Nanostructured Electrochemical Biosensors for Label-Free Detection

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Nanostructured electrochemical biosensors for labelfree detection of water- and food-borne pathogens Nekane Reta, Christopher P Saint, Andrew Michelmore, Beatriz Prieto-Simon, and Nicolas H. Voelcker ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b13943 • Publication Date (Web): 25 Jan 2018 Downloaded from http://pubs.acs.org on January 26, 2018

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ACS Applied Materials & Interfaces

Nanostructured electrochemical biosensors for label-free detection of water- and food-borne pathogens Nekane Reta,a Christopher P. Saint,b Andrew Michelmore,a,c Beatriz Prieto-Simon,*,a,d and Nicolas H. Voelcker

a

*,a,d,e

Future Industries Institute, University of South Australia, Mawson Lakes, South Australia 5095, Australia b

Natural & Built Environments Research Centre, School of Natural & Built Environments, University of South Australia, Mawson Lakes, South Australia 5095, Australia

c

School of Engineering, University of South Australia, Mawson Lakes, South Australia 5095, Australia d

Monash Institute of Pharmaceutical Sciences, Monash University, Parkville, Vic 3052, Australia

e

Melbourne Centre for Nanofabrication, Victorian Node of the Australian National Fabrication Facility, Clayton, Victoria, 3168, Australia

Corresponding authors: Prof. Nicolas H. Voelcker, Dr. Beatriz Prieto-Simon Drug Delivery, Disposition and Dynamics, Monash Institute of Pharmaceutical Sciences, Monash University, 381 Royal Parade, Parkville, Victoria, 3052, Australia Tel: +61 3 99039230

email: [email protected] [email protected]

Keywords: pathogen detection, label-free detection, environmental monitoring, food safety, nanoscale materials,

nanochannels,

electrochemical

ACS Paragon Plus Environment

biosensing

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Abstract The emergence of nanostructured materials has opened new horizons in the development of next generation biosensors. Being able to control the design of the electrode interface at the nanoscale combined with the intrinsic characteristics of the nanomaterials engenders novel biosensing platforms with improved capabilities. The purpose of this review is to provide a comprehensive and critical overview of the latest trends in emerging nanostructured electrochemical biosensors. A detailed description

and

discussion

of

recent

approaches

to

construct

label-free

electrochemical

nanostructured electrodes is given with special focus on pathogen detection for environmental monitoring and food safety. This includes the use of nanoscale materials such as nanotubes, nanowires, nanoparticles and nanosheets, as well as porous nanostructured materials including nanoporous anodic alumina, mesoporous silica, porous silicon and polystyrene nanochannels. These platforms may pave the way towards the development of point-of-care portable electronic devices for applications ranging from environmental analysis to biomedical diagnostics.

2 ACS Paragon Plus Environment

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1 Introduction

The presence of human pathogens in water and food is a major concern due to their rapid growth and harmful effects on health. The World Health Organization reported in 2007 that worldwide waterand food-borne infections account for about 4 billion episodes of diarrhea per year, resulting in about 1

1 to 2 million deaths . Pathogens responsible for water- and food-borne outbreaks include bacteria, viruses and toxins 2-6. Figure 1 shows images of the most relevant pathogens that have been related to cause adverse effects on human health. Bacteria such as Salmonella typhi (S. typhi), Escherichia coli (E. coli), Listeria monocytogenes (L. monocytogenes), and Campylobacter are generally responsible for most illnesses, in conjugation with viruses such as hepatitis (A, B) and norovirus (NoV). They can be found on ready-to eat food, raw or undercooked meat, shellfish and in poor sanitation environments. Amongst toxins, mycotoxins and enterotoxins have been associated with diseases. Mycotoxins are secondary metabolites produced by fungi and fungi-like organisms, 7

aflatoxin being one of the most harmful . Aflatoxins (AFs), generated by certain Aspergillus, Penicillium and Fusarium species, can be found in crops (i.e. corn, wheat) and nuts 8. Other mycotoxins that can cause adverse effects are ochratoxins, sterigmatocystin, trichlothecenes, zearelenone and fumonisins 9. Enterotoxins, peptides generally released by Gram-positive bacteria, have been associated with food poisoning and toxic shock syndrome

10

. Enterotoxins produced by

Staphylococcus aureus (S. aureus) are responsible for the most food poisoning cases as a result of contaminated food ingestion (dairy and meat products) and improperly stored food products

11-12

.

