Superior Sensitivity of Copper-Based Plasmonic Biosensors

University of Southern Denmark Superior Sensitivity of Copper-Based Plasmonic Biosensors Stebunov, Yury V; Yakubovsky, Dmitry I; Fedyanin, Dmitry Yu; Arsenin, Aleksey V; Volkov, Valentyn S Published in: Langmuir DOI: 10.1021/acs.langmuir.8b00276 Publication date: 2018 Document version Accepted manuscript Citation for pulished version (APA): Stebunov, Y. V., Yakubovsky, D. I., Fedyanin, D. Y., Arsenin, A. V., & Volkov, V. S. (2018). Superior Sensitivity of Copper-Based Plasmonic Biosensors. Langmuir, 34(15), 4681-4687. https://doi.org/10.1021/acs.langmuir.8b00276 Terms of use This work is brought to you by the University of Southern Denmark through the SDU Research Portal. Unless otherwise specified it has been shared according to the terms for self-archiving. 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Fedyanin, Aleksey V Arsenin, and Valentyn S Volkov Langmuir, Just Accepted Manuscript • DOI: 10.1021/acs.langmuir.8b00276 • Publication Date (Web): 26 Mar 2018 Downloaded from http://pubs.acs.org on March 27, 2018 Just Accepted “Just Accepted” manuscripts have been peer-reviewed and accepted for publication. They are posted online prior to technical editing, formatting for publication and author proofing. The American Chemical Society provides “Just Accepted” as a service to the research community to expedite the dissemination of scientific material as soon as possible after acceptance. “Just Accepted” manuscripts appear in full in PDF format accompanied by an HTML abstract. “Just Accepted” manuscripts have been fully peer reviewed, but should not be considered the official version of record. They are citable by the Digital Object Identifier (DOI®). “Just Accepted” is an optional service offered to authors. 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However, no copyright claim is made to original U.S. Government works, or works produced by employees of any Commonwealth realm Crown government in the course of their duties. Page 1 of 16 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 Langmuir Superior sensitivity of copper-based plasmonic biosensors Yury V. Stebunov,*,1,2 Dmitry I. Yakubovsky,1 Dmitry Yu. Fedyanin,1 Aleksey V. Arsenin,1,2 Valentyn S. Volkov1,3 1 Laboratory of Nanooptics and Plasmonics, Moscow Institute of Physics and Technology, 9 Institutsky Lane, Dolgoprudny, 141700, Russia 2 GrapheneTek, 7 Nobel Street, Skolkovo Innovation Center, 143026, Russia 3 SDU Nano Optics, Mads Clausen Institute, University of Southern Denmark, Campusvej 55, DK75230, Odense, Denmark Plasmonic biosensing has been demonstrated to be a powerful technique for quantitative determination of molecular analytes and kinetic analysis of biochemical reactions. However, interfaces of most plasmonic biosensors are made of noble metals, such as gold and silver, which are not compatible with industrial production technologies. This greatly limits biosensing applica7 tions beyond biochemical and pharmaceutical research. Here, we propose and investigate copper7based biosensor chips fully fabri7 cated with a standard complementary metal7oxide7semiconductor (CMOS) process. The protection of thin copper films from oxida7 tion is achieved with SiO2 and Al2O3 dielectric films deposited onto the metal surface. In addition, the deposition of dielectric films with thicknesses of only several tens of nanometers significantly improves the biosensing sensitivity, owing to better localization of electromagnetic field above the biosensing surface. According to surface plasmon resonance (SPR) measurements, the copper bio7 sensor chips coated with thin films of SiO2 (25 nm) and Al2O3 (15 nm) show 55% and 75% higher sensitivity to refractive index changes, respectively, in comparison to pure gold sensor chips. To test biomolecule immobilization, the copper7dielectric biosensor chips are coated with graphene oxide linking layers and used for the selective analysis of oligonucleotide hybridization. The pro7 posed plasmonic biosensors make SPR technology more affordable for various applications and provide the basis for compact bio7 sensors integrated with modern electronic devices. Plasmonic biosensors have found numerous applications in the areas of scientific and pharmaceutical research, medical diagnostics, veterinary practice, and food and safety con7 trols.1,2,3 SPR biosensing based on the Kretschmann configura7 tion4,5 has been achieved in many commercial instruments, providing researchers with an indispensable tool for the kinetic analysis of biochemical reactions.6 In the last years, various compact plasmonic biosensors, which exploits plasmonic ef7 fects in systems comprising metal nanoparticles, nanostruc7 