the sensitivity is limited (Figure 2). The bulk refractive index sensitivity increased by 6% from 164.57 °/RIU for the bare Ti/Au film to 174.98 °/RIU for the substrate with 2 nm Ti / 10.5 nm SiO2 and by 25% to 205.52 °/RIU for the substrate with 3.5 nm TiO2 / 10.5 nm SiO2. This increase (rather than a decrease) may be attributed to a modified mode profile due to the presence of overlayers, and is in agreement with earlier theoretical studies [15, 35]. We also performed simulations for other adhesion layers (0 nm Ti, 0.5 nm Ti, 1 nm Ti, 1.5 nm Ti and 2 nm TiO2) (data not shown) and found sensitivities in the range 154.23 °/RIU to 182.81 °/RIU.
Page 5 of 15
Figure 2 3.2
Experimental results
us
cr
ip t
3.2.1 Substrate characterization The SiO2-covered substrates were characterized with TEM as shown in Figure 3. The thickness of the Au film was found to be 45.3 nm ± 1.3 nm. The observed thickness of the central Ti layer is 3.8 nm ± 0.2 nm for the regions where the Ti layer was clearly distinguished from the SiO2. According to the contrast, the Ti layer is possibly completely oxidized. This is also supported by the thickness of the film which corresponds to the expected thickness when a 2 nm Ti film is fully oxidized (see Appendix A). The SiO2 had a thickness of 10.4 nm ± 2.3 nm. The standard deviations on the thickness of the Au layer and the SiO2 layer indicate some thickness variations over the substrates. Additionally, the roughness of both the Au substrates and the SiO2-covered substrates was checked with atomic force microscopy (AFM) (Figure A.1). The SiO2-covered Au substrates showed a slightly higher roughness compared to the Au-substrates (RMSAu = 0.98 nm and RMSSiO2 = 1.01 nm) but the difference was minimal.
an
Figure 3
Ac ce p
te
d
M
3.2.2 Experimental bulk sensitivity The SPR-performance was evaluated experimentally on a Biacore 3000 tool. Since the Biacore software only gives angles in pixel units and angular shifts in RU (“response units”), these arbitrary units were calibrated against our FDTD simulations as follows: (1) By comparing the response (in RU) extracted by the Biacore software with the position of the dip in the raw spectra (in pixel units), a slope of 742 RU/pixel is readily obtained (Figure A.3). (2) Using different glycerol-water mixtures, with refractive index ranging from 1.3330 to 1.3509, we compared the experimental angular shifts (in pixel units) for the Ti/Au substrate with the simulated shifts (in °) for the corresponding bulk refractive index (Figure A.4), which gives a scale factor of 8.42 pixels/°. This results in 6250 RU/° or 1 RU = 0.00016°, a slight deviation from the approximate rule-of-thumb figure of 1 RU = 0.0001° which is commonly quoted in literature. (3) The horizontal offset was determined based on the dip position for the Ti/Au substrate. The shape and width of the experimental dip in intensity versus pixel position for the bare Au substrate was in good agreement with theoretical predictions (Figure 1). An SPR angle (θSPR) shift was visible after the addition of the Ti and the SiO2 layer (~ 10 pixels and ~ 26 pixels respectively). This shift was caused by the refractive index change at the Au surface by the additional SiO2 layer. The θSPR shift also induced a smaller dynamic range for the SiO2-covered substrates, as the maximal detectable angle in the tool is 60 pixel. The experimental shift is in better agreement with the theoretical shift of the substrate with 2 nm Ti as adhesion layer compared to the theoretical dip shift of the substrate with 3.5 nm TiO2 as adhesion layer. The additional layer also resulted in a wider dip (FWHMAu ~ 16.5 pixels and FWHMSiO2 ~ 20 pixels). This experimental dip width is in better agreement with the theoretical width of the substrate with 3.5 nm TiO2 as adhesion layer compared to the theoretical dip width of the substrate with 2 nm Ti as adhesion layer which is significantly broader. A consequence of the higher FWHM, was a slightly higher noise level as shown in supplementary data (Figure A.2). The noise level was
Page 6 of 15
us
cr
ip t
measured over a time span of 10 min and resulted in a root mean square (RMS) of 0.29 RU and 0.39 RU for the Au substrate and the SiO2-covered Au substrate respectively and peak-to-peak variation of 1.8 RU and 2.3 RU respectively. A change in depth of the intensity dip was also detectable. The Au substrates showed a dip that bottoms out at ~ 10000 Biacore Units (BU) while the minimum of the SiO2-substrates is ~ 13500 BU. This could be explained by the fact that the θSPR is not measured at a single position but along a section of the sensor surface. The monitored dips are an average of the measured θSPR values. The higher substrate complexity (additional Ti and SiO2 layers) could result in a more heterogeneous layer thickness, causing refractive index heterogeneity and thereby heterogeneous θSPR resulting in a more shallow dip [28]. The disagreement between experimental and theoretical results for a specific substrate containing one adhesion layer probably indicates that the experimental adhesion layer is a mixed oxide or non-stoichiometric oxide.
M
an
Different glycerol-water mixtures, with refractive index ranging from 1.3330 to 1.3509, were injected using the Biacore 3000 tool. Both the responses (Figure 2) and the raw angular spectra (Figure A.5) were recorded. Dedicated responses were plotted as a function of the refractive index (Figure 2) and the resulting sensitivity was found to be only minimally affected by the additional layers (from 1.014x106 RU/RIU to 1.049x106 RU/RIU). This is in good agreement with the simulation results that predicted only a minor change in sensitivity caused by a 10.5 nm SiO2 overlayer, irrespective of the exact composition of the thin Ti-based adhesion layer.
Ac ce p
te
d
3.2.3 Biological recognition experiments To compare their surface sensitivities, an anti-PSA/PSA-assay on both the bare Au and the SiO2-covered Au substrates was performed. The substrates were functionalized with the dedicated SAMs and an in-flow antibody (Ab) immobilization was performed in the Biacore 3000 system. Comparable signals were obtained for both substrates: 3679 ± 264 RU on the Au substrates and 3541 ± 181 RU on the SiO2-covered Au substrates. The PSA dose-response curves are shown in Figure 4. The responses for PSA are the same (~ 320 RU) at the saturation level which was expected due to the similar Ab-immobilization degree. The substrates displayed a small difference in affinity (Ka = 5.0x107 M-1 for the Au substrates and 3.7x107 M-1 for the SiO2covered Au substrates). A possible explanation for the deviation is a difference in Ab orientation or structure due to different electrostatic contributions of the underlying Au or SiO2 films [36]. Note that these differences are not related to the SPR-performance. For the bare Au and the SiO2-covered Au substrate the limit of detection of the PSA assay were 1.5x10-9 M and 2.5x10-9 M respectively. The difference can be attributed to the higher noise level of the SiO2-covered Au substrates (discussed previously), resulting in a higher standard deviation of the blank signal (3.8 RU for the bare Au substrates compared to 5.2 RU for the SiO2-covered Au substrates). The observed shift of the dose-response curve to the right, induced a slightly higher LOD. The contribution of the slightly lower sensitivity (as discussed in previous sections) was rather minimal. The standard deviations at the saturation levels were quite high. This can be partially explained by differences between channels. It was noticed that within one experiment, the differences can be as high as 25 % for the saturation values. Another possible reason for the standard deviations might be linked to some variations in the surface chemistries. Figure 4
Page 7 of 15
ip t
Overall, the anti-PSA/PSA assay on both the bare Au substrate and the SiO2-covered Au substrate show similar results. This was expected as for both substrates, the same antibody coupling chemistry was used resulting in similar coverages. Minor differences might be explained by the difference in noise, electrostatic interactions and possible irreproducibility of the surface chemistries.
an
us
cr
4 Conclusions In this manuscript we demonstrated the reproducible fabrication of Au substrates and SiO2covered Au substrates for SPR-applications with standard evaporation techniques. Both substrates were stable in aqueous solutions and could be utilized in a commercial Biacore tool. The additional SiO2 layer and adhesion Ti layer result in an angular shift due to the refractive index change near the Au film. The SiO2-covered Au substrates showed similar sensitivities to the bare Au substrates, both theoretically and experimentally. These results were further corroborated by a PSA antibody immobilization assay on both bare and SiO2-covered Au substrates which resulted in similar measured affinities for the two substrates.