Table 1 summarizes the sources of the most significant water- and food-borne pathogens, the number of microorganisms or amount of toxin required to cause an infection, as well as the associated symptoms.



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Figure 1. Electron microscopy images of (A) S. typhi, (B) E. coli (C) L. monocytogenes (D) Campylobacter, (E) hepatitis A virus, (F) NoV, (G) Aspergillus parasiticus (A. parasiticus) and (H) S. aureus. Reproduced with permission from: (A) ref 13. Copyright 2006 Elsevier, (B, H) ref 14. Copyright 2000 American Association for the Advancement of Science, (C) ref 15. Copyright 2015 Multidisciplinary Digital Publishing Institute, (D) ref 16. Copyright 2011 American Society for Microbiology, (E) ref 17. Copyright 2016 Pixnio, (F) ref 18. Copyright 2010 PLOS, (G) ref 19. Copyright 2003 Oxford Academics.

Conventional methods of pathogen detection rely on specific microbiological and biochemical identification. They are based on culture and colony counting, immunological interactions and the polymerase chain reaction (PCR). Among all of these methods, culture and colony-based confirmation assays are the most accurate and reliable for pathogen detection. Although these methods are highly sensitive and often allow the detection of a single bacterium, they are time consuming and labor intensive as they may take up to 7-10 days to perform 3. Immunology-based methods rely on antibody and antigen interactions and have been widely used for the detection of bacterial cells, viruses and toxins 20. Those require shorter assay times compared to traditional culture methods. However, they still lack the ability to detect pathogens in real time and can be potentially affected by interferences. Additionally, the limited affinity and specificity of the antibody to the microorganism in some cases leads to low sensitivity and selectivity, respectively

21

. PCR techniques

have distinct advantages over culture and immunological methods, including high specificity,


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sensitivity and accuracy. In spite of their advantages, PCR requires a skilled operator and can be expensive to perform

22

. Thus, the development of new, less expensive tools providing equally reliable

results over much shorter time frames is key to preventing health and safety problems caused by pathogen contamination in food or drinking water. Electrochemical biosensors provide high sensitivity, selectivity and robustness at low cost

23-24

. When

a specific sensing strategy and detection technique are combined, biosensors can be developed for label-free detection, which presents shorter analysis time and simplicity over labelled strategies

25

.

Another important feature is their ability to be miniaturized and therefore potential for portable in-situ analysis. All these characteristics taken together with the unique properties of nanostructured materials provide an attractive means for the development of novel platforms with improved sensitivity for water- and food-borne pathogens 26. Nanostructured materials have emerged in the last decade as powerful materials for electrode fabrication in the biosensing field

27

. The novel features of nanomaterials have enabled significant

improvements in the sensing capabilities of biosensors.

28

. Here, the latest trends to design label-free

nanostructured electrodes will be discussed. This includes the use of nanoscale materials including nanotubes, nanowires, nanoparticles and nanosheets, as well as porous nanostructures featuring arrays of nanochannels such as nanoporous anodic alumina, mesoporous silicon, porous silicon and polystyrene nanochannels. The most relevant strategies for the detection of water- and food-borne contamination will also be highlighted.


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Table 1. Common water- and food-borne pathogens.

Pathogen

Source

Infectious dose

Symptoms

Ref.

Stomach pain, diarrhea,

3, 29-33

(no microorganisms) S. typhi

● Raw milk and dairy products, raw or undercooked

15 – 20

poultry and meat, seafood, chocolate and salads.

nausea, headache, fever, chills.

● Lack of safe water, poor sanitation and hygiene. E. coli O157:H7

● Raw milk and dairy products, raw or undercooked eggs,

< 10

poultry and meat, seafood and leafy vegetables.

Stomach pain, diarrhea,

3, 33-34

nausea, headache, fever, chills.

● Lack of safe water, poor sanitation and hygiene. L. monocytogenes

● Raw milk, soft cheese, raw poultry and meat, seafood

< 1000

and leafy vegetables.

Fever, chills, headache,

3, 31, 35

diarrhea.

● Lack of safe water, poor sanitation and hygiene. Campylobacter

● Raw milk, raw or undercooked meat and poultry.