tured metal films, waveguides, interferometers, optical fibers, and photonic crystals, have also been demonstrated.7–14 How7 ever, the further spread of this approach into personal diagnos7 tics and other areas of in situ biosensing analysis is limited because of the costs and complexity of modern analytical de7 vices. The main obstacle behind the mass7production of plas7 monic biosensors is the obligate and specific nature of materi7 als for plasmonic structures, which should simultaneously possess plasmonic properties and be compatible with existing manufacturing technologies. Currently, electronic and photon7 ic integrated circuits are mass7produced using low7cost com7 plementary metal7oxide7semiconductor (CMOS) processes, which allow one to fabricate planar multilayered structures and form nanostructured designs. The implementation of these technologies will open up the possibilities for the integration of biosensing components into consumer electronic devises like smartphones and wearables. Metals with plasmonic properties in the visible and near7 infrared ranges include gold, silver, copper, and aluminum, as well as various metal alloys.15–29 The most common material used for plasmonic biosensors is gold, which demonstrates excellent optical properties, resistance to oxidation, and ease in nanopatterning. In addition, the optical properties of gold have been widely investigated and considered regarding the influence of multiple factors such as deposition conditions, the annealing of metal films, the presence of an adhesion layer, grain sizes, and film thicknesses.15718 The drawbacks of gold in plasmonics include its high price and incompatibility with microelectronic technological processes. Silver7based plas7 monic devices show superior performance due to low optical losses.16,19–20 However, silver components need to be protected from oxidation when used in biosensors for the detection of biological and chemical agents. An SPR biosensor based on silver films covered with protecting dielectric films demon7 strates both the stability of the conducted analyses and an in7 crease in sensitivity due to the influence of the dielectric films on the SPR properties.21723 Aluminum is one more prospective material for plasmonic biosensors and is compatible with CMOS processes and suitable for devices operating in the UV range.24–26 Unfortunately, high optical losses in most of the visible range and in IR limit its biosensing applications. In addition, other plasmonic materials have been proposed for infrared applications, such as highly doped semiconductors, transparent conducting oxides, metal nitrides and 2D materi7 als.30733 Due to their compatibility with CMOS processes, some of them could form the basis of compact plasmonic bio7 sensors. Moreover, the ability to tune their optical and struc7 tural properties offers many practical advantages for biosens7 ing applications.33 ACS Paragon Plus Environment Langmuir 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 In this work, we chose copper as a plasmonic material for designing biosensing interfaces. Besides its excellent optical properties, copper is the most common metal used in CMOS processes. Compared with gold, copper is inexpensive and has lower optical losses in the visible and NIR ranges, which were investigated by ellipsometry and SPR measurements, as well as through the analysis of surface plasmon polaritons’ propa7 gation along the surface of the copper films.16,20,34–35 The ad7 vantages of copper in plasmonic applications were exploited in ultralow7loss CMOS copper plasmonic waveguides, which could be used for the manipulation of subwavelength optical fields and signal processing.36 The biosensing applications of copper plasmonics have also been reported; for example, high7 ly sensitive plasmonic biosensors can exploit the localized SPR in copper nanoparticles and the SPR in copper films de7 posited on a photonic crystal fiber.37–38 One of the main obsta7 cles for copper7based plasmonic biosensing is the rapid oxida7 tion of the metal. A possible way to overcome this issue is to protect the underlying metal surface with a barrier coating of graphene, which at the same time produces minimal impact on the optical quality of the interface between the sensor surface and the chemical or biological systems to be studied.20 Here, we propose SPR sensor chips based on plasmonic copper films covered with different dielectric layers (Fig. 1). Thin copper films were deposited on the surfaces of glass sub7 strates by electron beam evaporation, which is an essential part of standard CMOS processing. The additional deposition of dielectric films in the same fabrication cycle prevents