M
We conclude that the SiO2-covered Au substrates can be utilized for assay optimization studies in a commercial Biacore tool, which makes them an ideal tool for benchmarking to other SiOxbased biosensing systems.
te
d
5 Vitae Jef Ryken graduated in 2010 at the University of Hasselt as a master in Biomedical Sciences while specializing in Bioelectronics and Nanotechnology. For his master’s thesis he worked on graphene-based DNA-hybridization sensors. He started his PhD at Imec in 2010 on biosensors for detecting cell secreted proteins.
Ac ce p
Jiaqi Li received his bachelor’s degree in physics at Nanjing University, and master’s in Material Engineering in Shanghai Institute of Ceramics, Chinese Academy of Sciences. For his master’s study, he focused on the preparation of semiconductor quantum dots loaded mesoporous thin films, and investigation of their nonlinear optical properties. Now he is a PhD researcher at Imec, Belgium, devoted to the 2D plasmonic cystals for the optimum biosensor performance design. Tim Steylaerts graduated in 2008 at Group T Engineering School with a Master degree in Biochemical Engineering. He has been working as a research engineer at Imec since 2007, focusing on non-contact printing of biomolecules and interactions of oligonucleotides and immunoglobulins with self-assembled monolayers, deposited on different substrate materials. He has been part of several industrial collaborations for biosensor development. Rita Vos received her Master’s degree and her PhD in Physical Chemistry at the University of Leuven in 1991 and 1995 respectively. She joined Imec in 1995 where she was working in ultraclean processing and wet benching of advanced FEOL materials and device structures. In 2011, she joined the Life Science Technology group at Imec and her current focus is on the biosensor surface preparation and site-selective coupling of bio-recognition molecules at biosensor
Page 8 of 15
substrates. She has authored and co-authored more than 20 journal papers, 7 patents and more than 90 conference proceedings contributions.
ip t
Josine Loo received her bachelor of Science in Applied Physics in 1998 and worked as process integration for Philips Research and since 2005 for Imec. This research was successively for 0.18 µm non-volatile memories, advances (sub-25 nm) CMOS devices, analog-RFCMOS for 45 nm generation and Bio-electronic systems. Since 2013 she is process engineer ebeam lithography at Imec.
us
cr
Karolien Jans graduated in 2004 at the University of Leuven as a master in Chemistry. For her work in the field of interface chemistry and surface functionalization for biosensors she obtained her PhD in 2008. In her current position of project leader at IMEC, she is involved in industrial and European research projects concerning interface chemistry for biosensors.
M
an
Willem Van Roy obtained the Master and PhD degrees in Electrical Engineering from the University of Leuven (Belgium), in collaboration with Imec (Leuven, Belgium) in 1990 and 1995, working on III-V semiconductor-based magnetoelectronics. After a postdoctoral stay at JRCAT (Tsukuba, Japan, 1995-1999) he moved back to Imec as a senior scientist responsible for semiconductor and metal spintronics. Since 2008 his interest has shifted to the lifescience area, and he is currently responsible for activities in optical and electrical biosensing, with further interests in ionics, fluidics, and medical imaging.
Ac ce p
te
d
Tim Stakenborg graduated in 1998 at the University of Leuven as a master in engineering in chemistry and biochemistry. His first research project in the field of vaccinology has resulted in a patent and was honored with a laureate prize. For his work in the field of molecular biology, he was granted a PhD in 2005 at the University of Ghent after which he joined Imec as a postdoctoral fellow. In his current position of project leader, he is involved in several European and national research projects concerning biosensors, lab on chip and nanotechnology in general. Pol van Dorpe received his PhD in electrical engineering from the University of Leuven in 2006 and after a short stay in Stanford university, he pioneered the activities on plasmonics for biosensing applications as an FWO post-doctoral fellow in Imec. Currently he is a principal scientist of Imec and a part-time professor at the physics dept of the University of Leuven. His work has led to over 60 peer reviewed papers and multiple patents. His current research interests are focused on the application of integrated photonics to miniaturize biophotonic devices. Jeroen Lammertyn is MSc in Applied Biological Sciences and MSc in Biostatistics. He obtained his PhD in Applied Biological Engineering at University of Leuven in 2001. In 2002–2003 he was research associate at the Pennsylvania State University, USA. Since October 2005 he is Full Professor at the University of Leuven and head of the Biosensor group. His main research interests involve the development of novel bio-molecular detection concepts and miniaturized analysis systems: bio-assay development (e.g., aptamers, biofunctionalized nanomaterials), optical sensors (e.g., fiber optic SPR sensors, FT-IR and NIR spectroscopy), micro- and nanofluidics (e.g., lab-on-a-chip technology).
Page 9 of 15
cr
ip t
Liesbet Lagae received her PhD from the University of Leuven, Belgium for her work on Magnetic Random Access Memories in 2003. As a postdoctoral researcher she pioneered life science technologies based on nanobiochips at Imec. Currently she works as R&D manager in the Life Science Technologies group at Imec and she built up an extensive, multidisciplinary group of more than 50 researchers working on silicon based biochips, microfluidics and biochemistry in a 200m2 state of the art biolab infrastructure. She has (co-) authored >200 peerreviewed papers in international journals and conference proceedings and holds > 15 patents in the field.
Acknowledgements
an
us
The authors thank the colleagues from IMEC for their assistance during the preparation of this manuscript. We also thank Kim Baumans for the synthesis of the substrates, David Cheyns for the ellipsometry measurements and Jef Geypen for the TEM measurements. References
M
[1] B. Liedberg, C. Nylander, I. Lundström, Biosensing with surface plasmon resonance—how it all started, Biosens. Bioelectron., 10 (1995) i-ix.
d
[2] B.-K. Oh, Y.-K. Kim, K.W. Park, W.H. Lee, J.-W. Choi, Surface plasmon resonance immunosensor for the detection of Salmonella typhimurium, Biosens. Bioelectron., 19 (2004) 1497-1504.
te
[3] X.D. Hoa, A.G. Kirk, M. Tabrizian, Towards integrated and sensitive surface plasmon resonance biosensors: a review of recent progress, Biosens. Bioelectron., 23 (2007) 151-160.
Ac ce p
[4] B. Liedberg, C. Nylander, I. Lunström, Surface plasmon resonance for gas detection and biosensing, Sens. Actuators, 4 (1983) 299-304. [5] C. Nylander, B. Liedberg, T. Lind, Gas detection by means of surface plasmon resonance, Sens. Actuators, 3 (1983) 79-88. [6] R.P.H. Kooyman, H.E. de Bruijn, R.G. Eenink, J. Greve, Surface plasmon resonance as a bioanalytical tool, J. Mol. Struct., 218 (1990) 345-350. [7] A.J. Thiel, A.G. Frutos, C.E. Jordan, R.M. Corn, L.M. Smith, Surface Plasmon Resonance Imaging measurements of DNA Hybridization Adsorption and Streptavidin/DNA Multilayer Formation at Chemically Modified Gold Surfaces, Anal. Chem., 69 (1997) 4948-4956. [8] L. He, M.D. Musick, S.R. Nicewarner, F.G. Salinas, S.J. Benkovic, M.J. Natan, C.D. Keating, R. April, Colloidal Au-Enhanced Surface Plasmon Resonance for Ultrasensitive Detection of DNA Hybridization, J. Am. Chem. Soc., 122 (2000) 9071-9077.