400 – 500

Fever,

headache,

diarrhea, ●Lack of safe water, poor sanitation and hygiene. abdominal pain.


nausea,

3, 36-37

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Pathogen

Source

Infectious dose

Symptoms

Ref.

Fever, malaise, anorexia,

3, 38-40

o

(n microorganisms)

Hepatitis A virus

● Milk and dairy products, shellfish, fruits, vegetables and

100 – 1000

iced drinks.

nausea, jaundice.

● Lack of safe water, poor sanitation and hygiene.

NoV

● Raw oysters or shellfish, berries and salads.

Median

dose

that

Nausea,

vomiting,

causes infection in 50%

diarrhea,

stomach

of individuals (ID50) =

crumps, muscle cramps.

41-48

● Contaminated water, ice and frozen products.

AFs


18 virus particles

● Corn, cereals, nuts (pistachio, peanut), spices (black

< 15 µg/kg, per total

Hemorrhage, acute liver

pepper, chilies), dried fruits and in a lesser amount in milk

AFs (B1, B2, G1, G2)

damage,

and dairy products.

49-53

edema,

abdominal pain.


7



Pathogen

Source

Infectious dose

Page 8 of 59

Symptoms

Ref.

o

(n microorganisms)

Enterotoxin from S.

● Raw milk, cheese, meat products, salads, pastries and

aureus

custards.

20 – 100 ng

Nausea,

diarrhea,

10-11, 54-59

abdominal pain, cramps.

● Food improperly handled and stored.

8


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2 Electrochemical biosensors based on nanoscale materials The use of nanoscale materials in the design of electrochemical biosensors is a rapidly expanding area. Nanoscale materials such as nanotubes, nanowires, nanoparticles and nanosheets have been extensively incorporated in electrode construction. This is attributed to their remarkable characteristics such as high surface-to-volume ratio and excellent electrical properties, which have led to the development of sensing platforms with outstanding performance. The following section will provide a comprehensive overview of the recent trends for the fabrication of these nanoscale biosensors. The latest approaches will be discussed in detail and the most significant platforms for water- and foodborne pathogen detection will be described.

2.1

Nanotubes

Nanotubes as building blocks in the design of electrochemical sensors have attracted much interest over the past few years. Nanotubes consisting of organic or inorganic materials and exhibiting conductive or semi-conductive properties have been prepared. Amongst all of them, carbon nanotubes (CNTs) have paved the way for promising electrochemical biosensors since their discovery in the 1990s. Other nanotube structures based on titanium dioxide have also been employed, to a lesser extent, in the electrochemical biosensing field. The following section will introduce the different types of CNTs and their structures, as well as different approaches employed for the design of labelfree CNT-based electrochemical biosensors. Some promising examples will be highlighted with special focus on pathogen detection. Finally, strategies employing titanium dioxide nanotubes for electrochemical biosensors will be described. CNTs can be divided into two classes: single-walled carbon nanotubes (SWCNTs), which consist of a single cylindrical hollow tube with 0.4 to 2 nm diameter, and multi-walled carbon nanotubes (MWCNTs), which are made of multiple concentric tubes 3.4 Å apart with 2 – 100 nm diameters

60

.

The electrical behavior varies between both types, where MWCNTs behave similar to metals and SWCNTs can feature both metallic and semi-conductor character depending on their geometric structure such as chirality and tube diameter

61

. The chirality is related to the angle at which the



Page 10 of 59

graphene sheet is rolled up to form a SWCNT and can be defined by a lattice vector with integers (n,m) (Figure 2). Nanotubes with metallic properties are in “armchair” configuration while “chiral” and “zigzag” structures are semiconductors. The diameter of the SWCNT also plays an important role in the electronic behavior, where small diameter nanotubes only possess conductive or semi-conductive properties. As the diameter of the tube increases the band gap tends to zero, resulting in a zero-gap semiconductor.

62

.The high electronic conductivity is particularly interesting for electrochemical

detection platforms as it enhances the electrocatalytic activity or the electron transfer reaction between redox species and the electrode

63

2

. Furthermore, their large surface area (1315 m /g)

64

provides a favorable environment to immobilize biomolecules, as well as, remarkable mechanical strength. All these features of CNTs have been exploited for the development of highly sensitive electrochemical platforms 65.