the oxi7 dation of the copper and additionally increases its biosensing sensitivity. The performance of the proposed sensor chips was investigated both theoretically and experimentally, with the aim of optimizing the multilayered sensor chip configuration. The analysis of the angular spectra of the light reflectance from the proposed sensor chips provides quantitative data on both the SPR characteristics and the dependence of the sensi7 tivity on the material properties of the protecting films and their thicknesses. Experimentally, the sensitivity of the pro7 posed sensor chips was derived from the SPR signal response upon the injections of different salt solutions with predeter7 mined refractive indices. In addition, the performance of the dielectric7coated sensor chips was validated in the biosensing analysis of oligonucleotide hybridization. For this analysis, the sensor chips were coated with graphene oxide (GO) linking layers and thereafter immobilized with neutravidin protein, which is selective towards biotinylated ligands. The GO sub7 strates provide improved biosensing sensitivity compared to conventional hydrogel7based linking layers.39 In addition, they can be deposited on various dielectric and metal surfaces using the same procedure. EXPERIMENTAL SECTION Unless otherwise stated, all chemicals were pur7 chased from Sigma7Aldrich (Carlsbad, CA). Materials for evaporation, including copper, titanium, silicon dioxide and aluminum oxide, were purchased from the Kurt J. Lesker Company (Hastings, UK). The glass substrates used were D 263® T Eco7Friendly Thin Glass (SCHOTT AG, Mainz, Germany) and had the dimensions of 14x12x0.4 mm. An aqueous solution of GO with a concentration of 500 Gg/mL was purchased from Graphene Laboratories, Inc. (NY, USA) and synthesized by the Hummers method.40 The following oligonucleotides were used: 1) biotinylated 56 bp single7 stranded DNA sequence (D1) (5′7/5Biosg/ TCT CTC TGA Page 2 of 16 GTG GCC AAA ATT TCA TCT CTG AAT TCA GGG ATG ATG ATA ACA AAT GC73′) and 2) the 50 bp single7stranded DNA sequence (D2) (5′7GCA TTT GTT ATC ATC ATC CCT GAA TTC AGA GAT GAA ATT TTG GCC ACT CA73′). The oligonucleotides were synthesized by Integrated DNA Technologies (Coralville, USA). Neutravidin was purchased from Thermo Fisher Scientific (Waltham, MA). All solutions were prepared in ultrapure water (18.3 MOhm cm). Figure 1. Schematic representation of the SPR biosensor comprising the SPR sensor chip based on plasmonic copper films coated with a dielectric layer to protect against oxida7 tion. The prism and sensor chip substrates are made of the same type of glass, which allows for an efficient optical con7 nection. The immobilization of biomolecules on the biosensor surface can be achieved using a graphene oxide linking layer deposited atop the dielectric layer. SPR sensor chips were fabricated using a NEE74000 E7Beam Evaporating system produced by NANO7MASTER, Inc. (Austin, TX). Clean SCHOTT’s glass substrates were placed in the vacuum cham7 ber of the e7beam evaporator at a pressure of 3·1076 Torr. After that, the following films were deposited in a single process: 1) A 1.57nm thick Ti layer to ensure the adhesion of the copper to the substrate, 2) a 277nm7thick copper film, and 3) SiO2 or Al2O3 protecting layers of different thicknesses (2.5 nm 7 35 nm). The deposition rate was approximately 2 A/s. Additional7 ly, the thicknesses of the deposited films were confirmed by AFM measurements using an AFM Ntegra Aura produced by NT7MDT (Moscow, Russia). Finally, 57nm7thick graphene oxide linking layers were spray7coated onto the surfaces of the copper SPR chips to perform a biosensing assay.39 Ellipsometric measurements were conducted using the VASE Ellipsometer produced by J.A. Woollam Co. (Lincoln, NE). The spectral range was 3007 1500 nm, and the angles of incidence of the light beam were 60, 65 70 and 75 degrees. For the ellipsometric measurements, copper and gold films with thicknesses of 26 and 25 nm, re7 spectively, were deposited using e7beam evaporation on the surfaces of silicon wafers, both capped with a top layer of 27 nm7thick SiO2. The thickness of the metal films was measured by AFM, which allowed for the direct determination of the complex permittivity of the single7layer materials from ellip7 sometry data. Due to the partial oxidation of the copper, the multilayered structure with the copper film (used for ellipsom7 etry analysis) consisted of the layers of copper and copper (II) oxide, with thicknesses of 25.5 and 0.5 nm, respectively. The ACS Paragon Plus Environment Page 3 of 16 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 Langmuir permittivities of copper and gold in the near7infrared (8007 1500 nm) region were approximated by the Drude model:18 ε = ε∞ − ωp2 ω 2 + Γω (1) , where the dielectric function at an infinite frequency ε ∞ , the plasma frequency ωp , and the scattering rate Γ are fitting parameters. For the theoretical and experi7 mental analysis of copper biosensor chips, we reason that the SPR is excited in metal films by means of the Kretschman’s geometry, which is used in most of the commercial SPR in7 struments.4 In resonance, the reflection of p7polarized mono7 chromatic radiation is at its minimum, which results from the direct coupling into the surface plasmon polariton modes that travel along a metal7dielectric interface. Therefore, the pa7 rameter characterizing the biosensing performance is the sen7 sitivity to refractive index (RI) changes: RI = , (2) is the change in biosensor signal induced by the where .41 change of RI near the biosensor chip Theoretical descriptions of SPR excitation are based on the transfer matrix method (TMM), which uses Fresnel’s coeffi7 cients to describe the light reflection from the multilayered structure.42 Using TMM, we investigated the SPR excitation in the copper biosensor chips at 6357nm wavelength ( λ = 635 nm ), which was also used in the SPR experiments. The multilayered structure with the following layers was used for the simulations: 1) semi7infinite glass substrate with an RI 0 = 1.521 , 2) Ti layer with a thickness 1 = 1.5 nm and RI 1 = 2.709 + 3.771 , 3) Cu layer with a thickness and 2 2 = 50 nm = 0.0549 + 4.406 , 4) different protecting layers of varying thickness made of silicon ( silicon carbide ( con nitride ( SiN silicon oxide ( SiC Si = 3.879 + 0.0192 ), = 2.634 ), zinc sulfide ( = 2.01 ), aluminum oxide ( SiO2 ZnS = 2.35 ), sili7 Al2 O3 = 1.766 ), or = 1.457 ), and 5) semi7infinite aqueous solution with RI 4 = 1.33 . The minimum angular reflectance from the above7described structures corresponds to the SPR excitation. The change of the RI of the top aqueous solution layer by n leads to the shift of the SPR angle, which was de7 fined as the change of biosensor signal Pin equation (2), which was the calculations of the sensitivity SRI for the particulate biosensor chip. In the calculations, n was taken as 0.005. The experimental investigation of SPR sensitivity to RI changes was performed using a commercial SPR instrument, the Accolade 404SA produced by the company BiOptix (Boulder, USA). According to the manufacturer, this instru7 ment can detect RI changes of less than 0.5 ⋅10−6 . During the procedure, the SPR chips (fabricated by e7beam evaporation as described above) were inserted into the instrument and conse7 quently rinsed by two solutions: 0.17M phosphate7buffered saline (PBS) with a pH of 7.4 and 0.5% NaCl solution in the same PBS buffer. The difference between the RIs of these = 0.88 ⋅10−3 , which resulted in a solutions is approximately difference in the absolute SPR signals (measured in Volts).43 Obtained sensorgrams were corrected for linear drift caused by interactions between the sensor chips and the salt solutions. SPR biosensing measurements were conducted using copper7based biosensor chips comprising a 277nm7thick copper film coated with a 157nm7thick Al2O3 protecting layer, which were fabricated according to the above7described procedure. For biomolecule immobilization, a GO linking layer with a thickness of 5 nm was deposited on the surface of a copper SPR chip by spraying it with 1 mL of the GO aqueous solution with a concentration of 25 Gg/mL. Afterwards, neutravidin protein was immobilized on the sur7 face of the GO layer directly in the flow cell of the SPR in7 strument using amine coupling.6 For neutravidin immobiliza7 tion, the carboxyl groups of the GO sensor chip were activated with a mixture of 0.47M 17ethyl737(37(dimethylamino)propyl) carbodiimide hydrochloride (EDC) and 0.17M N7 Hydroxysuccinimide (NHS) solutions in a 47 morpholineethanesulfonic acid (MES) buffer with a pH of 6.0 for 7.5 min. Then, a 1007Gg/mL neutravidin solution in a 50 mM of sodium acetate buffer with a pH of 4.5 was injected for 30 min. Then, the sensor chip surface was deactivated by ex7 posure to a 17M tris(hydroxymethyl)7aminomethane hydro7 chloride (Tris7HCl) solution for 5 min. This procedure created a sensor surface selective towards different biotinylated mole7 cules. To evaluate the selectivity, we analyzed the adsorption of oligonucleotides D1 and D2 onto the SPR sensor chip with pre7immobilized neutravidin. Oligonucleotides were dissolved in the running PBS buffer (used in all SPR measurements) at a concentration of 50 nM. All SPR responses in the biosensing assay were within the dynamic range (105) of the Acco7 lade 404SA. RESULTS AND DISCUSSION ! The excitation of SPR in the Kretschmann geometry is associated with the opti7 cal coupling of laser radiation to a surface electromagnetic wave propagating along the interface between a thin metal film and dielectric medium. The optical properties of the metal film determine the possibility of SPR excitation in a particular system, the resonance characteristics, and the sensitivity of SPR7based biosensing.44 The permittivity of the copper and gold films used in the proposed SPR biosensor chips was de7 termined by spectroscopic ellipsometry in the wavelength range from 300 nm to 1500 nm. For this purpose, copper and gold films with thicknesses of 26 and 25 nm, respectively, were deposited on the surfaces of silicon wafers. The thick7 nesses of the metal films were confirmed by AFM measure7 ments, which were also used to estimate the surface roughness (Fig. 2). Electron7beam evaporation allows the deposition of copper films with a root7mean square roughness of 0.5 nm, which is sufficient for SPR applications. Figure 2. Atomic force microscopy images of (a) the scratch on the surface of a thin copper film deposited by electron ACS Paragon Plus Environment Langmuir 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 beam evaporation and (b) copper crystallites formed during the deposition. According to the above7described ellipsometric model, the permittivities of the copper and gold films were directly de7 termined from the ellipsometric data (Fig. 3). Table 1 shows the parameters of the Drude model (1) for both films. The approximate equality of the plasma frequencies of the copper and gold films results in the same SPR angles. Therefore, SPR sensor chips based on both copper and gold films can be used with the same SPR instruments without any adjustments of their optical configurations. In addition, copper displays a level of optical losses comparable to those of gold, which im7 plies the equality of the biosensing sensitivities of copper7 and gold7based sensor chips with the same multilayered configura7 tion. The permittivity of copper was 712.03+1.580i at 635 nm, which was the operating wavelength of the laser diode used in the SPR instrument. This permittivity value was subsequently used for SRI modeling. Table 1. The parameters of the Drude model for thin copper and gold films obtained by the fitting of their permittivities. Material ε∞ ωp [1016 /s] Γ [1013 /s] Cu 4.68 1.32 10.5 Au 5.08 1.35 12.8 Figure 3. Real (blue lines) and imaginary (red lines) parts of the permittivity of thin copper (solid lines) and gold (dashed lines) films obtained by spectroscopic ellipsometry. " . The sensitivity to the RI changes of copper SPR sensor chips covered with various oxidation7 protecting layers was investigated both theoretically and ex7 perimentally. The TMM was used for modeling the angular reflection, which gives the angles corresponding to SPR exci7 tation (Fig. 4(a)). The RI change in the upper layers leads to a shift in SPR angular curves, which is the basis of SPR biosens7 ing. Assuming = 0.005 , the RI was calculated for pro7 tecting layers composed of Si, SiC, ZnS, Al2O3, or SiO2 thin films with different thicknesses (Fig. 4(b)). For the high7RI protecting layers, the RI could be improved by nearly 4 times by adjusting the thickness of the protecting layer. The optimal thickness of the protecting layers was determined to be in the Page 4 of 16 range between several nanometers to several tens nanometers, where a smaller thickness corresponded to the larger RI of the layer. The sensitivity enhancement was a result of a change in the plasmonic field distribution in the way that more electro7 magnetic energy was concentrated in the dielectric layer above the chip’s surface. Further RI improvement was limited due to the SPR angular shift towards the high angle region, which made it impossible to efficiently excite SPR in structures with thick protecting layers. Experimentally, RI was estimated for two types of copper SPR chips covered with thin films of SiO2 and Al2O3 with different thicknesses. SPR measurements were conducted on a commercial SPR instrument, which utilizes a constant excita7 tion wavelength and angle of incidence to detect of phase changes in the reflected beam. RI was measured by the anal7 ysis of the SPR signal response to the injections of salt solu7 tions with a 0.88 ⋅10−3 difference in RI. The thicknesses of the SiO2 and Al2O3 protecting layers are in the ranges of 5735 nm and 2.5725 nm, respectively (Fig. 5). Figure 4. (a) SPR curves of copper sensor chips coated with Al2O3 protecting layers of various thicknesses for two sensing media with refractive indices of 1.33 (solid lines) and 1.335 (dashed lines). (b) The sensitivity to refractive index changes of the copper sensor chips coated with various protecting lay7 ers depending on their thicknesses. For all salt injections, the SPR signal demonstrated a stable baseline, which means there was no degradation of the copper plasmonic films. Therefore, the proposed dielectric coatings for the SPR chips provided a sufficient level of oxidation pro7 ACS Paragon Plus Environment Page 5 of 16 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 Langmuir tection for most SPR experiments. However, different thick7 nesses of the protecting layers corresponded to different RI values, which requires that the final structure of the copper SPR chips will need to be optimized. As shown in Fig. 5, the RI of the SPR chips with thin SiO2 and Al2O3 coatings in7 creased with the thicknesses of these layers, reaching peak values at 25 nm and 15 nm, respectively. Our results clearly show that the proposed dielectric coatings not only protected the plasmonic films from oxidation but also promoted an overall improvement in the biosensing sensitivity of the SPR chips. Thus, the corresponding peak values of RI (obtained for Al2O3 and SiO2 coatings) were 75% and 55% higher, re7 spectively, than those of the RI of the bare gold SPR chip. These results were found to be consistent with the results of the theoretical modeling of the SPR angle curves (see Fig. 4). # $% & ' % ' % The development of a selective inter7 face immobilized with ligands is an important part of the SPR biosensing assay. Traditionally, SPR chips use linking layers based on thiol chemistry due to the ability of sulfur7containing molecules to form strong bonds to metal surfaces.45 This ap7 proach is not applicable to the dielectric7layer covered copper SPR chips. Instead, each dielectric layer will generally require its own immobilization procedure. However, graphene materi7 als can help overcome this problem. Various SPR interfaces based on graphene and GO have been proposed in recent years. Due to the large surface area of graphene materials and their diverse chemical properties, these interfaces can be ap7 plied to the analysis of a wide range of biochemical interac7 tions, providing and even improving the immobilization ca7 pacity compared to the thiol7based linking layers.40,46750 Here, we demonstrated the development of a GO linking layer on the surface of SPR biosensor chips coated with dielectric layers for the first time. An aqueous solution of GO was spray7coated onto the sur7 face of a copper SPR chip with an Al2O3 protecting layer. Thereafter, this SPR chip was used for the SPR analysis of a DNA hybridization reaction. For this analysis, the GO surface was immobilized with a selective layer of neutravidin mole7 cules, which possess four binding sites for biotinylated lig7 ands. The neutravidin immobilization was performed in the flow cell of the SPR instrument using an amine coupling pro7 cedure, which includes the activation of carboxyl groups on the surface of the GO and subsequent covalent attachment of the neutravidin to these groups. The SPR signal corresponding to the adsorption of neutravidin was 4000 RU (Fig. 6(a)). For convenience, all SPR signals are given in units of RI, where 1 RU = 1076. The next stage of the SPR biosensing assay included the in7 vestigation of the selectivity of the developed neutravidin7 coated SPR surface as well as the possibility of using this in7 terface to analyze DNA hybridization. For this purpose, three solutions of DNA sequences – D2, D1, and D2 – were injected sequentially. D1 is an oligonucleotide sequence with a bioti7 nylated 5’7end, and D2 is non7biotinylated and complementary to D1. The SPR signals corresponding to the three consecutive DNA injections were 25 RU, 270 RU, and 85 RU, respectively (Fig. 6(B)). The amount of D2 adsorbed via hybridization is more than three times higher than the level of non7specific binding of D2 to the pure neutravidin surface, which clearly shows the selectivity of the developed SPR biosensor chips. In addition, the successful hybridization of complementary oli7 gonucleotide strands confirms the possibility to use the pro7 posed neutravidin7GO SPR interfaces for the investigation of many other types of biochemical interactions. Figure 5. Sensorgrams corresponding to the injections of 0.5% NaCl solution in a running buffer over the copper SPR sensor chips coated with (a) Al2O3 and (b) SiO2 protecting layers of various thicknesses. (c) Sensitivity to refractive index changes of the copper SPR sensor chips coated with Al2O3 (red squares) and SiO2 (blue circles) layers depending on their thicknesses. The dashed line shows the sensitivity level of a bare gold SPR chip. ACS Paragon Plus Environment Langmuir 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 Figure 6. (a) Covalent immobilization of neutravidin on the surface of the graphene oxide (GO) linking layer deposited on the copper surface plasmon resonance (SPR) sensor chip pro7 tected by 157nm7thick Al2O3 film. Immobilization procedure includes the activation of carboxyl groups of GO by the mix7 ture of 0.47M 17ethyl737(37(dimethylamino)propyl) car7 bodiimide hydrochloride (EDC) and 0.17M N7 Hydroxysuccinimide (NHS) solutions and deactivation of car7 boxyl groups after neutravidin adsorption by 17M Tris solu7 tion. (b) Adsorption of oligonucleotides D1 and D2 on the sur7 face of neutravidin7GO copper SPR chip. D1 is biotinylated and complementary to non7biotinylated D2. PBS, phosphate7 buffered saline. CONCLUSIONS Page 6 of 16 respectively, than for gold SPR chips without a dielectric layer on top of them. The subsequent biomolecule adsorption on the surfaces of the proposed SPR chips was achieved using the GO7based linking layers. GO is a promising material for bio7 sensing due to its high surface area, diverse biochemical prop7 erties, and cost7effective production, which allows the deposi7 tion of GO matrices with high immobilization capacity on any biosensing interface. The selectivity of the SPR biosensing assay was achieved by the immobilization of neutravidin on the surface of the GO. Neutravidin molecules possessing four binding sites for biotin residues were covalently attached to the carboxyl groups of the GO using amine coupling. The SPR corresponding to the neutravidin adsorption was 4000 RU, which provides sufficient sensitivity for the analysis of most biochemical interactions. The performance of the neutravidin7 GO surface, which is suitable for the immobilization of vari7 ous biotinylated ligands, was assessed for the analysis of a DNA hybridization reaction. The SPR signals corresponding to the binding of biotinylated oligonucleotide and oligonucleo7 tide hybridization were 11 and 3 times higher, respectively, than for the non7specific binding of non7biotinylated oligonu7 cleotide to the pure neutravidin7GO surface. In addition to their use in traditional SPR biosensing devices, copper plas7 monic biosensors with GO linking layers can be utilized as integrated components of various analytical devices with ap7 plications ranging from medical diagnostics to food and envi7 ronmental safety controls. Traditional semiconductor7 fabrication technologies are suitable for the mass7production of the optical transduction part of such biosensors, whereas the GO linking layers can enable the implementation of various qualitative and quantitative biosensing assays. AUTHOR INFORMATION Corresponding Author * E7mail: [email protected]. Phone: +774987744765782. ORCID Yury V. Stebunov: 0000700027925779595 Notes The authors declare no competing financial interest. ACKNOWLEDGMENT We have proposed SPR biosensor chips based on a copper7 dielectric plasmonic interface. Thin copper films support the excitation of surface plasmons, which can efficiently couple with external laser radiation. Moreover, copper biosensor chips can be adapted for any commercial SPR instrument that was developed for gold plasmonic interfaces without any ad7 justments to the optical configuration. This finding was con7 firmed both by SPR and ellipsometric measurements, which demonstrate the excitation of SPR in copper7 and gold7based optical structures under approximately the same conditions. For SPR biosensing, copper can provide improved biosensing sensitivity compared to gold, owing to lower optical losses. The main drawback of copper usage in biosensing is the rapid oxidation of the metal in most biological solutions. However, the deposition of thin dielectric layers on the copper surface can protect biosensor chips from oxidation and, additionally, significantly improve its biosensing sensitivity. According to SPR measurements, the copper SPR biosensor chips coated with 25 nm SiO2 and 15 nm Al2O3 films provided maximum sensitivity to RI changes, which are 55% and 75% higher, Y.V.S., D.I.Y., A.V.A and V.S.V. acknowledge support from the Russian Science Foundation (17779720345). D.Y.F. acknowledges support from the Ministry of Education and Science of the Russian Federation (8.9898.2017/6.7). The authors thank Slava Petropavlovskikh from the BiOptix Ana7 lytical LLC (Louisville, CO, USA) for expert technical assis7 tance with the experiment. REFERENCES (1) Cooper, M. A. Optical Biosensors in Drug Discovery. ())(, , 515–528. (2) Karlsson, R. SPR for Molecular Interaction Analysis: A Review of Emerging Application Areas. ())*, , 151–161. (3) Homola, J. 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