Page 10 of 15
[9] A. Iguchi, N. Fukuda, T. Takahashi, T. Watanabe, H. Matsuda, H. Nagase, T. Bando, H. Sugiyama, K. Shimizu, K., RNA binding properties of novel gene silencing pyrrole-imidazole polyamides, Biol. Pharm. Bull., 36 (2013) 1152-1158.
ip t
[10] M. Malmqvist, Surface plasmon resonance for detection and measurement of antibodyantigen affinity and kinetics, Curr. Opin. Immunol., 5 (1993) 282-286. [11] R. Karlsson, A. Fält, Experimental design for kinetic analysis of protein-protein interactions with surface plasmon resonance biosensors, J. of Immunol. Methods, 200 (1997) 121-133.
us
cr
[12] D.K. Weeraratne, J. Lofgren, S. Dinnogen, S.J. Swanson, Z.D. Zhong, Development of a biosensor-based immunogenicity assay capable of blocking soluble drug target interference, J. Immunol. Methods, 396 (2013) 44-55. [13] Y. Iwasaki, T. Horiuchi, O. Niwa, Detection of electrochemical enzymatic reactions by surface plasmon resonance measurement, Anal. Chem., 73 (2001) 1595-1598.
M
an
[14] E. Wijaya, C. Lenaerts, S. Maricot, J. Hastanin, S. Habraken, J.-P. Vilcot, R. Boukherroub, S. Szunerits, Surface plasmon resonance-based biosensors: From the development of different SPR structures to novel surface functionalization strategies, Curr. Opin. Solid St. M., 15 (2011) 208-224.
d
[15] S. Szunerits, A. Shalabney, R. Boukherroub, I. Abdulhalim, Dielectric coated plasmonic interfaces: their interest for sensitive sensing of analyte-ligand interactions, Rev. Anal. Chem., 31 (2012) 15-28.
te
[16] B. Jang, A. Hassibi, Biosensor systems in standard CMOS processes: fact or fiction?, Trans. Ind. Electron., 56 (2009) 979-985.
Ac ce p
[17] B.Y. Lee, S.M. Seo, D.J. Lee, M. Lee, J. Lee, J.-H. Cheon, E. Cho, H. Lee, I.-Y. Chung, Y.J. Park, S. Kim, S. Hong, Biosensor system-on-a-chip including CMOS-based signal processing circuits and 64 carbon nanotube-based sensors for the detection of a neurotransmitter, Lab Chip, 10 (2010) 894-898. [18] H. Yang, Y. Zhu, A high performance glucose biosensor enhanced via nanosized SiO2, Anal. Chim. Acta, 554 (2005) 92-97. [19] S. Krishnamoorthy, A.A. Iliadis, T. Bei, G.P. Chrousos, An interleukin-6 ZnO/SiO2/Si surface acoustic wave biosensor, Bioesens.Bioelectron., 24 (2008) 313-318. [20] A. Chandekar, S.K. Sengupta, J.E. Whitten, Thermal Stability of Thiol and Silane Monolayers: A Comparative Study, Appl. Surf. Sci., 256 (2010) 2742-2749. [21] H.B. Liao, W. Wen, G.K.L. Wong, Preparation and optical characterization of AuÕSiO2 composite films with multilayer structure 2003, J. Appl. Phys., 93, 4485.
Page 11 of 15
[22] D.K. Kambhampati, T.A.M. Jakob, J.W. Robertson, M. Cai, J.E. Pemberton, W. Knoll, Novel Silicon Dioxide Sol-Gel Films for Potential Sensors Applications: A Surface Plasmon Resonance Study, Langmuir, 17 (2001) 1169-1175.
ip t
[23] A. Granéli, J. Rydstrom, B. Kasemo, F. Höök, Formation of Supported Lipid Bilayer Membranes on SiO2 from Proteoliposomes Containing Transmembrane Proteins, Langmuir, 19 (2003) 842-850.
cr
[24] S. Szunerits, R. Boukherroub, Preparation and characterization of thin organosilicon films deposited on SPR chips, Langmuir, 22 (2006) 1660-1663.
us
[25] S. Szunerits, R. Boukherroub, Preparation of Electrochemical and Surface Plasmon Resonance Active Interfaces: Deposition of Indium Tin Oxide on Silver Thin Films, Electrochem. Commun., 8 (2006) 439-444.
an
[26] Z. Salamon, H. A. Macleod, G. Tollin, Coupled Plasmon-Waveguide Resonators: A New Spectroscopic Tool for Probing Proteolipid Film Structure and Properties, Biophys. J., 73 (1997) 2791-2797.
M
[27] N. Skiveses, R. Horvath, H. C. Pedersen, Optimization of metal-clad waveguide sensors, Sens. Actuators B, 106 (2004) 668-676. Press,
2001,
137-170
d
[28] P.A. Van der Merwe, Oxford University http://wiki.phy.queensu.ca/shughes/images/d/de/SPR_1.pdf
te
[29] A.E. Herr, D.J. Trockmorton, A.A. Davenport, A.K. Singh, On-chip native gel electrophoresis-based immunoassays for tetanus antibody and toxin, Anal. Chem., 77 (2005) 585-590.
Ac ce p
[30] J.N. Miller, J.C. Miller, Statistics and Chemometrics for Analytical Chemistry, fifth ed., Prentice Hall, London, 2000. [31] P.B. Johnson, R.W. Christy, Optical constants of noble metals, Phys. Rev. B, 6 (1972) 4704379. [32] E.D. Palik, Handbook of Optical Constants of Solids, Boston, 1985. [33] M. Suihua, L. Le, Z. Zhong, H. Yonghong, L. Zhiyi, M. Hui, G. Jihua, W. Zhangshan, Experimental study on the influence of the metal film thickness on surface plasmon resonance biosensors, Proc. of SPIE, 7192 (2009) 719211-7. [34] H.R. Gwon, S.H. Lee, Spectral and Angular Responses of Surface Plasmon Resonance Based on the Kretschmann Prism Configuration, Mater. Trans., 51 (2010) 1150-1155. [35] A. Shalabney, I. Abdulhalim, Improving the Performances of Surface Plasmon Resonance Sensor in the Infrared Region by Adding Thin Dielectric Over-Layer, 27th Convention of Electrical and Electronics Engineers in Israel, 27 (2012) 1-5.
Page 12 of 15
[36] W. Norde, Adsorption of proteins from solution at the solid-liquid interface, Advances in Colloids and Interface Science, 25 (1986) 267-340.
M
an
us
cr
ip t
Figures
Ac ce p
te
d
Figure 1: The dip positions of both the Au-substrate (black) and the SiO2-covered substrate (red). The theoretical spectra (dashed lines) are plotted on the left and bottom axes, the experimental data (full lines) on the right and top axis. The experimental and theoretical angles were calibrated as described in the text BU: Biacore Units (intensity units as given by the Biacore software). The theoretical θAu ~ 73.9 ° and θSiO2 ~ 76.8 ° for the substrate with 2 nm as adhesion layer 2 nm Ti and 77.5 ° for the substrate with 3.5 nm TiO2 as adhesion layer. The theoretical widths of the dips are FWHMAu ~ 2.47 °, FWHMSiO2 ~ 6.01 ° for the substrate with 2 nm Ti adhesion layer and 3.6 ° for the substrate with 3.5 nm TiO2 adhesion layer. The minima of the dips are 0.176 for the bare Au substrate and 0.169 and 0.029 for both SiO2-covered Au substrates. The experimental dips have a position of 10 pixel for the bare Au substrate and 26 pixel for the SiO2-covered Au substrate, a FWHM of 16.5 pixel for the bare Au substrate and 20 pixel for the SiO2-covered Au substrate, and a dip minimum of 10000 BU for the bare Au substrate and 13500 BU for the SiO2-covered Au substrate.
Page 13 of 15
ip t cr us
Ac ce p
te
d
M
an
Figure 2: Theoretical sensitivity (open symbols) and experimental responses (solid symbols) for both the Au-substrate (□,■) and SiO2-covered substrate (○,∆,●) for refractive indexes ranging from 1.3330 to 1.3509. Linear fits are shown for both experimental (solid line) and theoretical (dashed line) results. The bare Au data were used to obtain an absolute calibration of the arbitrary experimental response units as described in the text.
Figure 3: Transmission Electron Microscopy (TEM) images of the SiO2-covered SPR substrates with (A) an image showing the Au-layer, the Ti-layer and the SiO2-layer, (B) a zoom of the Tibased adhesion layer which is likely to be completely oxidized during the SiO2 deposition.
Page 14 of 15
ip t cr us
Ac ce p
te
d
M
an
Figure 4: Dose-response curves of the anti-PSA/PSA bio-assay on the bare Au substrates (squares, solid line and black error bars) and the SiO2-covered Au substrates (circles, dashed line and grey error bars) The average and standard deviations were taken from 3 different experiments with 4 different flowpaths each .
Page 15 of 15
S0925-4005(14)00471-7 http://dx.doi.org/doi:10.1016/j.snb.2014.04.060 SNB 16826
To appear in:
Sensors and Actuators B
Received date: Revised date: Accepted date:
20-2-2014 8-4-2014 17-4-2014
Please cite this article as: J. Ryken, J. Li, T. Steylaerts, R. Vos, J. Loo, K. Jans, W. Van Roy, T. Stakenborg, P. Van Dorpe, J. Lammertyn, L. Lagae, Biosensing with SiO2 covered SPR substrates in a commercial SPR-tool, Sensors and Actuators B: Chemical (2014), http://dx.doi.org/10.1016/j.snb.2014.04.060 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Biosensing with SiO2-covered SPR substrates in a commercial SPR-tool
1
Imec, Kapeldreef 75, B-3001 Leuven, Belgium
2
ip t
Jef Ryken1,2,*, Jiaqi Li1,3, Tim Steylaerts1, Rita Vos1, Josine Loo1, Karolien Jans1, Willem Van Roy1, Tim Stakenborg1, Pol Van Dorpe1,3, Jeroen Lammertyn2, Liesbet Lagae1,3
Department of Biosystems (MeBioS), KU Leuven, Willem de Croylaan 42, B-3001 Leuven, Belgium
3
cr
Department of Physics and Astronomy, KU Leuven, Celestijnenlaan 200d, B-3001 Leuven, Belgium
an
us
*Corresponding author. Tel: +32 16 28 78 37. E-mail address: [email protected]
M
Highlights:
Stable SPR substrates were fabricated with standard evaporation techniques.
•
Au substrates and SiO2-covered Au substrates were compared.
•
Minimal experimental bulk sensitivity change with additional layers was observed.
•
Experimental results were confirmed by theoretical simulations.
•
Similar bio-assay results for both substrate types.
Ac ce p
te
d
•
Abstract
Surface Plasmon Resonance (SPR) is a well-established technique for studying binding kinetics and is extensively used in assay development as well as in drug discovery. Many biosensors contain an oxide surface instead of the conventional Au surface typically used in SPR sensing, which may introduce additional variables when using established protocols. Therefore, SiOxcovered SPR substrates are of great interest as a benchmarking tool for silicon-based biosensors. Moreover, SiOx has multiple advantages over Au, for instance with respect to the thermal stability of commonly used surface coupling strategies. In this paper, the bulk sensitivity of SiO2-covered Au substrates was evaluated for use in SPR. Both theoretical simulations and experimental results showed that the presence of ~10 nm of SiO2 resulted in minimal loss of bulk sensitivity compared to Au substrates. This was proven for a prostate specific antigen (PSA) recognition immuno-assay. Thus we clearly demonstrated that SiO2-covered Au
Page 1 of 15
substrates can be used for biosensing applications and do not generate significant differences compared to the original Au substrates. Keywords: Surface Plasmon Resonance, Biacore, Silicon Dioxide, Bulk Sensitivity, Biosensing
an
us
cr
ip t
1 Introduction Surface Plasmon Resonance (SPR) is a spectroscopic technique that is commonly used to study biomolecular events at the sensor surface [1, 2, 3]. A beam of p-polarized light is reflected on a conducting layer (e.g. Au) at the interface of a high refractive index glass substrate and an external medium with low refractive index. In the Kretschmann configuration, the light illuminates the conducting film from the glass slide. At a specific incident angle (θSPR), the evanescent wave, that penetrates through the metal film, excites surface plasmons, resulting in a reduction of the intensity of the reflected light (intensity dip) [4, 5, 6]. The resonance angle varies with the refractive index of the medium outside the Au film. Possible causes of refractive index changes are buffer effects and binding events occurring at the Au sensor surface. Therefore, SPR is widely used to study DNA- and RNA-hybridization [7, 8, 9], antibody-antigen interactions [10, 11, 12], enzymatic reactions [13], etc.
Ac ce p
te
d
M
SPR is commonly used as a benchmark to evaluate assays for biosensing applications, but as its use is currently limited to Au substrates, SiO2-covered Au substrates for SPR would be of great interest in the field [14, 15]. Recently, there has been a growing interest in using complementary metal-oxide semiconductor (CMOS) compatible materials for biosensor fabrication [16, 17]. SiO2 is a good candidate because of the ease and cost effectiveness of its processing [18, 19]. Moreover, it has several advantages over Au as a biosensor surface material. An important advantage is the thermal stability of the most commonly used surface biocoupling agents employed for immobilization of capture molecules. Silanes show larger thermal stability than thiols, which are typically used for Au-biofunctionalization, because of the stronger bonds with the SiOx compared to Au, respectively [20]. Different strategies, to deposit SiOx thin films on top of Au SPR-sensor chips in a simple and cost-effective manner, have been reported. The different deposition strategies applied include multitarget magnetron sputtering of Au/SiO2 composite multilayer films [21], sol-gel techniques [22] or vapor deposition approaches [23, 24, 25]. All these studies investigated the effect of the additional SiOx-layer on the SPR dip and its stability using experimental SPR set-ups. None of the above studies show the functionality of the substrates in real bio-assays. Different studies have been performed in order to increase the sensitivity of SPR tools by means of dielectric overcoatings onto the metal films. This principle uses low density dielectric coatings with thicknesses ranging from 100 nm to 750 nm and are referred to as coupled plasmon-waveguide resonance [26, 27]. These studies however cannot be performed with a commercial Biacore tool but need specific equipment. In this paper, standard evaporation and sputtering techniques were used for the preparation of the Au substrates and SiO2-covered Au substrates. The bulk sensitivities of these substrates are compared using a commercial Biacore 3000 tool. Experimentally obtained sensitivities are
Page 2 of 15
2
ip t
compared to theoretically simulated sensitivities and for the first time to our knowledge, applicability of the substrates for biosensing is demonstrated with a bio-assay to detect prostate specific antigen (PSA). Hereby, the use of a standardized SPR system, currently still one of the best and most widely used tools for bio-assay optimization and kinetics studies on Au, was extended towards dielectric coatings, strongly increasing the field of use.
Materials and Methods
d
M
an
us
cr
2.1 Materials All materials were used as received. Glycerol was obtained from Fluka analytical (St.-Gallen, Switzerland). Sodium L-ascorbate, toluene, N, N-diisopropylethylamine (DIPEA), sodium chloride (NaCl), glycine, sodium phosphate dibasic (Na2HPO4) and potassium phosphate monobasic (KH2PO4) were purchased from Sigma Aldrich (Missouri, USA). Acetone, ethanol (EtOH), hydrogen peroxide (H2O2) and sulfuric acid (H2SO4) were obtained from Honeywell (Minneapolis, USA). 11-azidoundecane-1-thiol was bought from Prochimia (Sopot, Poland). Hydrogenchloride (HCl) was purchased from OM Group (Ohio, USA). 11azidoundecyltriethoxysilane was obtained from ABCR (Karlsruhe, Germany). S1828 photoresist was purchased from micro resist technology (Berlin, Germany). Dibenzylcyclooctyne polyethylene glycol N-hydroxysuccimide ester (DBCO-PEG4-NHS) was purchased from Jena Bioscience (Jena, Germany). Sodium acetate was bought from Roth (Karlsruhe, Germany). The PSA (MW 30kDa) protein was bought from Scipac (Kent, England). The monoclonal mouse antibody against PSA (Anti-PSA) was obtained from Fujirebio Diagnostics (Pennsylvania, USA)
2.3
Ac ce p
te
2.2 Substrate preparation Both Ti and Au were deposited on a 5x5 cm quartz substrate (Schott, Germany) using thermal evaporation (CIT-Alcatel model SCM600) under vacuum (4x10-6 Pa). Ti (2 nm) was deposited at a rate of 0.1-0.2 nm/s followed by 46 nm ± 3 nm at a rate of 0.4-0.6 nm/s. For the SiO2substrates, the same Ti and Au deposition was used and additionally, 2 nm of Ti was sputtered (1x10-8 Pa, 300 W) (Pfeiffer Spider 630 tool) followed by sputtering of 10.5 nm of SiO2 (1x10-8 Pa, 250 W). Afterwards, photoresist (S1828) was deposited by spin-coating, as a protection layer, before dicing the substrates to 1x1 cm pieces. The SiO2-covered Au substrates were characterized with Transmission Electron Microscope (TEM) (Tecnai G2 F30). Multiple thickness measurements were conducted on different positions of the substrate for the different layers. Substrate cleaning and surface functionalization
2.3.1 Au substrates Prior to self-assembled monolayer (SAM) deposition, the substrates were rinsed with acetone and ethanol to remove the photoresist, followed by UV-ozone treatment for 15 min to remove all organic contaminants. Immediately after cleaning, the substrates were immersed in the 11azidoundecane-1-thiol solution (1mM thiol in EtOH). After overnight incubation, the substrates were rinsed with EtOH and dried with N2. 2.3.2 SiO2-covered Au substrates Substrate cleaning was performed as above. Afterwards, the SiO2-covered Au substrates were transferred to a controlled N2 atmosphere (glovebox) and immersed in a 11-
Page 3 of 15
azidoundecyltriethoxysilane solution (2% in toluene/DIPEA mixture). After overnight incubation, the substrates were rinsed with toluene and dried with N2. Finally, the substrates were baked for 1 h at 110 °C.
us
cr
ip t
2.4 SPR measurements SPR sensitivity measurements were performed on a Biacore 3000 system (GE Healthcare, United Kingdom). SPR reflectance dip positions are automatically converted into arbitrary response units (RU). The following approximate correspondences are often found in literature: 1 RU ≈ 0.0001° shift in SPR angle and 1 RU ≈ 10-6 RIU (refractive index units) [28]. These are however no absolute values but merely guidelines. The actual sensitivity depends on the detail of the SPR substrate, and a calibration for our specific substrates is given in the body of the paper.
M
an
2.5 Glycerol measurements Glycerol measurements were performed on a Biacore 3000 system at a temperature of 20 °C. The substrates were docked and primed twice with de-ionized water (DIW) as the running buffer. Afterwards, the freshly prepared glycerol-DIW solutions (0 %, 1 %, 2 %, 3 %, 4 %, 5 %, 10 % and 15 %) were injected at a flow rate of 20 µl/min. During the injection phase, the dips were monitored. The reported responses were monitored during injection relative to the response before injection. All sensitivity measurements were performed in triplex for both the Au substrates and the SiO2-covered Au substrates.
Ac ce p
te
d
2.6 Bio-assay The immobilization of the capture antibodies was performed in the Biacore 3000 tool at a temperature of 20 °C. After docking the substrate, a priming step with phosphate buffered saline (PBS) running buffer (150 mM NaCl, 50 mM potassium phosphate, pH 7,4) was conducted. Prior to immobilization, the thiol-functionalized Au substrates and silane-functionalized SiO2covered Au substrates were activated with an injection of 50 µl DBCO-PEG4-NHS in EtOH/DIW (1/1) at a flow rate of 5 µl/min. Afterwards, 150 µl of the anti-PSA antibody solution (250 µg/ml in 15 mM acetate buffer, pH 5.5) was injected at the same flow rate. Two short pulses of glycine/HCl (5 µl, 10 mM, pH 2.2) were used to remove non-covalently coupled antibodies from the surface. For PSA detection, the flow rate was set to 20 µl/min. Different two-fold dilutions of PSA, ranging from 0 µg/ml to 20 µg/ml (5.88x10-7 M), in 200 µl running buffer were injected. Injection of 5 µl Glycine/HCl was used to regenerate the antibody surface in between the PSA injections. The PSA binding signals were calculated relative to the baseline before each injection. The average responses and standard deviations were calculated from triplex measurements on all four flowpaths. The obtained responses were fitted using a sigmoidal fit [29] in Origin 8.1 (OriginLab, USA). The dissociation constants (Kd), the concentration where the signal reaches 50% of its saturation value, were determined to calculate the affinity constants (Ka) = 1/Kd. The limit of detection (LOD) was determined as the antigen concentration that gave a response equal to three times the standard deviation of the blank signal (0 µg/ml PSA) after baseline correction [30].
Page 4 of 15
3
an
us
cr
ip t
2.7 Simulations Two-dimensional finite-difference time-domain (2D FDTD)-based simulations were performed with FDTD Solutions (Lumerical Solutions, Vancouver, Canada). A sweep function was used, with the incidence angle of a plane wave source varying from 70° to 85°. The refractive index parameters for Au and SiO2, preloaded in the software, were used: Au from the Johnson and Christy handbook [31] and SiO2 from Palik [32]. The optical parameters, at a wavelength of 760 nm, of both Ti layers were obtained with ellipsometry (nTi1 = 2.97, kTi1 = 2.84, nTi2 = 3.19 and kTi2 = 3.15). The mesh size of the simulations was 0.5 nm. Two different layer structures for the SiO2-covered Au films were simulated. The first one corresponds to the nominal structure 2.0 nm Ti, 46.0 nm Au, 2 nm Ti, 10.5 nm SiO2, where we took the Au and SiO2 thickness as observed by TEM (see section 3.2.1), and where the Ti thicknesses were the nominal values as deposited. However, TEM suggested that the central Ti layer was ~3.5 nm thick (see section 3.2.1). Therefore, also the following structure was simulated: 2.0 nm Ti, 46 nm Au, 3.5 nm TiO2, 10.5 nm SiO2.
Results and Discussion
Ac ce p
Figure 1
te
d
M
3.1 Theoretical study Previous studies investigated the optimal thickness of the Au layer with results ranging from 42 nm [33] to 50 nm [34]. To synthesize stable SiO2-covered Au substrates, a Ti adhesion layer is required. The thickness of the Ti layer was chosen to be approximately 2 nm to minimize the decay of the SPR-signal. The SiO2 layer also needed to have a sufficient thickness to cover the complete surface but also needed to be thin enough to maintain original sensitivity, 10 nm was selected.
To understand the impact of the SiO2-surface layer on top of Au, 2D-FDTD simulations of the expected reflectance spectra and refractive index sensitivities were performed. The simulated reflection dips for the bare Au and SiO2-covered Au substrates in contact with DIW are shown in Figure 1. As expected, a significant shift in resonance angle by the addition of the SiO2 layer was observed. The shift does not only depend on the SiO2 layer, but also on the adhesion layer, with a larger shift (3.8°) for the 3.5 nm TiO2 layer than for the 2 nm Ti layer (3.0°). The width and the depth were also strongly affected by overlayer, with a much broader and less deep dip for the metallic (hence lossy) Ti adhesion layer as compared to the insulating TiO2 layer. However, the influence of the overlayers on the sensitivity is limited (Figure 2). The bulk refractive index sensitivity increased by 6% from 164.57 °/RIU for the bare Ti/Au film to 174.98 °/RIU for the substrate with 2 nm Ti / 10.5 nm SiO2 and by 25% to 205.52 °/RIU for the substrate with 3.5 nm TiO2 / 10.5 nm SiO2. This increase (rather than a decrease) may be attributed to a modified mode profile due to the presence of overlayers, and is in agreement with earlier theoretical studies [15, 35]. We also performed simulations for other adhesion layers (0 nm Ti, 0.5 nm Ti, 1 nm Ti, 1.5 nm Ti and 2 nm TiO2) (data not shown) and found sensitivities in the range 154.23 °/RIU to 182.81 °/RIU.
Page 5 of 15
Figure 2 3.2
Experimental results
us
cr
ip t
3.2.1 Substrate characterization The SiO2-covered substrates were characterized with TEM as shown in Figure 3. The thickness of the Au film was found to be 45.3 nm ± 1.3 nm. The observed thickness of the central Ti layer is 3.8 nm ± 0.2 nm for the regions where the Ti layer was clearly distinguished from the SiO2. According to the contrast, the Ti layer is possibly completely oxidized. This is also supported by the thickness of the film which corresponds to the expected thickness when a 2 nm Ti film is fully oxidized (see Appendix A). The SiO2 had a thickness of 10.4 nm ± 2.3 nm. The standard deviations on the thickness of the Au layer and the SiO2 layer indicate some thickness variations over the substrates. Additionally, the roughness of both the Au substrates and the SiO2-covered substrates was checked with atomic force microscopy (AFM) (Figure A.1). The SiO2-covered Au substrates showed a slightly higher roughness compared to the Au-substrates (RMSAu = 0.98 nm and RMSSiO2 = 1.01 nm) but the difference was minimal.
an
Figure 3
Ac ce p
te
d
M
3.2.2 Experimental bulk sensitivity The SPR-performance was evaluated experimentally on a Biacore 3000 tool. Since the Biacore software only gives angles in pixel units and angular shifts in RU (“response units”), these arbitrary units were calibrated against our FDTD simulations as follows: (1) By comparing the response (in RU) extracted by the Biacore software with the position of the dip in the raw spectra (in pixel units), a slope of 742 RU/pixel is readily obtained (Figure A.3). (2) Using different glycerol-water mixtures, with refractive index ranging from 1.3330 to 1.3509, we compared the experimental angular shifts (in pixel units) for the Ti/Au substrate with the simulated shifts (in °) for the corresponding bulk refractive index (Figure A.4), which gives a scale factor of 8.42 pixels/°. This results in 6250 RU/° or 1 RU = 0.00016°, a slight deviation from the approximate rule-of-thumb figure of 1 RU = 0.0001° which is commonly quoted in literature. (3) The horizontal offset was determined based on the dip position for the Ti/Au substrate. The shape and width of the experimental dip in intensity versus pixel position for the bare Au substrate was in good agreement with theoretical predictions (Figure 1). An SPR angle (θSPR) shift was visible after the addition of the Ti and the SiO2 layer (~ 10 pixels and ~ 26 pixels respectively). This shift was caused by the refractive index change at the Au surface by the additional SiO2 layer. The θSPR shift also induced a smaller dynamic range for the SiO2-covered substrates, as the maximal detectable angle in the tool is 60 pixel. The experimental shift is in better agreement with the theoretical shift of the substrate with 2 nm Ti as adhesion layer compared to the theoretical dip shift of the substrate with 3.5 nm TiO2 as adhesion layer. The additional layer also resulted in a wider dip (FWHMAu ~ 16.5 pixels and FWHMSiO2 ~ 20 pixels). This experimental dip width is in better agreement with the theoretical width of the substrate with 3.5 nm TiO2 as adhesion layer compared to the theoretical dip width of the substrate with 2 nm Ti as adhesion layer which is significantly broader. A consequence of the higher FWHM, was a slightly higher noise level as shown in supplementary data (Figure A.2). The noise level was
Page 6 of 15
us
cr
ip t
measured over a time span of 10 min and resulted in a root mean square (RMS) of 0.29 RU and 0.39 RU for the Au substrate and the SiO2-covered Au substrate respectively and peak-to-peak variation of 1.8 RU and 2.3 RU respectively. A change in depth of the intensity dip was also detectable. The Au substrates showed a dip that bottoms out at ~ 10000 Biacore Units (BU) while the minimum of the SiO2-substrates is ~ 13500 BU. This could be explained by the fact that the θSPR is not measured at a single position but along a section of the sensor surface. The monitored dips are an average of the measured θSPR values. The higher substrate complexity (additional Ti and SiO2 layers) could result in a more heterogeneous layer thickness, causing refractive index heterogeneity and thereby heterogeneous θSPR resulting in a more shallow dip [28]. The disagreement between experimental and theoretical results for a specific substrate containing one adhesion layer probably indicates that the experimental adhesion layer is a mixed oxide or non-stoichiometric oxide.
M
an
Different glycerol-water mixtures, with refractive index ranging from 1.3330 to 1.3509, were injected using the Biacore 3000 tool. Both the responses (Figure 2) and the raw angular spectra (Figure A.5) were recorded. Dedicated responses were plotted as a function of the refractive index (Figure 2) and the resulting sensitivity was found to be only minimally affected by the additional layers (from 1.014x106 RU/RIU to 1.049x106 RU/RIU). This is in good agreement with the simulation results that predicted only a minor change in sensitivity caused by a 10.5 nm SiO2 overlayer, irrespective of the exact composition of the thin Ti-based adhesion layer.
Ac ce p
te
d
3.2.3 Biological recognition experiments To compare their surface sensitivities, an anti-PSA/PSA-assay on both the bare Au and the SiO2-covered Au substrates was performed. The substrates were functionalized with the dedicated SAMs and an in-flow antibody (Ab) immobilization was performed in the Biacore 3000 system. Comparable signals were obtained for both substrates: 3679 ± 264 RU on the Au substrates and 3541 ± 181 RU on the SiO2-covered Au substrates. The PSA dose-response curves are shown in Figure 4. The responses for PSA are the same (~ 320 RU) at the saturation level which was expected due to the similar Ab-immobilization degree. The substrates displayed a small difference in affinity (Ka = 5.0x107 M-1 for the Au substrates and 3.7x107 M-1 for the SiO2covered Au substrates). A possible explanation for the deviation is a difference in Ab orientation or structure due to different electrostatic contributions of the underlying Au or SiO2 films [36]. Note that these differences are not related to the SPR-performance. For the bare Au and the SiO2-covered Au substrate the limit of detection of the PSA assay were 1.5x10-9 M and 2.5x10-9 M respectively. The difference can be attributed to the higher noise level of the SiO2-covered Au substrates (discussed previously), resulting in a higher standard deviation of the blank signal (3.8 RU for the bare Au substrates compared to 5.2 RU for the SiO2-covered Au substrates). The observed shift of the dose-response curve to the right, induced a slightly higher LOD. The contribution of the slightly lower sensitivity (as discussed in previous sections) was rather minimal. The standard deviations at the saturation levels were quite high. This can be partially explained by differences between channels. It was noticed that within one experiment, the differences can be as high as 25 % for the saturation values. Another possible reason for the standard deviations might be linked to some variations in the surface chemistries. Figure 4
Page 7 of 15
ip t
Overall, the anti-PSA/PSA assay on both the bare Au substrate and the SiO2-covered Au substrate show similar results. This was expected as for both substrates, the same antibody coupling chemistry was used resulting in similar coverages. Minor differences might be explained by the difference in noise, electrostatic interactions and possible irreproducibility of the surface chemistries.
an
us
cr
4 Conclusions In this manuscript we demonstrated the reproducible fabrication of Au substrates and SiO2covered Au substrates for SPR-applications with standard evaporation techniques. Both substrates were stable in aqueous solutions and could be utilized in a commercial Biacore tool. The additional SiO2 layer and adhesion Ti layer result in an angular shift due to the refractive index change near the Au film. The SiO2-covered Au substrates showed similar sensitivities to the bare Au substrates, both theoretically and experimentally. These results were further corroborated by a PSA antibody immobilization assay on both bare and SiO2-covered Au substrates which resulted in similar measured affinities for the two substrates.
M
We conclude that the SiO2-covered Au substrates can be utilized for assay optimization studies in a commercial Biacore tool, which makes them an ideal tool for benchmarking to other SiOxbased biosensing systems.
te
d
5 Vitae Jef Ryken graduated in 2010 at the University of Hasselt as a master in Biomedical Sciences while specializing in Bioelectronics and Nanotechnology. For his master’s thesis he worked on graphene-based DNA-hybridization sensors. He started his PhD at Imec in 2010 on biosensors for detecting cell secreted proteins.
Ac ce p
Jiaqi Li received his bachelor’s degree in physics at Nanjing University, and master’s in Material Engineering in Shanghai Institute of Ceramics, Chinese Academy of Sciences. For his master’s study, he focused on the preparation of semiconductor quantum dots loaded mesoporous thin films, and investigation of their nonlinear optical properties. Now he is a PhD researcher at Imec, Belgium, devoted to the 2D plasmonic cystals for the optimum biosensor performance design. Tim Steylaerts graduated in 2008 at Group T Engineering School with a Master degree in Biochemical Engineering. He has been working as a research engineer at Imec since 2007, focusing on non-contact printing of biomolecules and interactions of oligonucleotides and immunoglobulins with self-assembled monolayers, deposited on different substrate materials. He has been part of several industrial collaborations for biosensor development. Rita Vos received her Master’s degree and her PhD in Physical Chemistry at the University of Leuven in 1991 and 1995 respectively. She joined Imec in 1995 where she was working in ultraclean processing and wet benching of advanced FEOL materials and device structures. In 2011, she joined the Life Science Technology group at Imec and her current focus is on the biosensor surface preparation and site-selective coupling of bio-recognition molecules at biosensor
Page 8 of 15
substrates. She has authored and co-authored more than 20 journal papers, 7 patents and more than 90 conference proceedings contributions.
ip t
Josine Loo received her bachelor of Science in Applied Physics in 1998 and worked as process integration for Philips Research and since 2005 for Imec. This research was successively for 0.18 µm non-volatile memories, advances (sub-25 nm) CMOS devices, analog-RFCMOS for 45 nm generation and Bio-electronic systems. Since 2013 she is process engineer ebeam lithography at Imec.
us
cr
Karolien Jans graduated in 2004 at the University of Leuven as a master in Chemistry. For her work in the field of interface chemistry and surface functionalization for biosensors she obtained her PhD in 2008. In her current position of project leader at IMEC, she is involved in industrial and European research projects concerning interface chemistry for biosensors.
M
an
Willem Van Roy obtained the Master and PhD degrees in Electrical Engineering from the University of Leuven (Belgium), in collaboration with Imec (Leuven, Belgium) in 1990 and 1995, working on III-V semiconductor-based magnetoelectronics. After a postdoctoral stay at JRCAT (Tsukuba, Japan, 1995-1999) he moved back to Imec as a senior scientist responsible for semiconductor and metal spintronics. Since 2008 his interest has shifted to the lifescience area, and he is currently responsible for activities in optical and electrical biosensing, with further interests in ionics, fluidics, and medical imaging.
Ac ce p
te
d
Tim Stakenborg graduated in 1998 at the University of Leuven as a master in engineering in chemistry and biochemistry. His first research project in the field of vaccinology has resulted in a patent and was honored with a laureate prize. For his work in the field of molecular biology, he was granted a PhD in 2005 at the University of Ghent after which he joined Imec as a postdoctoral fellow. In his current position of project leader, he is involved in several European and national research projects concerning biosensors, lab on chip and nanotechnology in general. Pol van Dorpe received his PhD in electrical engineering from the University of Leuven in 2006 and after a short stay in Stanford university, he pioneered the activities on plasmonics for biosensing applications as an FWO post-doctoral fellow in Imec. Currently he is a principal scientist of Imec and a part-time professor at the physics dept of the University of Leuven. His work has led to over 60 peer reviewed papers and multiple patents. His current research interests are focused on the application of integrated photonics to miniaturize biophotonic devices. Jeroen Lammertyn is MSc in Applied Biological Sciences and MSc in Biostatistics. He obtained his PhD in Applied Biological Engineering at University of Leuven in 2001. In 2002–2003 he was research associate at the Pennsylvania State University, USA. Since October 2005 he is Full Professor at the University of Leuven and head of the Biosensor group. His main research interests involve the development of novel bio-molecular detection concepts and miniaturized analysis systems: bio-assay development (e.g., aptamers, biofunctionalized nanomaterials), optical sensors (e.g., fiber optic SPR sensors, FT-IR and NIR spectroscopy), micro- and nanofluidics (e.g., lab-on-a-chip technology).
Page 9 of 15
cr
ip t
Liesbet Lagae received her PhD from the University of Leuven, Belgium for her work on Magnetic Random Access Memories in 2003. As a postdoctoral researcher she pioneered life science technologies based on nanobiochips at Imec. Currently she works as R&D manager in the Life Science Technologies group at Imec and she built up an extensive, multidisciplinary group of more than 50 researchers working on silicon based biochips, microfluidics and biochemistry in a 200m2 state of the art biolab infrastructure. She has (co-) authored >200 peerreviewed papers in international journals and conference proceedings and holds > 15 patents in the field.
Acknowledgements
an
us
The authors thank the colleagues from IMEC for their assistance during the preparation of this manuscript. We also thank Kim Baumans for the synthesis of the substrates, David Cheyns for the ellipsometry measurements and Jef Geypen for the TEM measurements. References
M
[1] B. Liedberg, C. Nylander, I. Lundström, Biosensing with surface plasmon resonance—how it all started, Biosens. Bioelectron., 10 (1995) i-ix.
d
[2] B.-K. Oh, Y.-K. Kim, K.W. Park, W.H. Lee, J.-W. Choi, Surface plasmon resonance immunosensor for the detection of Salmonella typhimurium, Biosens. Bioelectron., 19 (2004) 1497-1504.
te
[3] X.D. Hoa, A.G. Kirk, M. Tabrizian, Towards integrated and sensitive surface plasmon resonance biosensors: a review of recent progress, Biosens. Bioelectron., 23 (2007) 151-160.
Ac ce p
[4] B. Liedberg, C. Nylander, I. Lunström, Surface plasmon resonance for gas detection and biosensing, Sens. Actuators, 4 (1983) 299-304. [5] C. Nylander, B. Liedberg, T. Lind, Gas detection by means of surface plasmon resonance, Sens. Actuators, 3 (1983) 79-88. [6] R.P.H. Kooyman, H.E. de Bruijn, R.G. Eenink, J. Greve, Surface plasmon resonance as a bioanalytical tool, J. Mol. Struct., 218 (1990) 345-350. [7] A.J. Thiel, A.G. Frutos, C.E. Jordan, R.M. Corn, L.M. Smith, Surface Plasmon Resonance Imaging measurements of DNA Hybridization Adsorption and Streptavidin/DNA Multilayer Formation at Chemically Modified Gold Surfaces, Anal. Chem., 69 (1997) 4948-4956. [8] L. He, M.D. Musick, S.R. Nicewarner, F.G. Salinas, S.J. Benkovic, M.J. Natan, C.D. Keating, R. April, Colloidal Au-Enhanced Surface Plasmon Resonance for Ultrasensitive Detection of DNA Hybridization, J. Am. Chem. Soc., 122 (2000) 9071-9077.
Page 10 of 15
[9] A. Iguchi, N. Fukuda, T. Takahashi, T. Watanabe, H. Matsuda, H. Nagase, T. Bando, H. Sugiyama, K. Shimizu, K., RNA binding properties of novel gene silencing pyrrole-imidazole polyamides, Biol. Pharm. Bull., 36 (2013) 1152-1158.
ip t
[10] M. Malmqvist, Surface plasmon resonance for detection and measurement of antibodyantigen affinity and kinetics, Curr. Opin. Immunol., 5 (1993) 282-286. [11] R. Karlsson, A. Fält, Experimental design for kinetic analysis of protein-protein interactions with surface plasmon resonance biosensors, J. of Immunol. Methods, 200 (1997) 121-133.
us
cr
[12] D.K. Weeraratne, J. Lofgren, S. Dinnogen, S.J. Swanson, Z.D. Zhong, Development of a biosensor-based immunogenicity assay capable of blocking soluble drug target interference, J. Immunol. Methods, 396 (2013) 44-55. [13] Y. Iwasaki, T. Horiuchi, O. Niwa, Detection of electrochemical enzymatic reactions by surface plasmon resonance measurement, Anal. Chem., 73 (2001) 1595-1598.
M
an
[14] E. Wijaya, C. Lenaerts, S. Maricot, J. Hastanin, S. Habraken, J.-P. Vilcot, R. Boukherroub, S. Szunerits, Surface plasmon resonance-based biosensors: From the development of different SPR structures to novel surface functionalization strategies, Curr. Opin. Solid St. M., 15 (2011) 208-224.
d
[15] S. Szunerits, A. Shalabney, R. Boukherroub, I. Abdulhalim, Dielectric coated plasmonic interfaces: their interest for sensitive sensing of analyte-ligand interactions, Rev. Anal. Chem., 31 (2012) 15-28.
te
[16] B. Jang, A. Hassibi, Biosensor systems in standard CMOS processes: fact or fiction?, Trans. Ind. Electron., 56 (2009) 979-985.
Ac ce p
[17] B.Y. Lee, S.M. Seo, D.J. Lee, M. Lee, J. Lee, J.-H. Cheon, E. Cho, H. Lee, I.-Y. Chung, Y.J. Park, S. Kim, S. Hong, Biosensor system-on-a-chip including CMOS-based signal processing circuits and 64 carbon nanotube-based sensors for the detection of a neurotransmitter, Lab Chip, 10 (2010) 894-898. [18] H. Yang, Y. Zhu, A high performance glucose biosensor enhanced via nanosized SiO2, Anal. Chim. Acta, 554 (2005) 92-97. [19] S. Krishnamoorthy, A.A. Iliadis, T. Bei, G.P. Chrousos, An interleukin-6 ZnO/SiO2/Si surface acoustic wave biosensor, Bioesens.Bioelectron., 24 (2008) 313-318. [20] A. Chandekar, S.K. Sengupta, J.E. Whitten, Thermal Stability of Thiol and Silane Monolayers: A Comparative Study, Appl. Surf. Sci., 256 (2010) 2742-2749. [21] H.B. Liao, W. Wen, G.K.L. Wong, Preparation and optical characterization of AuÕSiO2 composite films with multilayer structure 2003, J. Appl. Phys., 93, 4485.
Page 11 of 15
[22] D.K. Kambhampati, T.A.M. Jakob, J.W. Robertson, M. Cai, J.E. Pemberton, W. Knoll, Novel Silicon Dioxide Sol-Gel Films for Potential Sensors Applications: A Surface Plasmon Resonance Study, Langmuir, 17 (2001) 1169-1175.
ip t
[23] A. Granéli, J. Rydstrom, B. Kasemo, F. Höök, Formation of Supported Lipid Bilayer Membranes on SiO2 from Proteoliposomes Containing Transmembrane Proteins, Langmuir, 19 (2003) 842-850.
cr
[24] S. Szunerits, R. Boukherroub, Preparation and characterization of thin organosilicon films deposited on SPR chips, Langmuir, 22 (2006) 1660-1663.
us
[25] S. Szunerits, R. Boukherroub, Preparation of Electrochemical and Surface Plasmon Resonance Active Interfaces: Deposition of Indium Tin Oxide on Silver Thin Films, Electrochem. Commun., 8 (2006) 439-444.
an
[26] Z. Salamon, H. A. Macleod, G. Tollin, Coupled Plasmon-Waveguide Resonators: A New Spectroscopic Tool for Probing Proteolipid Film Structure and Properties, Biophys. J., 73 (1997) 2791-2797.
M
[27] N. Skiveses, R. Horvath, H. C. Pedersen, Optimization of metal-clad waveguide sensors, Sens. Actuators B, 106 (2004) 668-676. Press,
2001,
137-170
d
[28] P.A. Van der Merwe, Oxford University http://wiki.phy.queensu.ca/shughes/images/d/de/SPR_1.pdf
te
[29] A.E. Herr, D.J. Trockmorton, A.A. Davenport, A.K. Singh, On-chip native gel electrophoresis-based immunoassays for tetanus antibody and toxin, Anal. Chem., 77 (2005) 585-590.
Ac ce p
[30] J.N. Miller, J.C. Miller, Statistics and Chemometrics for Analytical Chemistry, fifth ed., Prentice Hall, London, 2000. [31] P.B. Johnson, R.W. Christy, Optical constants of noble metals, Phys. Rev. B, 6 (1972) 4704379. [32] E.D. Palik, Handbook of Optical Constants of Solids, Boston, 1985. [33] M. Suihua, L. Le, Z. Zhong, H. Yonghong, L. Zhiyi, M. Hui, G. Jihua, W. Zhangshan, Experimental study on the influence of the metal film thickness on surface plasmon resonance biosensors, Proc. of SPIE, 7192 (2009) 719211-7. [34] H.R. Gwon, S.H. Lee, Spectral and Angular Responses of Surface Plasmon Resonance Based on the Kretschmann Prism Configuration, Mater. Trans., 51 (2010) 1150-1155. [35] A. Shalabney, I. Abdulhalim, Improving the Performances of Surface Plasmon Resonance Sensor in the Infrared Region by Adding Thin Dielectric Over-Layer, 27th Convention of Electrical and Electronics Engineers in Israel, 27 (2012) 1-5.
Page 12 of 15
[36] W. Norde, Adsorption of proteins from solution at the solid-liquid interface, Advances in Colloids and Interface Science, 25 (1986) 267-340.
M
an
us
cr
ip t
Figures
Ac ce p
te
d
Figure 1: The dip positions of both the Au-substrate (black) and the SiO2-covered substrate (red). The theoretical spectra (dashed lines) are plotted on the left and bottom axes, the experimental data (full lines) on the right and top axis. The experimental and theoretical angles were calibrated as described in the text BU: Biacore Units (intensity units as given by the Biacore software). The theoretical θAu ~ 73.9 ° and θSiO2 ~ 76.8 ° for the substrate with 2 nm as adhesion layer 2 nm Ti and 77.5 ° for the substrate with 3.5 nm TiO2 as adhesion layer. The theoretical widths of the dips are FWHMAu ~ 2.47 °, FWHMSiO2 ~ 6.01 ° for the substrate with 2 nm Ti adhesion layer and 3.6 ° for the substrate with 3.5 nm TiO2 adhesion layer. The minima of the dips are 0.176 for the bare Au substrate and 0.169 and 0.029 for both SiO2-covered Au substrates. The experimental dips have a position of 10 pixel for the bare Au substrate and 26 pixel for the SiO2-covered Au substrate, a FWHM of 16.5 pixel for the bare Au substrate and 20 pixel for the SiO2-covered Au substrate, and a dip minimum of 10000 BU for the bare Au substrate and 13500 BU for the SiO2-covered Au substrate.
Page 13 of 15
ip t cr us
Ac ce p
te
d
M
an
Figure 2: Theoretical sensitivity (open symbols) and experimental responses (solid symbols) for both the Au-substrate (□,■) and SiO2-covered substrate (○,∆,●) for refractive indexes ranging from 1.3330 to 1.3509. Linear fits are shown for both experimental (solid line) and theoretical (dashed line) results. The bare Au data were used to obtain an absolute calibration of the arbitrary experimental response units as described in the text.
Figure 3: Transmission Electron Microscopy (TEM) images of the SiO2-covered SPR substrates with (A) an image showing the Au-layer, the Ti-layer and the SiO2-layer, (B) a zoom of the Tibased adhesion layer which is likely to be completely oxidized during the SiO2 deposition.
Page 14 of 15
ip t cr us
Ac ce p
te
d
M
an
Figure 4: Dose-response curves of the anti-PSA/PSA bio-assay on the bare Au substrates (squares, solid line and black error bars) and the SiO2-covered Au substrates (circles, dashed line and grey error bars) The average and standard deviations were taken from 3 different experiments with 4 different flowpaths each .
Page 15 of 15