Figure 2. (A) Schematic of the folding procedure to form different configurations of SWCNTs from planar graphene sheet according to their chiral angle (ɸ). C is a roll-up vector that is defined as: C=na1+ma2, where a1 and a2 are the primitive lattice vectors of the hexagonal lattice and n and m are integers. (B) MWCNT structure with several concentric tubes. Reproduced with permission from ref 66. Copyright 2015 Frontiers.

Several strategies have been employed to design electrochemical biosensors incorporating CNTs. These include CNT-coated electrodes, CNT-composite electrodes, vertically-aligned CNT (VA-CNT)


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electrodes and CNT-based field-effect transistors (FET) (Figure 3)

67

. Coating or depositing CNTs

layers on conventional electrodes such as glassy carbon electrodes (GCE) has been done by spraying, drop casting or electrodeposition techniques

68

. When using this method, it is extremely 69

important to obtain a uniform CNT coating in order to enhance the electrocatalytic activity

. The

performance of CNT-composite electrodes also depends on the homogeneity of the CNT dispersion within the binder employed. Coatings on electrodes are typically fabricated by mixing the CNTs with a polymer matrix using a variety of binders (mineral oil, Teflon or epoxy resins)

63

. A way to obtain an

effective CNT coating from an electrochemical perspective is by using VA-CNTs

65

. Electrodes

modified with VA-CNTs have shown excellent performance due to an increase in the charge transfer rate when compared to the CNTs randomly oriented

70

. VA-CNTs can be prepared by physical or

chemical methods. Physical methods are based on growing CNTs directly on certain substrates employing template-assisted or template-free procedures

71

, while chemical methods rely on

assembling CNTs by covalent bonds, metal-assisted chelation and electrostatic interaction

72

. CNTs

have also been widely integrated in FETs, mainly via chemical vapor deposition (CVD). Typically, these devices consist of a single CNT or a network of CNTs acting as a conduction channel between a source and a drain electrode, and a change in conductivity upon the interaction of the biomolecule with the CNT surface is the common biosensing mechanism 73.

Figure 3. Strategies to fabricate CNT-based biosensors: (A) CNT-coated electrode, (B) CNTcomposite, (C) VA-CNTs and (D) CNT-based FET.



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Modifications of the CNT-based electrodes to attach biomolecules as recognition elements have been perfomed by non-covalent and covalent routes

74-75

. Non-covalent approaches include physical

adsorption or entrapment on the CNTs using polymer hydrogels and sol-gel chemistry. Covalent attachments are generally performed by conjugating the carboxyl groups on oxidized CNTs to form amide bonds or by non-selective attack using highly reactive species such as nitrenes and aryl diazonium salts 66. Zelada-Guillen developed highly sensitive potentiometric SWCNT-based sensors modified with aptamers for real-time measurements of different bacteria, including E. coli typhi

76

, S. aureus

77

and S.

78

. These aptasensors were fabricated by depositing a 30 µm-thick SWCNT layer on a GCE by

spray deposition. Then, an aptamer specific to one of the target bacteria mentioned above was covalently bound to the carboxylic groups of the external CNT walls via carbodiimide chemistry. The sensing principle is based on ion-to-electron transduction. Binding of the bacteria to the aptamer provokes a conformational change in the aptamer, separating the negatively charged phosphate groups on the aptamer from the SWCNT walls. This causes a change in the charge of the SWCNT that is recorded as a change in the electrode potential. The aptasensor for S. typhi detection was the most sensitive, being able to detect a single colony-forming unit (cfu) in less than a minute 78. Figure 4 shows the potentiometric response of the sensor when it was exposed to a stepwise concentration of 6

S. typhi from 0.2 cfu/mL to 10

cfu/mL and the corresponding linear relation between the

electromotive force (EMF) measured between the working and reference electrodes and the logarithm of S. typhi concentration. The slope greatly decreased after 103 cfu/mL, reaching a plateau at 106 cfu/mL, due to possible saturation of the available binding sites. E. coli and Lactobacillus casei, showed no significant potentiometric response, demonstrating good selectivity of the sensor. Moreover, the aptasensor was regenerated by incubation in 2 M NaCl for 30 min, and was stable over three months. This platform represents a very promising development towards a point-of-analysis electrochemical device.


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Figure 4. (A) Potentiometric response of the aptasensor for the increasing concentration of S. typhi bacteria, and (B) corresponding calibration curve. Reproduced with permission from ref 78. Copyright 2009 ACS Wiley-VCH.

Andrade et al. developed an impedimetric biosensor for Gram-negative bacteria such as Klebsiella pneumoniae (K. pneumoniae) and E. coli

79

. In this approach, an Au electrode was functionalized with

cysteine to introduce amine groups by a self-assembled monolayer (SAM). The carboxylic groups of the

MWCNTs

were

activated

by

1-(3-dimethylaminopropyl)-3-ethylcarbodiimide

(EDC)/N-

hydroxysuccinimide (NHS) to bind first to the amine groups on the Au electrode and then to clavanin A, an antimicrobial peptide used as a capture probe, forming stable amide bonds. The performance of these biosensors was studied using electrochemical impedance spectroscopy (EIS) in a threeelectrode cell containing the CNT-modified Au as working electrode (WE), platinum as auxiliary electrode (AE) and Ag/AgCl as reference electrode (RE). These impedance spectra were fitted to an equivalent circuit model that enabled the calculation of the charge-transfer resistance (RCT), the circuit element commonly employed to follow affinity binding. RCT increased with increasing concentration of both Gram-negative bacteria (E. coli and K. pneumoniae), while the tested Gram-positive bacteria (Enterococcus faecalis (E. faecalis) and Bacillus subtilis (B. subtilis) gave no significant change in the 2

5

biosensor response. The developed biosensors exhibited a linear working range from 10 to 10 and 102 to 106 cfu/mL for E. coli and K. pneumoniae, respectively, with 100 cfu/mL being the lowest concentration detected for both bacterial species.



Page 14 of 59

A CNT-chitosan biocomposite electrode was prepared to detect sterigmatocystin, a carcinogenic mycotoxin found in food 80. SWCNTs were dispersed in chitosan solution and cast on the surface of a Au electrode. Then, aflatoxin oxidase (AFO) enzymes were immobilized onto the composite via electrostatic and hydrophobic interactions. The CNT-chitosan biocomposite provided a favorable environment to maintain the native conformation and electrocatalytic activity of the enzyme, while enhancing the direct electron transfer between the enzyme and the electrode surface. The enzymatic sensor, based on the voltammetric detection of the enzymatic product H2O2, exhibited a limit of detection (LOD) of 3 ng/mL that was comparable to those obtained by chromatographic detection techniques but was achieved within a shorter time (~10 s). Moreover, this sensor exhibited good stability in a dry state at 4 °C for about 1 month with a retained enzymatic activity of 85.6 % and good reproducibility. Other sol-gel CNTs enzymatic composites have also been fabricated to detect mycotoxins 81-82. CNTs have also been widely integrated in FETs to target E. coli 86-87

and T7 bacteriophage

85

. Villazar et al.

86

83-85

, Salmonella infantis (S. infantis)

fabricated the selective CNT network-FET

immunosensor for S. infantis detection. The CNT networks were fabricated on a silicon dioxide layer by CVD. Both source and drain electrodes were screen-printed with Ag ink, and an Al layer on the back of the silicon was used as the gate electrode to monitor the effect of gate voltage on current flow upon the bacterial binding. Current decreased with increasing concentration of S. infantis from 100 to 500 cfu/mL, while no changes were observed upon incubation with Shigella bacterium. The LOD of the device was not determined, but considerable current changes were recorded after incubating the immunosensor with100 cfu/mL of S. infantis for 1 h. Further experiments would be required to determine if this device can detect S. infantis in real samples and if the working range could be further extended. Titanium dioxide or titania nanotubes (TNT) have also been employed to detect various analytes

88

.

Mandal et al. prepared screen-printed carbon electrodes (SPCEs) with TNT by the drop casting method, followed by the attachment of the antibody to detect penicillin-binding protein, a marker for methicillin-resistant S. aureus

89

. The TNT-modified immunosensor monitors protein-antibody


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interaction by measuring changes in current by cyclic voltammetry (CV), allowing detection of 1 ng/µL, and being unaffected by other proteins such as bovine serum albumin (BSA).

2.2

Nanowires

Progress and development using nanowires as biosensing platforms has rapidly grown over the last decade

90

. Nanowires are defined as one-dimensional fibril-like nanostructures of diameters ranging

from tens of nanometers to a couple of hundred nanometers

91

. Several types of nanowires including

those based on semiconductors (e.g. Si, InP, InAs, GaN), metals (e.g. Au) and metal oxides (e.g. TiO2, ZnO) have been employed to construct sensing platforms

92-93

. The excellent electrical

properties, tunable morphology and ease of functionalization of silicon nanowires (SiNWs) have led to the development of a wide range of SiNW-based biosensors over the years. Most of these biosensors use nanowires configured as FETs 73, 94. SiNW-based biosensors can be fabricated by bottom-up and top-down approaches, metal-catalyzed CVD and metal-assisted electroless etching being the most common representative techniques, respectively

90

. The perfomance of a SiNW biosensor is infuenced by several factors such as

nanowire diameter, carrier density and mobility, and surface chemistry 95. It has been found that small diameters and lower doping density result in more sensitive FET-based biosensors

96-97

. Thus, it is

particularly important to fabricate SiNWs with appropiate dimension and doping density in order to develop a highly sensitive biosensing platform. The surface of SiNWs can be modified by silanization or photochemical hydrosilylation in order to introduce functional groups that can be further reacted with different bioreceptors. Silanization is commonly employed to introduce amine-, thiol- or aldehydeterminated functional groups 90, while photochemical hydrosilylation is used to form stable Si-C bonds 98

. Several examples of SiNW-FET biosensors for both bacteria and viruses detection are discussed

below. Patolsky et al. fabricated a real-time SiNW-FET array for the electrical detection of both influenza-A virus and adenovirus

99

. This study demonstrated single virus detection with high selectivity. It

consisted of two nanowires modified with anti-influenza A and anti-adenovirus antibodies, respectively. Figure 5A shows the sensing principle consisting of monitoring the change on the SiNW



Page 16 of 59

surface upon the virus particle binding and release event. Briefly, specific binding of a single virus particle to the antibody produces a change in the conductance that is characteristic to the surface charge of the virus particle. Upon an increase in the pH, the virus particle releases from the SiNW surface and conductance returns to the baseline value. Figure 5B shows real-time conductance changes when the SiNW-based influenza immunosensor was incubated with a single influenza virus particle (1 and 4 arrows), while in Figure 5C no changes were seen upon non-specific virus incubation. Moreover, the capability of multiplexed detection was also demonstrated by parallel measurements of influenza A and adenovirus. This sensor thus shows promise as a device for the simultaneous detection of large numbers of different viral threats.

Figure 5. (A) Schematic of virus particle binding to and release from SiNW-based immunosensors (left), with the corresponding changes in conductance as a function of time (right). Conductance change of the anti-influenza antibody-modified SiNW when incubating with (B) influenza A virus particle and (C) paramyxovirus and adenovirus. Arrows 1 and 4 indicate introduction of the virus and the remaining arrows represent incubation with buffer. Reproduced with permission from ref 99. Copyright 2004 Proceedings of the National Academy of Sciences, U.S.A.


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Recently, Shen et al. also reported a highly selective immunosensor for influenza A virus detection using a SiNW-FET biosensor

100

. This platform was able to detect 29 virus particles/µL within minutes

in clinical exhaled breath condensate samples, without significant effect of other interfering viruses or particles. Another highly sensitive platform based on a SiNW-FET device was developed by Zhang et al. to detect Dengue serotype 2 (DEN-2) virus

101

. This biosensor was combined with the reverse-

transcription PCR (RT-PCR) technique to select and amplify a specific fragment of DEN-2 genome sequences. Here, the detection principle was based on DNA hybridization. Briefly, the SiNW surface was first silanized with (3-aminopropyl)trimethoxysilane (APTES) to immobilize a specific peptide nucleic acid (PNA), a neutral peptide to minimize repulsion and facilitate hybridization event with the complementary and negatively charged DNA sequence. Hybridization of PNA and a specific DNA fragment of DEN-2 caused an accumulation of negative charge on the surface, inducing an increase in the resistance of the nanowire. This SiNW sensor achieved low detection levels (10 fM of the amplicons) within 30 minutes. Although SiNWs are the most common type of nanowires for biosensing applications, wires made of TiO2 have also been employed for direct electrochemical detection of bacteria

102

. A TiO2 nanowire bundle-based immunosensor was connected to an Au

microelectrode for the sensitive impedimetric detection of L. monocytogenes. This platform exhibited good performance, with a LOD of 500 cfu/mL achieved within 1 h assay time and without significant changes in the impedance response with other foodborne pathogens such as E. coli O157:H7, S. typhi and S. aureus.

2.3

Nanoparticles

Nanoparticles (NPs) are generally defined as particles smaller than 100 nm in size. Their physical, chemical and electronic properties have attracted great interest in the construction of novel sensing platforms. These properties depend on the kind and number of atoms that form the particle

103

. Many

types of NPs such as metal, metal oxide and semiconductors have been used to design electrochemical biosensors

104

. Metal NPs (MNPs) such as Au (AuNPs), Ag (AgNPs) and Pt (PtNPs)

are the most frequently employed

105

. This is attributed to their excellent conductivity, large surface-to-

volume ratio, inert nature and biocompatibility. MNPs provide a suitable microenvironment for biomolecule immobilization retaining their biological activity, and promoting electron transfer between



the immobilized biomolecules and the electrode surface

Page 18 of 59

106-107

. Furthermore, they have been used as

labels or tags to amplify the signal of the biorecognition event, as well as, for multiplexed detection. However, this review focuses on label-free detection and thus this will not be covered here and has been reviewed elsewhere

28, 103, 108-109

. The following section will focus on describing different

approaches to design MNP-modified biosensors for label-free electrochemical detection. MNP morphology, assembly and modification play a significant role in the performance of the MNPbased biosensors

110

. In general, they are prepared by chemical reduction of the corresponding metal

salt in aqueous organic phase and in the presence of a stabilizer/surface protector. This surface protector binds to their surface to avoid their aggregation by improving stability and solubility, as well as, providing the desired charge and chemical groups

111-112

. Colloidal AuNPs are typically prepared

by reducing chloroauric acid with sodium acetate in aqueous media, and stabilized by ions, biological molecules and polymers

113

. Compared to AuNPs, the synthesis of monodisperse AgNPs has been

more challenging since they are prone to corrosion and aggregation in solution

1t4

. Reduction of silver

nitrate by sodium borohydride in the presence of different citrate, polymer and biological molecules, is the most common method to prepare AgNPs

115-117

. PtNPs are frequently synthesized by

hexachloroplatinic acid reduction with citric acid and modified with biological molecules to increase stability

118

. Although impressive progress has been made in the synthesis of MNPs, precise control

over monodispersity, morphology and surface chemistry remains challenging 110. MNPs have been directly assembled onto the electrode surface, or incorporated with other nanomaterials and polymers to form sophisticated electrode architectures. One of the preferred strategies to modify electrode surfaces with MNPs is by means of self-assembly. This is a simple and versatile approach to form highly ordered monolayers with different functional groups

110, 119

. The

majority of the self-assembled monolayers (SAMs) are based on the affinity between thiol/amine groups and noble metal surfaces, providing a high degree of control over the molecular architecture of the biorecognition interface

110

. Recent work from Liu et al. has demonstrated a stable and robust

AuNP-modified GCE using SAMs

120

. This resulted in excellent electron transfer ability and low

background signal. Scheme 1 shows the different strategies for the fabrication on AuNP-modified GCE for sensing applications. Briefly, a GCE was modified either with 4-aminophenyl or 4-thiophenol to introduce amine- or thiol- terminal groups, respectively. AuNPs were subsequently tethered to the


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surface, forming S-Au and NH-Au bonds. Moreover, the amine-terminated electrode was also modified to form diazonium groups, followed by AuNP immobilization to form phenyl-Au bonds (GCPh-AuNPs). Amongst all of them, the GC-Ph-AuNP electrode exhibited the highest stability in an aqueous environment, with less particle losses over time. Other SAM strategies using cysteamine have been used to immobilize AuNPs and AgNPs on Au electrodes 121-123.

Scheme 1. Fabrication of stable AuNP-modified GCEs via self-assembly. Reproduced with permission from ref 120. Copyright 2015 ACS Publications.

MNPs have also been incorporated on electrode surfaces by the layer-by-layer technique

124-126

. This

technique relies on the electrostatic interaction between anionic and cationic polyelectrolytes and provides highly ordered architectures with accurate control over the composition, number of layers, as well as, the thickness of the multilayer assemblies at nanoscale level 127. As an example, Sungwoo et al. developed a multilayer structure based on catalase-encapsulated AuNPs by electrostatic assembly of anionic poly(sodium 4-styrene sulfonate) and cationic poly(allylamine hydrochloride) polyelectrolyte 126

. This multilayer assembly allowed electrostatic charge reversal and structure modifications by



Page 20 of 59

adjusting the pH. Near the isoelectric point of catalase (pH 5.2), dispersed catalase-encapsulated AuNPs could be altered to form colloidal or network architectures. Besides, high loading of catalase, as well as, effective electron transfer, high catalytic activity toward H2O2 was achieved. To further enhance the electronic, electrochemical and mechanical properties of the electrode surface, MNPs have been combined with several nanomaterials and polymers

110,

128-129

.

Nanomaterials such as graphene and CNTs have become ideal candidates to pair with MNPs for the development of electrochemical biosensors. Wang et al. fabricated AuNP-modified graphene impedimetric immunosensors to detect E. coli O157:H7 in food samples

130

. AuNPs were deposited

on a graphene-modified nitrocellulose membrane filter and functionalized with streptavidin followed by biotinylated antibody immobilization and subsequent blocking with BSA. EIS was employed to detect E. coli O157:H7, exhibiting a wide linear detection range from 1.5 x 102 to 1.5 x 107 cfu/mL and a LOD of 150 cfu/mL in 30 min. Moreover, this platform was succesfully tested with E. coli O157:H7 contaminated food samples such as beef and cucumber, showing LOD of 1.5 x 104 cfu/mL and 1.5 x 3

10 cfu/mL, respectively. This platform also possessed a high tolerance to mechanical stress. Wang et al. reported an ultrasensitive AuNP-SWCNT composite biosensor in which SWCNTs arrays were coated with AuNPs by electrodeposition

131

. This composite was functionalized with DNA to

detect human hepatitis B and papilloma viruses and showed the ability to detect as low as 1 aM of complementary 21- and 24-base hepatitis B and papilloma DNA, respectively. The low detection limit was achieved due to the synergistic effect of AuNPs and CNTs. Other MNPs such as PtNPs are often combined with CNTs to further improve the surface-to-volume ratio and electrocatalytic activity

129

132

133

Here, PtNPs were embedded in CNTs arrays

or simply decorated the surface of the CNTs

PtNP-modified CNTs can also be entrapped in a sol-gel matrix on the electrode surface

.

.

134

.

Incorporation of MNPs with polymer matrices to form three-dimensional molecular networks offers the possibility to engineer biosensors with improved electrical and mechanical properties because they provide several advantages such as prevention of MNP oxidation and coalescence, and stability enhancement of the nanocomposite 135. Au and AgNPs are the most common MNPs entrapped in solgels including silica gels for electrochemical sensing 136-139. Commercial AuNP-coated SPCEs have been employed to detect murine norovirus (MNV)

140

. Here,

AuNPs were modified with a thiolated MNV-specific DNA aptamer. Binding of the virus to the


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immobilized aptamer caused a decrease in the current intensity that was measured by square wave voltammetry (SVW). These aptasensors achieved a LOD of 180 virus particles in buffer, and were unaffected by other gastrointestinal viruses including vesicular stomatitis virus and vaccinia virus and human serum albumin. A similar aptasensor was also reported by the same group for the detection of vaccinia virus 141.

2.4

Nanosheets

Nanosheets or layered materials are two-dimensional nanostructures with thicknesses from about 1 to 100 nm

142

. Graphene, a single-atom thick sheet of carbon atoms packed into a two-dimensional

hexagonal lattice, is the most commonly employed nanosheet for electrochemical sensing. Graphene exhibits distinct features that are very attractive for sensing, such as excellent electrical conductivity, 2

large surface area (2630 m /g) and remarkable mechanical strength

143-144

. The material can be

produced by mechanical exfoliation, epitaxial growth of silicon carbide, chemical reduction of graphite oxide and unzipping of CNTs

145-146

. Of these methods, graphite oxide reduction is one of the most

economically viable one for mass production

147

. This method offers the possibility of introducing

several functionalities to the electrode, as well as to form hybrid nanomaterials during the reduction procedure due to the ease of functionalization with carboxyl, epoxy, hydroxyl and carbonyl reactive groups

143

. The epoxy and hydroxyl groups are located on the faces of each graphene oxide (GO)

sheet, while the carboxyl groups are usually found at the edge

148

.

Strategies for incorporating graphene in biosensing are similar to those employed for CNTs, and include graphene-coated conventional electrodes, graphene composites and using graphene in FETbased devices

149

. