Nonlithographic Fabrication of Plastic-Based Nanofibers Integrated

Subscriber access provided by READING UNIV

Article

Nonlithographic Fabrication of Plastic-based Nanofibers Integrated Microfluidic Biochip for Sensitive Detection of Infectious Biomarker Brince Paul K, Asisa Kumar Panigrahi, Vikrant Singh, and Shiv Govind Singh ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b11331 • Publication Date (Web): 31 Oct 2017 Downloaded from http://pubs.acs.org on October 31, 2017

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 free 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 accessible to all readers and citable by the Digital Object Identifier (DOI®). “Just Accepted” is an optional service offered to authors. Therefore, the “Just Accepted” Web site may not include all articles that will be published in the journal. After a manuscript is technically edited and formatted, it will be removed from the “Just Accepted” Web site and published as an ASAP article. Note that technical editing may introduce minor changes to the manuscript text and/or graphics which could affect content, and all legal disclaimers and ethical guidelines that apply to the journal pertain. ACS cannot be held responsible for errors or consequences arising from the use of information contained in these “Just Accepted” manuscripts.

ACS Applied Materials & Interfaces is published by the American Chemical Society. 1155 Sixteenth Street N.W., Washington, DC 20036 Published by American Chemical Society. Copyright © American Chemical Society. 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 39

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

ACS Applied Materials & Interfaces

Nonlithographic

Fabrication

of

Plastic-based

Nanofibers Integrated Microfluidic Biochip for Sensitive Detection of Infectious Biomarker

Brince Paul K, †, ‡ Asisa Kumar Panigrahi, ‡ Vikrant Singh,Γ Shiv Govind Singh ‡,*

‡ Department of Electrical Engineering, Indian Institute of Technology, Hyderabad, India † Department of Biomedical Engineering, Indian Institute of Technology, Hyderabad, India. Γ School of Medicine, University of California Davis, USA

KEYWORDS: Biochip, Tune transfer method, Sensing platform, C-ZnO nanofibers, Infectious diseases

ABSTRACT We report fabrication of a fully integrated plastic based microfluidic biochip for biosensing application. The microfluidic channels were fabricated by tune transfer method and integrated with the pre-functionalized sensing platform. This approach to assemble microchannel into prefunctionalized sensing substrate facilitates controlled functionalization and prevents damages on functionalized surface. The sensing platform comprised of a three electrode system, in which the sensing electrode was integrated with antibody immobilized carbon nanotubes-zinc oxide (C-

ACS Paragon Plus Environment


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

ZnO) nanofibers. Electrospinning technique was used to synthesize C-ZnO nanofibers and the surface of the nanofibers was covalently conjugated with histidine rich protein II antibodies (AntiHRP II) toward detection of infectious malarial specific antigen, namely histidine rich protein II (HRP II). The analytical performance of the fabricated biochip was evaluated by differential pulse voltammetry method. The device exhibited a high sensitivity of 1.19 mA/ g mL-1 /cm2 over a wide detection range (10 fg/mL - 100 µg/mL) with a low detection limit of 7.5 fg/mL towards HRP II detection. This fully integrated biochip offers a promising cost-effective approach for detection of several other infectious disease biomarkers.

1. INTRODUCTION The integration of the microfluidic systems with point of care diagnostics (POC) has recently gained a considerable attention owing to their several advantages over traditional diagnostic instruments in respect

of miniaturization and compactness, enhanced mass transport and

diffusion, reduced consumption of reagents, and protection of the sensing area from the environment.1,2 Numerous microfluidic integrated biosensors based on optical,3 fluorescence,4 surface plasmon resonance (SPR)5, and surface enhanced raman scattering (SERS)6 have been reported. However, these detection methods requires bulky set up, expensive source and detectors, and additional signal extraction process.7 On the other hand, electrochemical detection technique facilitates the benefits of easiness and high sensitivity.8 Furthermore, the fabrication process of these sensors is less complicated and can be directly interfaced with most measurement systems.9 Many groups have reported microfluidic integrated electrochemical immunosensors.10-13 A critical step in fabrication process involves sealing the devices by 2 ACS Paragon Plus Environment

Page 2 of 39

Page 3 of 39

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


bonding channel to a sensing substrate without damages on the bio-fictionalized sensing surface. Although, surface modification and antibody conjugation can be adopted after the device sealing, the channel and nonspecific areas of the device would be contaminated by the reagents used for functionalization.14 In this concern, assembling a microfluidic channel to a pre-functionalized sensing substrate continues to be a challenge. The development of microfluidic based devices integrated with one-dimensional (1D) metal oxide nanostructured materials have fascinated growing interest in biosensing because of their mechanical stability, enhanced electro catalytic

properties and diffusivity, and high

surface to volume ratio.15,16 These 1D nanostructures act as ‘electronic wires’ for the conduction of electrons, and also help in increased loading of the biomolecules.17,18 Among the metal oxide semiconductors, zinc oxide (ZnO) offer attractive properties like high electron transfer ability, biocompatibility, non-toxicity, and high free-exciton binding energy (60 meV).19,20 Furthermore, the high isoelectric point (IEP=9.5) of ZnO aids to immobilize a biomolecule with low IEP.21 ZnO nanostructures based immunosensors have widely been explored.22-27 However, inherently poor electrical conductivity and further surface functionalization for bioconjugation are the foremost pitfalls of ZnO that needs to be addressed while developing biosensors.28 In this context, the sensing characteristics of ZnO based biosensors can be enhanced by doping or grafting carbon nanotubes(CNTs) in to ZnO nanostructures.29,30 A few study available on CNTs- ZnO matrix integrated biosensors have proven that the incorporation of CNTs with ZnO can improve sensing characteristics of biosensors compared with the biosensors based on ZnO or CNTs alone.31-36 In a previous study, we have demonstrated carbon nanotubes embedded ZnO nanowires modified conventional bulk glassy carbon electrode based biosensor for cancer biomarker detection.37 3 ACS Paragon Plus Environment


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

We have also developed chemiresistive biosensor based on these nanowires.38 However, the integration of 1D CNTs- ZnO (C-ZnO) nanostructures with microfluidic device has great potential for the improvement of biosensor efficiency in terms of sensitivity and miniaturization. There has been an intensive effort in the integration of nanostructures with microfluidic devices in POC diagnostics, such as nickel oxide nanorods based biochip for cholesterol estimation,18 nickel oxide-carbon nanotube composite based microfluidic chip for cholesterol detection,39 chitosan functionalized titanium dioxide nanoparticles integrated microfluidic device for analysis of total cholesterol,40 carbon nanotubes−nickel oxide nanocomposite based smart lab on a chip for lipoprotein detection,41 and graphene foam modified carbon-doped titanium dioxide nanofibers electrode based immune biochip for detection of cancer biomarker.13 However, these device fabrication and bonding requires clean room facilities, thereby resulting in increased production costs and time. Moreover, glass substrate based devices require extra care while handling, which bounds their application in medical laboratories. Hence, there is a need to replace the glass substrate based microfluidic POC device with an affordable, disposable, and user friendly substrate based device. In recent years, paper and plastic substrate fabricated devices have received great consideration for the POC application in terms of miniaturization and portability. Moreover, these substrates are simple to manufacture, inexpensive, mass producible, and readily available around the world.42,

43

There have been some studies for the

POC devices fabricated on these substrates. Josephine et al. fabricated paper-based paper analytical device for oligonucleotides and protein detection.44 Shenguang et al. developed gold nanoparticle integrated paper based electrode for Microcystin-LR immunoassay.45 Jan et al. designed paper built electrochemical device for glucose sensing.46 Zhihong et al. reported 4 ACS Paragon Plus Environment

Page 4 of 39

Page 5 of 39

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


fabrication of paper based electrochemical device for glucose estimation.47 Nandhinee et al. demonstrated polyimide substrate based electrochemical sensor platform for cardiac troponin monitoring.48 In most cases, soft lithography or wax printing technology was used to pattern microfluidic channels on paper devices. In this endeavor, we propose a fully integrated plastic substrate based microfluidic device for biosensing application without using wax printing or lithography technique. Malaria is one of the foremost infectious diseases caused by protozoan of the genus Plasmodium spread by female Anopheles species mosquitoes.49 According to the latest report (2016) from world health organization (WHO), there were 212 million cases of malaria and an estimated 429, 000 malaria deaths.50 Although more than 100 species of Plasmodium are available, Plasmodium falciparum (Pf) parasite can cause severe malaria in humans.51 histidinerich protein II (HRP II), found in several cell compartments of Pf, and assays based on pfHRP II shows the better sensitivity for the detection of Pf.52 At present, numerous methods, including mass spectrophotometry,53 molecular diagnostic,54 microscopic technique,55 lamp technique,56 polymerase chain reaction (PCR),57 blood cell analyzers,58 and flow cytometry, are usually used to detect Pf.59 However, these techniques are expensive, time consuming, and have need of expert personnel for conducting process . In the view of the above, the development of a simple, cheap, and fully integrated portable immunosensors are extremely required. In this work, we introduce the fabrication of a fully integrated plastic – based biochip for histidine rich protein II (HRP Ⅱ) detection. The microfluidic channels were fabricated by novel tune transfer method. Compared to the multi-process soft lithographic procedure, tune transfer method is simpler, cost-effective, facilitates rapid prototyping, and it doesn’t require cleanroom facilities. The fabricated micro channel was integrated with the plastic sensing substrate 5 ACS Paragon Plus Environment


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

Page 6 of 39

functionalized with carbon nanotubes-zinc oxide nanofibers (C-ZnONFs) and HRP II antibodies (AntiHRP II). The 1D carbon nanotubes-zinc oxide nanofibers (1D C-ZnONFs) were synthesized by electrospinning method. Among the various techniques, electrospinning is a remarkably simple, cost-effective method of producing long continuous 1D nanostructure in a large scale.60 The incorporation of carbon nanotubes not only enhances the conductivity of nanofibers but also generates the functional groups on the nanofibers surface during the calcination process of C-ZnONFs formation that facilitates covalent immobilization of HRP II antibodies devoid of additional surface modifications. The AntiHRP II was covalently immobilized to a C-ZnONFs modified sensing electrode. Assembling a microfluidic channel to a pre-functionalized chip may provide controllable functionalization and prevent damages on the biosensor. This fully integrated device holds promising opportunities to develop several other infectious diseases biomarker detection systems. 2. EXPERIMENTAL SECTION 2.1.

Materials and Reagents

Histidine rich protein II antigen (HRP II), histidine rich protein II monoclonal antibody (Anti HRP II), and CA 125 were procured from Fitzgerald, USA. Polyacrylonitrile (PAN, MW=150,000), myoglobin, bovine serum albumin (BSA), N, N-dimethyl-formamide(DMF), , N-ethyl-N-(3-dimethylaminopropyl phosphate

buffer

saline

(PBS,

carbodiimide) pH

7.4)

(EDC), tablets,

N-hydroxysuccinimide and

zinc

acetate

(NHS),

dehydrate

(Zn(CH3COO)22H2O,99%) were purchased from Sigma- Aldrich, USA. Immunoglobulin G (IgG) and creatine kinase were acquired from Abcam. Pristine multi walled carbon nanotubes (MWCNTs, diameter: 20–70 nm and length: 1-10µm) was supplied by Reinste Nano Ventures,

6 ACS Paragon Plus Environment

Page 7 of 39

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


India. Polydimethylsiloxane (PDMS) was acquired from Dow Corning, USA. PBS ( pH 7.4) was used for preparing stock solution and different concentrations of HRP II. 2.2.

Instruments

The characterization of the synthesized 1D nanofibers were performed using transmission electron microscope(TEM, FEI-Tecnai G2- F30, 300 KV), field emission scanning electron microscopy (FE-SEM; JOEL JSM-7600F), X-ray diffraction (XRD; Bruker Advance D8 X-ray diffractometer with Cu Kα radiation and λ = 0.154 nm), and X-ray photoelectron spectroscopy (XPS; Thermo Scientific Multilab 2000, Mg Kα X-rays as the source). The electrode modifications were observed using Energy dispersive X-ray spectroscopy (EDX; Oxford Instruments) and Atomic force microscopy (AFM;

Bruker Instruments). CHI660E (CH

Instruments, TX, USA) electrochemical analyzer was used for all electrochemical analysis and 0.01M PBS (PH 7.4) containing 5 mM ferro/ ferricyanide redox couple [Fe (CN)6]3-/4- was used as an electrolyte. 2.3.

Synthesis of C-ZnONFs

The electrospinning technique was used to synthesis C-ZnONFs using Zn(CH3COO)2.2H2O as a precursor material. Briefly, MWCNTs were dispersed in DMF via ultra-sonication, followed by the addition of 8 wt.% PAN. The mixture was vigorously stirred at 65˚C for 2 h. Then, 4 wt.% of Zn(CH3COO)2.2H2O was added into above solution by stirring at 60 ˚C for 3 h. The wt% of MWCNTs, PAN and Zn(CH3COO)2.2H2O were optimized to achieve an identical bead free fiber devoid of electrostatic repulsion hindrance. The prepared homogeneous precursor was filled into a syringe having 26 gauge metallic needle, which was connected to a high power source. The electric field (2 kV/cm) applied between the needle and collector (copper plate covered with aluminum foil) and flow rate (40 μL/min) were optimized to get a nanofiber mat. The thick free



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

standing PAN/MWCNTs/ Zn(CH3COO)2.2H2O polymeric nanofiber mat was collected after 3 h of electrospinning. Finally, the nanofiber mat was calcined at 400 °C for 2 h to form CZnONFs after the decomposition of PAN polymer. During this calcination process, the defective carbons present in the nanofibers were functionalized in to carboxylic groups that facilitates covalent immobilization of biomolecules.61,62 2.4.

Fabrication of plastic based sensing platform

In this work, the sensing platform comprises of three electrode system was fabricated on a 300 µm-thick transparent polyethylene terephthalate (PET) plastic sheet. These electrode system was made-up of Platinum (Pt, as counter electrode), Silver/silver chloride (Ag/AgCl, as reference electrode), and Gold (Au, as working electrode). Prior to the fabrication of electrodes, the PET substrates were thoroughly cleaned by successive washes in acetone, isopropanol (IPA), and DI water. A cleaned PET substrate was placed in a sputtering chamber and 200 nm- thick Au layer was sputtered through an aligned shadow mask having a pattern of working electrode. Chrome (Cr) layer of 20 nm thickness was used as an adhesion layer for the gold onto the PET substrate. The Ag (200nm) and Pt (200nm) layers were deposited using their respective shadow masks. Titanium layer (20 nm) was used as an adhesion layer for silver and platinum. The fabricated electrodes were cleaned by DI water and then they were dried. The Ag/AgCl reference electrode was developed by treating Ag electrode with 1M KCl. The Au working electrode had dimensions of 2 mm diameter, and counter and reference electrode had dimensions of 1×1.5 mm. The integration of 1D C-ZnONFs on the sensing surface (Au working electrode) was achieved by drop-casting process. In this step, the synthesized C-ZnONFs was dispersed in DMF (0.5 mg mL-1) using an ultrasonic bath for 1h. 10 µL of nanofiber dispersion was drop casted on


Page 8 of 39

Page 9 of 39

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


to the Au electrode surface and dried at 50 °C for 1h. This was followed by washing with DI water and dried. Finally, the AntiHRP II was covalently immobilized to the nanofiber integrated sensing surface. For this purpose, the carboxylic (–COOH) groups present on the C-ZnONFs surface were activated by treating with cross linker solution containing 0.4 M EDC and 0.1M NHS for 4 hours. After washing with PBS, the activated C-ZnONFs surface was incubated with 10 µL of AntiHRP II (200 µg/mL) overnight at 4 °C. The strong amide (C-N) bonds developed between the activated −COOH groups on nanofiber surface and primary amine (-NH2) groups of antibody enable covalent immobilization of AntiHRP II antibodies onto C-ZnONFs surface. Then, the antibody modified sensing surface was rinsed with PBS (pH 7.0) multiple times to remove all unbounded antibodies. The nonspecific binding sites of the sensor surface was blocked by incubating with BSA (1 wt %) for 1h. Again, the sensing surface of the fabricated device was washed with PBS (pH 7.0) and stored at 4 °C.

2.5.

Fabrication of microfluidic channel by tune transfer method and sensor integration

The PDMS micro channel of dimensions 2 cm × 500 µm × 30 µm was fabricated using tune transfer method. In brief, the channel pattern was designed by CAD software and printed on a glossy paper using an ordinary printer. Then, the printed channel pattern was tune transferred to a copper board (30 µm copper layer thickness). In this step, after glossy paper with printed pattern side down had been placed on a copper surface, a little pressure was applied all along the paper by moving an electric iron (with maximum temperature) around for about 5 min. The printed pattern on the glossy paper was transferred on to the copper surface due to heat from the iron. Next, the glossy paper was removed from the surface with water and the tune transferred



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

Page 10 of 39

board was immersed in ferric chloride (FeCl3) solution for 30 min to etch the copper. During this process, the transferred tuner acts as an etch resist (mask) in FeCl3 and unmasked areas of the copper are etched away, leaving a pattern under the masked regions. The patterned board was cleaned with acetone, DI water and allowed to dry. The PDMS precursor was mixed with curing agent in a weight ratio of 10:1, followed by degassing for 30 min under vacuum. Then, the degassed PDMS polymer was poured onto the patterned board, and cured at 80 ˚C for 1 h. Finally, the cured PDMS channel replica was released from the board and holes were punched for inlet and outlet tubing. The fabricated PDMS channel was adhered to a plastic sensing platform using 3M Scotch permanent double sided transparent tape. In this step, the channel pattern was cut in to the Scotch tape according to the channel dimension (2 cm × 500 µm). Then, the patterned tape was pasted on to the sensing platform in such a way that the antibody functionalized working electrode, counter and reference electrodes lie within the patterned channel area. Finally, the PDMS channel was aligned under a simple optical microscope and sealed to a sensing platform. The sealed chip was baked at 50 °C for 1h to enhance the bond strength and flexible tubes coupled with connecters were assembled to channel inlet/outlet ports to complete the packaging. The advantages of use of a transparent adhesive tape for bonding microfluidic channel are; (1) facilitates assembling the microchannel to a pre-bio functionalized sensing substrate without damages on functionalized region; (2) facilitates optical visualization of the channels for imaging.63 The Schematic representation of the nanofiber synthesis, fabrication of sensing platform and functionalization, and the fabrication steps for microfluidic channel and sensor integration is shown in Figure 1.


Page 11 of 39

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 1: Schematic of stepwise fabrication process and functionalization procedure of the plastic based biochip for infectious detection application.

3. Results and Discussions 3.1.

Morphological and Structural characterization

The FESEM imaging was used to analyze the morphological features of the nanofibers. Figure 2a shows the synthesized PAN/MWCNTs/ Zn(CH3COO)2.2H2O polymeric nanofibers, which have the diameter ranges from 300–400 nm. The electrospinning parameters (electric field, feed rate, and precursor concentrations) were optimized to attain polymeric nanofibers with desired morphology. Figure 2b shows the FESEM image of the C-ZnONFs, which were obtained after calcination of the above synthesized polymeric fibers. The resultant diameters of the nanofibers shrunk to about 180-200 nm. The shrinkage in nanofiber diameter can be attributed to the 11 ACS Paragon Plus Environment


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

Page 12 of 39

decomposition of metal precursor and selective removal of PAN polymer by calcination process. The formation of the nanostructures is further confirmed with TEM analysis. The TEM image of C-ZnONFs (Figure 2c) reveals the nanofiber possess 1D structure. The selected area electron diffraction (SAED) pattern (inset of Figure 2c) of C-ZnONFs appears with alternating black and bright rings indicating its polycrystalline nature. The crystalline structure of the synthesized CZnONFs was investigated by XRD Studies (Figure 2d). The diffraction peaks obtained at 2θ = 31.7°, 34.4°, 36.28°, 47.5° and 56.6° correspond to the (100), (002), (101), (102) and (110) ZnO planes, respectively. The structural analysis indicated that the formed ZnO are polycrystalline and have a wurtzite structure. The peak observed at 2θ =26° correspond to (100) graphitic plane and the peak at 2θ =43° correspond to (100) graphitic plane.64 Figure 2e shows the optical image of PET sensing platform comprises of nanofiber and AntiHRP Ⅱ modified Au working electrode, pt counter electrode, and Ag/AgCl reference electrode integrated with the microchannel. The inlet of the PDMS microchannel was connected to the syringe pump to control the flow though the microfluidic chip. Figure 2f shows the twodimensional AFM image of CZnONFs on the Au electrode surface, which are well aligned and distributed uniformly. The formation of Ag/AgCl reference electrode was confirmed with EDX spectroscopy analysis. The EDX spectra of silver electrode showed only the presence of Ag element, and after treating the silver electrode with KCl, an additional spectrum of Cl element appeared (Figure S1). These results revealed that the AgCl layer was effectively formed by KCl treatment with the silver electrode.


Page 13 of 39

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 2: FESEM image of (a) as-spun polymeric nanofibers (b) resultant C-ZnONFs (c) TEM image of nanofiber mat; Inset shows SEAD pattern (d) XRD results of C-ZnONFs conducted between 2θ angles 20o and 70o (e) Optical image of the fabricated device connected with inlet and outlet tubing (f) 2D AFM image of the C-ZnONFs modified sensing electrode.

To further confirm the functional groups present in the C-ZnONFs surface, element chemical state, and immobilization of Anti HRP II on the surface of C-ZnONFs, XPS studies was performed. Figure 3a shows wide scan spectra obtained for the C-ZnONFs and AntiHRP II conjugated C-ZnONFs. The peak at 1021 is assigned to Zn2p3/2, while the peak at 1043 eV is assigned to Zn2p1/2. The peak at 283.2 is attributed to C1s, and the peak at 530.4 eV is attributed to O1s of C-ZnONFs. The presence of C1s, O1s, and Zn2p peaks indicate the incorporation of carbon dopant in C-ZnONFs. The N1s peak observed in wide scan spectra of AntiHRP II conjugated C-ZnONFs reveal the successful immobilization of HRP II antibodies on the surface of C-ZnO nanofibers. Table S1 summarizes the atomic percentage of different elements existing 13 ACS Paragon Plus Environment


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

Page 14 of 39

in the C-ZnONFs and AntiHRP II modified C-ZnONFs. After the conjugation of antibodies on the C-ZnONFs, the atomic concentration of nitrogen was found to be 11.3%. Figure 3b shows the deconvolution of C1s core level XPS spectra for C-ZnONFs into characteristic peaks with a Gaussian profile. The peaks 284.5 eV is assigned to C-C (graphitic nature), and the peak at 286.6 eV is assigned to C-O groups. The peak at 285.5 eV is ascribed to hydroxyl (C-OH) group, while the peak found at 288.4 is attributed to the carboxylic (O-C=O) group present in C-ZnONFs. Figure 3c shows the C1s core level spectra for AntiHRP Ⅱ immobilized C-ZnONFs. The additional peak occurred at 287.13 eV is ascribed to the formation of amide (N-C=O) bonds between the -NH2 groups on AntiHRP II and the activated −COOH groups on C-ZnONF. However, the peaks related to graphitic, C-O, hydroxyl, and carboxylic groups are slightly shifted towards lower binding energy positions due to covalent functionalization of antibody. The slight shift in binding energy position may perhaps be attributed to the electron donation from the adjacent nitrogen atom of amide groups. This phenomenon can be explained as follows: The p orbital of nitrogen in the amide group and the carbonyl group oxygen contains nonbonding electrons. During conjugation of these p orbitals, the neighboring nitrogen atom donates an electron which results shift in binding energy position.65 The atomic percentage of different chemical groups existing in the CZnONFs and antiHRP Ⅱ modified C-ZnONFs , full width at half maximum(FWHM), and their corresponding binding energy are summarized in Table S2. After the functionalization of the antibodies on CZnONFs surface, the relative atomic percentage of O-C=O group present in C-ZnONFs reduced, indicating that most of the carboxylic groups are used in antibody immobilization. The XPS analysis of N1s core level spectra of antibody functionalized C-ZnONFs


Page 15 of 39

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 3: XPS analysis of (a) wide scan spectra obtained for the C-ZnONFs(i) and wide scan spectra obtained for Antibody conjugated C-ZnONFs(ii) (b) Core level C1s spectra for CZnONFs (c) ) Core level C1s spectra for antibody immobilized C-ZnONFs (d) N1s core level spectra of antibody functionalized C-ZnONFs.

is shown in Figure 3d. The peaks at 398 eV and 400.4 eV correspond to N1s (-N=) and existence of amide nitrogen (CO-NH), respectively. The peak appear at 401.4 eV is attributed to existence of quaternary N structure, which further endorses covalent functionalization. 15 ACS Paragon Plus Environment


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

3.2.

Page 16 of 39

Device characterization and sensing

The device performance at every stage of surface functionalization was analyzed using electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV) studies to investigate the electrochemical properties of the fabricated biochip. EIS studies were performed by applying a sinusoidal waveform with frequency ranging from 0.1Hz to 105 Hz. The electrochemical impedance Z (ω) of electrode reaction can be decomposed to their real (Z’) and imaginary part (Z”) as;

Z (ω) = (Z’) - j (Z”) with the phase angle (φ) = tan-1(Z”/ Z’)

where Z’ is equal to

and Z” is equal to

(1)

. The charge transfer

resistance (Rct) and double layer capacitance (Cd) at the electrode /electrolyte interface can be obtained from Randles electrical equivalent circuit modelling.66 Figure 4a shows the Nyquist plot of EIS curve obtained for devices corresponding to the different surface modification on sensing electrode. The Rct value obtained for the C-ZnONFs/Gold electrode (85.75 Ω) is found to be lower than that of the unmodified gold electrode (1.35 kΩ). This is ascribed to the electrode surface modification with C-ZnONFs enhances the electron transfer kinetics between electrode and electrolyte. After the conjugation of AntiHRP II and BSA molecules onto the nanofiber modified electrode surface , the Rct value increases to 543.5 Ω implying that the molecules immobilized electrode surface (BSA/AntiHRP II/C-ZnONFs/Gold) causes the hindrance to the diffusion of electrons towards the electrode.


Page 17 of 39

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


The heterogeneous electron transfer rate constant (ks) of the modified electrodes can be estimated using Eq.2 67

(2)

⁄

where R is the universal gas constant, T is the absolute temperature, n is the electron transferring constant, F is the Faraday’s constant, A is the effective surface area, and C is the concentration of redox probe. The lower ks of BSA/AntiHRP II/C-ZnONFs/Gold electrode (3.128 × 10-6 cms1

) compared to that of C-ZnONFs/Gold electrode (19.76 × 10-6 cms-1) indicates sluggish electron

charge transfer kinetics due to the electron transfer hindrance delivered by the BSA and antibody immobilized electrode. The binding site fraction for BSA/AntiHRP II/C-ZnONFs/Gold electrode is estimated to be 0.84, which indicates more than 84% the C-ZnONFs/Gold electrode surface covered with BSA-AntiHRP II molecules 22. The CV measurements were conducted to study the electrochemical redox behavior of the different surface functionalized electrodes in the potential range of -0.4 V to +0.8 V at 50 mVs-1 scan rate (Figure 4b). The peak current of C-ZnONFs integrated electrode was observed to be higher (206.7 µA) as compared to that of the bare electrode. This may be owing to the existence of nanofibers that provide pathways for the fast electron transfer kinetics toward the electrode surface. Further, it was found that the immobilization of AntiHRP II and BSA molecules reduces the peak current (185.5 µA) and increases peak potential separation (ΔE= 125 mV) because of the insulating nature of the proteins, which causes slower electron transfer toward the electrode. The diffusion coefficient (D) of various surface modified electrodes were calculated using Randles- Sevick Eq.3 68 17 ACS Paragon Plus Environment


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

⁄

where,

⁄

⁄

Page 18 of 39

(3)

is the redox peak current and

is the scan rate (V/s). The diffusion coefficient

obtained for the BSA/AntiHRP II/C-ZnONFs/Gold electrode was found to be 386.1 µcm2/s which is lower than that of C-ZnONFs/Gold electrode (479.5 µcm2/s). .

Figure 4: (a) EIS results for the C-ZnONFs, BSA and antibody modified device (b) CV results for the C-ZnONFs, BSA and antibody modified device (c) CV response of the device at different flow rate (d) diffusivity versus flow rate plot obtained for the device at different flow rate in PBS containing 5 mM [Fe (CN)6]3-/4-.


Page 19 of 39

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


This is because of the functionalization of antibody and BSA layer on electrode surface, which causes the sluggish diffusion of [Fe (CN)6]3−/4− ions at the electrode surface interface. The surface concentration (Γ) of the BSA/AntiHRP II/C-ZnONFs/Gold electrode was calculated using the Brown–Anson model as per Eq.4 69

(4)

where Ip/ν is the calibration plot slope. The surface concentration (Γ) was found to be 1.257 × 10-7 mol/cm2, indicating a high coverage of BSA-AntiHRP II molecules on the nanofiber modified electrode surface. The effect of flow rate of the fabricated device was investigated using CV technique at a flow rate from 0.01 µL/min to 15 µL/min. Figure 4c shows CV response of the nanofiber modified device at different flow rate. It was observed that the peak current increases with the rising increasing flow rate. The redox species diffusivity at different flow rate were estimated using the Randles- Sevick equation (Eq.3) and the corresponding diffusivity versus flow rate plot is shown in Figure 4d. The diffusivity increases with the increase in the flow rate due to enhanced mass transport at higher fluid velocity and it reaches saturation at a flow rate of 5 µL/min.70 Thus, all the electrochemical measurements were conducted at this optimum flow rate. To study the nature of the redox process, CV of the BSA/AntiHRP II /C-ZnONFs/Gold electrode was conducted at different scan rates (20-200 mVs-1) (Figure 5a). The anodic and cathodic peak currents was found to be increased linearly with the square root of the scan rate (Figure 5b), signifying that the quasi-reversible diffusion controlled process happening at the electrode



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

Page 20 of 39

surface. The correlation between peak current and square root of the scan rate can be expressed by Eq.5 and Eq.6 ⁄

[

⁄

⁄ ]

[

⁄

⁄ ]

(5)

⁄

(6)

The sensing characteristics of the fabricated biochip were explored as a function of HRP II concentration using differential pulse voltammetry (DPV) technique in the potential ranges from -0.2 V to +0.6 V. The various concentrations of HRP II proteins were injected into the microchannel. Figure 5c shows the DPV response of the device with HRP II concentration over the range of 10 fgmL-1 to 100 µgmL-1. It was observed that the magnitude of peak current decreases with increased concentration of target HRP II. This indicates that the formation of immuno- complexes between HRP II and AntiHPRP II at the electrode surface hinders redox probes diffusion towards the electrode surface. Control experiments for fabricated device were performed without immobilizing antibody on C-ZnO nanofiber modified working electrode surface. As shown in the Figure S2, there was no significant variation in the response signal obtained when increasing concentrations of HRP II protein as compared to the antibody immobilized device (Figure 5c). The insignificant variation in the response current perhaps owing to non-specific interactions of HRP II proteins on to nanofiber modified electrode surface. This indicates that the response in current was mainly due to the interaction between HRP II protein and AntiHRP II functionalized C-ZnO nanofiber surface. The correlation between the response current and logarithmic value of HRP II concentration in the range of 10 fg to 100


Page 21 of 39

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 5: (a) CV studies of BSA and antibody modified device at different scan rate (20 mV/s200 mV/s) (b) The correlation between peak current and square root of the scan rate (c) DPV response of the device in detecting various HRP II concentration (d) Calibration plot showing the response current versus logarithmic value of HRP II concentration.

µg/mL was obtained by fitting Rodbard’s four parameter model relating a sigmoidal behavior (Figure 5d) and is represented by Eq.7

(

(

)

(7) )



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

Page 22 of 39

The sensitivity of the device from the calibration curve was found to be 1.19 mA g-1 mL cm-2. The limit of detection (LOD) of the device was estimated using Eq.8 71

⁄

(8)

where EC50 is the concentration at 50% of the highest signal variation, A1 is the minimum asymptotes of curve, A2 is the maximum asymptotes of the curve, p is the slope at inflection point, and σ is the standard deviation of blank. The LOD of the device was obtained to be 7.5fgmL-1.The sensing performances of fabricated device in terms of detection limit, sensitivity, and fabrication methods compared to other electrochemical microfluidics devices are summarized in Table 1. It can be noted that the C-ZnO nanofibers integrated microfluidic biochip exhibits higher sensitivity and lower detection limit as compared to that of other microfluidic devices for biomarker detection. This may perhaps be due to the large surface feature of C-ZnO nanofibers facilities more loading of Anti HRP II antibodies for specific interaction with HRP II antigens. In addition, C-ZnO nanofibers provide a channel for higher charge transport to the electrode surface resulting in enhanced sensitivity.


Page 23 of 39

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


Table 1. Sensing Performance Comparison between Our Fabricated Device and Other Reported Electrochemical Microfluidic Devices

bioelectrode

detection technique

biomarker

limit of detection

sensitivity

CdSe@ZnS Qds

square wave voltammetry Amperometric

APoE

12.5 ng mL-1

-

PSA

0.23 pg mL−1

CA125

AuNPs GO film

pulse voltammetry

microfluidic fabrication method photolithography

ref

2.5 µA/mL fg-1 /cm2

machining

73

0.001 ng mL-1

-

photolithography

74

72

G/Thi/AuNPs

pulse voltammetry

CEA

10 pg mL-1

-

paper pattern

75

GF−nTiO2

pulse voltammetry

ErbB2

-

0.585 µA/µM

photolithography

13

Carbon

pulse voltammetry

TNFα

4.1 ng mL-1

-

PSA sheet cut

76

DTSP-SAM

cyclic voltammetry

Cortisol

10 pg mL-1

0.207 µA/ M

green tape pattern

77

C-ZnONF

pulse voltammetry

HRP II

7.5 fg mL-1

1.19 mA/ g mL-1 /cm2

tune transfer

Present work



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

Page 24 of 39

Figure 6: (a) DPV response of the device in presence of HRP II with various concentration of interference (b) Histogram plot showing the device response toward presences of various interference (c) Reproducibility studies of the different devices fabricated under similar conditions; Inset shows the histogram plot for the current outputs of the devices (d) stability test results of the device.

The selectivity study of the device was conducted in presence of HPR II (1 pg/mL) protein with various interferences such as IgG (10 ng /mL), myoglobin (10 ng /mL), CA125 (10 ng/ mL), creatine kinase (10 ng/mL), and mixture of all at the same flow rate (Figure 6a). A negligible effect in response current was observed due to the interference with respect to HRP II (Figure 6b). The low relative standard deviation (RSD=3.1 %) obtained reflects the good


Page 25 of 39

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


selectivity of the fabricated device. The reproducibility study was performed by observing the device response towards 1 pg /mL of HRP Ⅱ concentration detection using four various devices fabricated under similar conditions (Figure 6c). The fabricated devices exhibited good reproducibility for the detection of HRP Ⅱ protein as evident by low RSD (4.6%). The repeatability study of the device was carried out in the presence of 10 fg/mL of HRP II concentration using CV method. There was no significant variation in the response in current even after 12 times of repeated measurements (Figure S3). The relative standard deviation (RSD) was estimated to be as 1.66 %, which indicates that the fabricated device shows good repeatability. The storage stability of the device was studied by monitoring the change in magnitude of the current response at a regular interval of 5 days over 20 days (Figure 6d). The biochip was stored at 40C, when not in use. It was observed that device retains 95 % of the initial response after storage for 20 days, indicating acceptable stability of the device. In order to assess the practical use of the fabricated biochip, we executed the detection of HRP II in diluted human blood serum sample.

Serum sample spiked with 100 pg/mL

concentration of HRP II was applied to the biochip, and current response was recorded. Figure 7 shows the comparison of device performance in buffer and serum sample for 100 pg/ mL HRP II detection. The percentage (%) change in current response has been calculated using Eq.9 78 (9)

Where, I0 is the mean current at zero target concentration and Ic is the mean current at any target concentration. The negligible variation between percentage of change in current response



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

Page 26 of 39

(less than 5%) obtained for serum sample and the standard buffer indicates the potential application of fabricated biochip in real sample analysis.

Figure 7: Comparison of percentage change in current response towards 100 pg mL-1 HRP II detection in standard buffer and blood serum sample.

4. Conclusion In summary, we have demonstrated the fabrication of a fully integrated plastic based biochip for the label free detection of infectious malarial biomarker. The micro channels have been fabricated by tune transfer method and integrated with the antibody immobilized nanofiber modified sensing substrate. The C-ZnONFs produced by electrospiining method and the nanofibers surface covalently conjugated with antibodies via EDC-NHS chemistry. The microfluidic assembly process utilizes the simple and inexpensive bonding technique, which facilitates integrating a micro channel to a pre bio-functionalized sensing substrates without any


Page 27 of 39

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


damages on the sensing surface. The fabricated biochip exhibits good sensitivity of 1.19 mA/ g mL-1/cm2 over a wide detection range of 10 fg -100 µg/mLof HRP II protein with a detection limit of 7.5 fg/mL. Furthermore, the device also shows good selectivity, stability and reproducibility to the target protein. This fully integrated biochip has great potential towards developing smart point-of-care devices for detecting several significant infectious disease biomarkers.

ASSOCIATED CONTENT Supporting Information

Additional figures related to EDX spectra confirming Ag/AgCl formation, Repeatability and control study of the fabricated device, and atomic concentration of various elements present in the nanofiber and antibody conjugated nanofiber surface. AUTHOR INFORMATION Corresponding Author *E-mail address: [email protected]

ACKNOWLEDGMENTS The authors would like to express their sincerest gratitude to the INUP, IISC Bangalore and INUP, IITB for the support to material characterization facility. The authors would like to thank A.Soundararaj, Physics department, IITH for his valuable support during the initial characterization of fabricated device.



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

Page 28 of 39

REFERENCES

(1)

Choi, S.; Goryll, M.; Sin, L. Y. M.; Wong, P. K.; Chae, J. Microfluidic-Based Biosensors toward Point-of-Care Detection of Nucleic Acids and Proteins. Microfluid. Nanofluidics 2011, 10 (2), 231–247.

(2)

Gervais, L.; De Rooij, N.; Delamarche, E. Microfluidic Chips for Point‐of‐care Immunodiagnostics. Adv. Mater. 2011, 23 (24).

(3)

Ihara, T.; Mori, Y.; Imamura, T.; Mukae, M.; Tanaka, S.; Jyo, A. Colorimetric Multiplexed Immunoassay Using Specific Aggregation of Antigenic Peptide-Modified Luminous Nanoparticles. Anal. Chim. Acta 2006, 578 (1), 11–18.

(4)

Hauer, P.; Le Ru, E. C.; Willmott, G. R. Co-Ordinated Detection of Microparticles Using Tunable Resistive Pulse Sensing and Fluorescence Spectroscopy. Biomicrofluidics 2015, 9 (1), 014110.

(5)

Endo, T.; Kerman, K.; Nagatani, N.; Hiepa, H. M.; Kim, D.-K.; Yonezawa, Y.; Nakano, K.; Tamiya, E. Multiple Label-Free Detection of Antigen− Antibody Reaction Using Localized Surface Plasmon Resonance-Based Core− Shell Structured Nanoparticle Layer Nanochip. Anal. Chem. 2006, 78 (18), 6465–6475.

(6)

Yang, S.; Slotcavage, D.; Mai, J. D.; Guo, F.; Li, S.; Zhao, Y.; Lei, Y.; Cameron, C. E.; Huang, T. J. Electrochemically Created Highly Surface Roughened Ag Nanoplate Arrays for SERS Biosensing Applications. J. Mater. Chem. C Mater. Opt. Electron. Devices 2014, 2 (39), 8350.

(7)

Pregibon, D. C.; Toner, M.; Doyle, P. S. Multifunctional Encoded Particles for HighThroughput Biomolecule Analysis. Science 2007, 315 (5817), 1393–1396.


Page 29 of 39

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


(8)

Lee, H.-L.; Chen, S.-C. Microchip Capillary Electrophoresis with Electrochemical Detector for Precolumn Enzymatic Analysis of Glucose, Creatinine, Uric Acid and Ascorbic Acid in Urine and Serum. Talanta 2004, 64 (3), 750–757.

(9)

Daniels, J. S.; Pourmand, N. Label‐free Impedance Biosensors: Opportunities and Challenges. Electroanalysis 2007, 19 (12), 1239–1257.

(10) Dykstra, P. H.; Roy, V.; Byrd, C.; Bentley, W. E.; Ghodssi, R. Microfluidic Electrochemical Sensor Array for Characterizing Protein Interactions with Various Functionalized Surfaces. Anal. Chem. 2011, 83 (15), 5920–5927. (11) Narakathu, B. B.; Atashbar, M.; Bejcek, B. Improved Detection Limits of Toxic Biochemical Species Based on Impedance Measurements in Electrochemical Biosensors. Biosens. Bioelectron. 2010, 26 (2), 923–928. (12) Varshney, M.; Li, Y.; Srinivasan, B.; Tung, S. A Label-Free, Microfluidics and Interdigitated Array Microelectrode-Based Impedance Biosensor in Combination with Nanoparticles Immunoseparation for Detection of Escherichia Coli O157: H7 in Food Samples. Sens. Actuators B Chem. 2007, 128 (1), 99–107. (13) Ali, M. A.; Mondal, K.; Jiao, Y.; Oren, S.; Xu, Z.; Sharma, A.; Dong, L. Microfluidic Immuno-Biochip for Detection of Breast Cancer Biomarkers Using Hierarchical Composite of Porous Graphene and Titanium Dioxide Nanofibers. ACS Appl. Mater. Interfaces 2016, 8 (32), 20570–20582. (14) Huang, J.; Lee, I.; Luo, X.; Cui, X. T.; Yun, M. Shadow Masking for Nanomaterial-Based Biosensors Incorporated with a Microfluidic Device. Biomed. Microdevices 2013, 15 (3), 531–537.



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

(15) Mondal, K.; Ali, M. A.; Agrawal, V. V.; Malhotra, B. D.; Sharma, A. Highly Sensitive Biofunctionalized Mesoporous Electrospun TiO2 Nanofiber Based Interface for Biosensing. ACS Appl. Mater. Interfaces 2014, 6 (4), 2516–2527. (16) Huang, Y.; Duan, X.; Wei, Q.; Lieber, C. M. Directed Assembly of One-Dimensional Nanostructures into Functional Networks. Science 2001, 291 (5504), 630–633. (17) Min, J.; Baeumner, A. J. Characterization and Optimization of Interdigitated Ultramicroelectrode Arrays as Electrochemical Biosensor Transducers. Electroanalysis 2004, 16 (9), 724–729. (18) Ali, M. A.; Solanki, P. R.; Patel, M. K.; Dhayani, H.; Agrawal, V. V.; John, R.; Malhotra, B. D. A Highly Efficient Microfluidic Nano Biochip Based on Nanostructured Nickel Oxide. Nanoscale 2013, 5 (7), 2883–2891. (19) Ahmad, M.; Pan, C.; Luo, Z.; Zhu, J. A Single ZnO Nanofiber-Based Highly Sensitive Amperometric Glucose Biosensor. J. Phys. Chem. C 2010, 114 (20), 9308–9313. (20) Ahmad, M.; Pan, C.; Zhu, J. Electrochemical Determination of L-Cysteine by an Elbow Shaped, Sb-Doped ZnO Nanowire-Modified Electrode. J. Mater. Chem. 2010, 20 (34), 7169–7174. (21) Lin, S.-Y.; Chang, S.-J.; Hsueh, T.-J. ZnO Nanowires Modified with Au Nanoparticles for Nonenzymatic Amperometric Sensing of Glucose. Appl. Phys. Lett. 2014, 104 (19), 193704. (22) Ali, M. A.; Mondal, K.; Singh, C.; Malhotra, B. D.; Sharma, A. Anti-Epidermal Growth Factor Receptor Conjugated Mesoporous Zinc Oxide Nanofibers for Breast Cancer Diagnostics. Nanoscale 2015, 7 (16), 7234–7245.


Page 30 of 39

Page 31 of 39

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


(23) Vasudev, A.; Kaushik, A.; Bhansali, S. Electrochemical Immunosensor for Label Free Epidermal Growth Factor Receptor (EGFR) Detection. Biosens. Bioelectron. 2013, 39 (1), 300–305. (24) Das, M.; Sumana, G.; Nagarajan, R.; Malhotra, B. Application of Nanostructured ZnO Films for Electrochemical DNA Biosensor. Thin Solid Films 2010, 519 (3), 1196–1201. (25) Wei, Y.; Li, Y.; Liu, X.; Xian, Y.; Shi, G.; Jin, L. ZnO Nanorods/Au Hybrid Nanocomposites for Glucose Biosensor. Biosens. Bioelectron. 2010, 26 (1), 275–278. (26) Tak, M.; Gupta, V.; Tomar, M. Flower-like ZnO Nanostructure Based Electrochemical DNA Biosensor for Bacterial Meningitis Detection. Biosens. Bioelectron. 2014, 59, 200– 207. (27) Hong, H.; Shi, J.; Yang, Y.; Zhang, Y.; Engle, J. W.; Nickles, R. J.; Wang, X.; Cai, W. Cancer-Targeted Optical Imaging with Fluorescent Zinc Oxide Nanowires. Nano Lett. 2011, 11 (9), 3744–3750. (28) Paul, K. B.; Kumar, S.; Tripathy, S.; Vanjari, S. R. K.; Singh, V.; Singh, S. G. A Highly Sensitive Self Assembled Monolayer Modified Copper Doped Zinc Oxide Nanofiber Interface for Detection of Plasmodium Falciparum Histidine-Rich Protein-2: Targeted towards Rapid, Early Diagnosis of Malaria. Biosens. Bioelectron. 2016, 80, 39–46. (29) Gooding, J. J. Nanostructuring Electrodes with Carbon Nanotubes: A Review on Electrochemistry and Applications for Sensing. Electrochimica Acta 2005, 50 (15), 3049– 3060. (30) Liu, A. Towards Development of Chemosensors and Biosensors with Metal-Oxide-Based Nanowires or Nanotubes. Biosens. Bioelectron. 2008, 24 (2), 167–177.



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

Page 32 of 39

(31) Wang, Y.; Yu, L.; Wang, J.; Lou, L.; Du, W.; Zhu, Z.; Peng, H.; Zhu, J. A Novel LLactate Sensor Based on Enzyme Electrode Modified with ZnO Nanoparticles and Multiwall Carbon Nanotubes. J. Electroanal. Chem. 2011, 661 (1), 8–12. (32) Haghighi,

B.;

Bozorgzadeh,

S.

Fabrication

of

a

Highly

Sensitive

Electrochemiluminescence Lactate Biosensor Using ZnO Nanoparticles Decorated Multiwalled Carbon Nanotubes. Talanta 2011, 85 (4), 2189–2193. (33) Zhang, Z.; Shao, C.; Li, X.; Wang, C.; Zhang, M.; Liu, Y. Electrospun Nanofibers of PType NiO/N-Type ZnO Heterojunctions with Enhanced Photocatalytic Activity. ACS Appl. Mater. Interfaces 2010, 2 (10), 2915–2923. (34) Tak, M.; Gupta, V.; Tomar, M. Zinc Oxide–multiwalled Carbon Nanotubes Hybrid Nanocomposite Based Urea Biosensor. J. Mater. Chem. B 2013, 1 (46), 6392–6401. (35) Palanisamy, S.; Cheemalapati, S.; Chen, S.-M. Highly Sensitive and Selective Hydrogen Peroxide Biosensor Based on Hemoglobin Immobilized at Multiwalled Carbon Nanotubes–zinc Oxide Composite Electrode. Anal. Biochem. 2012, 429 (2), 108–115. (36) Fang, B.; Zhang, C.; Zhang, W.; Wang, G. A Novel Hydrazine Electrochemical Sensor Based on a Carbon Nanotube-Wired ZnO Nanoflower-Modified Electrode. Electrochimica Acta 2009, 55 (1), 178–182. (37) Paul, K. B.; Singh, V.; Vanjari, S. R. K.; Singh, S. G. One Step Biofunctionalized Electrospun Multiwalled Carbon Nanotubes Embedded Zinc Oxide Nanowire Interface for Highly Sensitive Detection of Carcinoma Antigen-125. Biosens. Bioelectron. 2017, 88, 144–152.


Page 33 of 39

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


(38) Paul, K.; Panigrahi, A.; Singh, V.; Singh, S. A Multi-Walled Carbon Nanotube-Zinc Oxide Nanofiber Based Flexible Chemiresistive Biosensor for Malaria Biomarker Detection. The Analyst 2017. (39) Ali, M. A.; Srivastava, S.; Solanki, P. R.; Reddy, V.; Agrawal, V. V.; Kim, C.; John, R.; Malhotra, B. D. Highly Efficient Bienzyme Functionalized Nanocomposite-Based Microfluidics Biosensor Platform for Biomedical Application. Sci. Rep. 2013, 3, 2661. (40) Ali, M. A.; Srivastava, S.; Mondal, K.; Chavhan, P. M.; Agrawal, V. V.; John, R.; Sharma, A.; Malhotra, B. D. A Surface Functionalized Nanoporous Titania Integrated Microfluidic Biochip. Nanoscale 2014, 6 (22), 13958–13969. (41) Ali, M. A.; Solanki, P. R.; Srivastava, S.; Singh, S.; Agrawal, V. V.; John, R.; Malhotra, B. D. Protein Functionalized Carbon Nanotubes-Based Smart Lab-on-a-Chip. ACS Appl. Mater. Interfaces 2015, 7 (10), 5837–5846. (42) Martinez, A. W.; Phillips, S. T.; Whitesides, G. M.; Carrilho, E. Diagnostics for the Developing World: Microfluidic Paper-Based Analytical Devices. 2009. (43) Shafiee, H.; Asghar, W.; Inci, F.; Yuksekkaya, M.; Jahangir, M.; Zhang, M. H.; Durmus, N. G.; Gurkan, U. A.; Kuritzkes, D. R.; Demirci, U. Paper and Flexible Substrates as Materials for Biosensing Platforms to Detect Multiple Biotargets. Sci. Rep. 2015, 5, 8719. (44) Cunningham, J. C.; Brenes, N. J.; Crooks, R. M. Paper Electrochemical Device for Detection of DNA and Thrombin by Target-Induced Conformational Switching. Anal. Chem. 2014, 86 (12), 6166–6170. (45) Ge, S.; Liu, W.; Ge, L.; Yan, M.; Yan, J.; Huang, J.; Yu, J. In Situ Assembly of Porous Au-Paper Electrode and Functionalization of Magnetic Silica Nanoparticles with HRP via



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

Page 34 of 39

Click Chemistry for Microcystin-LR Immunoassay. Biosens. Bioelectron. 2013, 49, 111– 117. (46) Lankelma, J.; Nie, Z.; Carrilho, E.; Whitesides, G. M. Paper-Based Analytical Device for Electrochemical Flow-Injection Analysis of Glucose in Urine. Anal. Chem. 2012, 84 (9), 4147–4152. (47) Nie, Z.; Nijhuis, C. A.; Gong, J.; Chen, X.; Kumachev, A.; Martinez, A. W.; Narovlyansky, M.; Whitesides, G. M. Electrochemical Sensing in Paper-Based Microfluidic Devices. Lab. Chip 2010, 10 (4), 477–483. (48) Shanmugam, N. R.; Muthukumar, S.; Prasad, S. Ultrasensitive and Low-Volume Point-ofCare Diagnostics on Flexible Strips–a Study with Cardiac Troponin Biomarkers. Sci. Rep. 2016, 6. (49) Cox, F. E. History of the Discovery of the Malaria Parasites and Their Vectors. Parasit. Vectors 2010, 3 (1), 5. (50) World Health Organization. World Malaria Report 2015; World Health Organization, 2016. (51) Rich, S. M.; Leendertz, F. H.; Xu, G.; LeBreton, M.; Djoko, C. F.; Aminake, M. N.; Takang, E. E.; Diffo, J. L.; Pike, B. L.; Rosenthal, B. M. The Origin of Malignant Malaria. Proc. Natl. Acad. Sci. 2009, 106 (35), 14902–14907. (52) Kifude, C. M.; Rajasekariah, H. G.; Sullivan, D. J.; Stewart, V. A.; Angov, E.; Martin, S. K.; Diggs, C. L.; Waitumbi, J. N. Enzyme-Linked Immunosorbent Assay for Detection of Plasmodium Falciparum Histidine-Rich Protein 2 in Blood, Plasma, and Serum. Clin. Vaccine Immunol. 2008, 15 (6), 1012–1018.


Page 35 of 39

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


(53) Demirev, P. A. Mass Spectrometry for Malaria Diagnosis. Expert Rev. Mol. Diagn. 2004, 4 (6), 821–829. (54) Fontecha, G. A.; Mendoza, M.; Banegas, E.; Poorak, M.; De Oliveira, A. M.; Mancero, T.; Udhayakumar, V.; Lucchi, N. W.; Mejia, R. E. Comparison of Molecular Tests for the Diagnosis of Malaria in Honduras. Malar. J. 2012, 11 (1), 119. (55) Hanson, J.; Lam, S. W.; Mahanta, K. C.; Pattnaik, R.; Alam, S.; Mohanty, S.; Hasan, M. U.; Hossain, A.; Charunwatthana, P.; Chotivanich, K. Relative Contributions of Macrovascular and Microvascular Dysfunction to Disease Severity in Falciparum Malaria. J. Infect. Dis. 2012, 206 (4), 571–579. (56) Yatsushiro, S.; Yamamura, S.; Yamaguchi, Y.; Shinohara, Y.; Tamiya, E.; Horii, T.; Baba, Y.; Kataoka, M. Rapid and Highly Sensitive Detection of Malaria-Infected Erythrocytes Using a Cell Microarray Chip. PLoS One 2010, 5 (10), e13179. (57) Poon, L. L.; Wong, B. W.; Ma, E. H.; Chan, K. H.; Chow, L. M.; Abeyewickreme, W.; Tangpukdee, N.; Yuen, K. Y.; Guan, Y.; Looareesuwan, S. Sensitive and Inexpensive Molecular Test for Falciparum Malaria: Detecting Plasmodium Falciparum DNA Directly from Heat-Treated Blood by Loop-Mediated Isothermal Amplification. Clin. Chem. 2006, 52 (2), 303–306. (58) Campuzano-Zuluaga, G.; Hänscheid, T.; Grobusch, M. P. Automated Haematology Analysis to Diagnose Malaria. Malar. J. 2010, 9 (1), 346. (59) Shapiro, H. M.; Mandy, F. Cytometry in Malaria: Moving beyond Giemsa. Cytometry A 2007, 71 (9), 643–645.



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

Page 36 of 39

(60) Koo, W.-T.; Jang, J.-S.; Choi, S.-J.; Cho, H.-J.; Kim, I.-D. Metal-Organic Framework Templated Catalysts: Dual Sensitization of PdO-ZnO Composite on Hollow SnO2 Nanotubes for Selective Acetone Sensors. ACS Appl. Mater. Interfaces 2017. (61) Li, Q.; Li, Y.; Zhang, X.; Chikkannanavar, S. B.; Zhao, Y.; Dangelewicz, A. M.; Zheng, L.; Doorn, S. K.; Jia, Q.; Peterson, D. E. Structure‐dependent Electrical Properties of Carbon Nanotube Fibers. Adv. Mater. 2007, 19 (20), 3358–3363. (62) Behler, K.; Osswald, S.; Ye, H.; Dimovski, S.; Gogotsi, Y. Effect of Thermal Treatment on the Structure of Multi-Walled Carbon Nanotubes. J. Nanoparticle Res. 2006, 8 (5), 615–625. (63) Thompson, C. S.; Abate, A. R. Adhesive-Based Bonding Technique for PDMS Microfluidic Devices. Lab. Chip 2013, 13 (4), 632–635. (64) Liu, Y.; Zhang, H.; An, X.; Gao, C.; Zhang, Z.; Zhou, J.; Zhou, M.; Xie, E. Effect of Al Doping on the Visible Photoluminescence of ZnO Nanofibers. J. Alloys Compd. 2010, 506 (2), 772–776. (65) Ramanathan, T.; Fisher, F.; Ruoff, R.; Brinson, L. Amino-Functionalized Carbon Nanotubes for Binding to Polymers and Biological Systems. Chem. Mater. 2005, 17 (6), 1290–1295.

(66) Kumar, S.; Sen, A.; Kumar, S.; Augustine, S.; Yadav, B. K.; Mishra, S.; Malhotra, B. D. Polyaniline Modified Flexible Conducting Paper for Cancer Detection. Appl. Phys. Lett. 2016, 108 (20), 203702.


Page 37 of 39

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


(67) Bardea, A.; Katz, E.; Willner, I. Probing Antigen–antibody Interactions on Electrode Supports by the Biocatalyzed Precipitation of an Insoluble Product. Electroanalysis 2000, 12 (14), 1097–1106. (68) Muhammad, A.; Yusof, N. A.; Hajian, R.; Abdullah, J. Construction of an Electrochemical Sensor Based on Carbon Nanotubes/Gold Nanoparticles for Trace Determination of Amoxicillin in Bovine Milk. Sensors 2016, 16 (1), 56. (69) Kano, K.; Uno, B. Surface-Redox Reaction Mechanism of Quinones Adsorbed on BasalPlane Pyrolytic Graphite Electrodes. Anal. Chem. 1993, 65 (8), 1088–1093. (70) Mulchandani, P.; Chen, W.; Mulchandani, A. Flow Injection Amperometric Enzyme Biosensor for Direct Determination of Organophosphate Nerve Agents. Environ. Sci. Technol. 2001, 35 (12), 2562–2565. (71) Chrouda, A.; Sbartai, A.; Bessueille, F.; Renaud, L.; Maaref, A.; Jaffrezic-Renault, N. Electrically Addressable Deposition of Diazonium-Functionalized Antibodies on BoronDoped Diamond Microcells for the Detection of Ochratoxin A. Anal. Methods 2015, 7 (6), 2444–2451. (72) Medina-Sánchez, M.; Miserere, S.; Morales-Narváez, E.; Merkoçi, A. On-Chip MagnetoImmunoassay for Alzheimer’s Biomarker Electrochemical Detection by Using Quantum Dots as Labels. Biosens. Bioelectron. 2014, 54, 279–284. (73) Chikkaveeraiah, B. V.; Mani, V.; Patel, V.; Gutkind, J. S.; Rusling, J. F. Microfluidic Electrochemical Immunoarray for Ultrasensitive Detection of Two Cancer Biomarker Proteins in Serum. Biosens. Bioelectron. 2011, 26 (11), 4477–4483.



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

(74) Wu, Y.; Xue, P.; Hui, K. M.; Kang, Y. A Paper-Based Microfluidic Electrochemical Immunodevice Integrated with Amplification-by-Polymerization for the Ultrasensitive Multiplexed Detection of Cancer Biomarkers. Biosens. Bioelectron. 2014, 52, 180–187. (75) Wang, Y.; Xu, H.; Luo, J.; Liu, J.; Wang, L.; Fan, Y.; Yan, S.; Yang, Y.; Cai, X. A Novel Label-Free Microfluidic Paper-Based Immunosensor for Highly Sensitive Electrochemical Detection of Carcinoembryonic Antigen. Biosens. Bioelectron. 2016, 83, 319–326. (76) Eletxigerra, U.; Martinez-Perdiguero, J.; Merino, S. Disposable Microfluidic ImmunoBiochip for Rapid Electrochemical Detection of Tumor Necrosis Factor Alpha Biomarker. Sens. Actuators B Chem. 2015, 221, 1406–1411. (77) Vasudev, A.; Kaushik, A.; Tomizawa, Y.; Norena, N.; Bhansali, S. An LTCC-Based Microfluidic System for Label-Free, Electrochemical Detection of Cortisol. Sens. Actuators B Chem. 2013, 182, 139–146. (78) Das, J.; Kelley, S. O. Protein Detection Using Arrayed Microsensor Chips: Tuning Sensor Footprint to Achieve Ultrasensitive Readout of CA-125 in Serum and Whole Blood. Anal. Chem. 2011, 83 (4), 1167–1172.


Page 38 of 39

Page 39 of 39

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


TOC 107x36mm (150 x 150 DPI)

ACS Paragon Plus Environment

Nonlithographic Fabrication of Plastic-Based Nanofibers Integrated

Recommend Documents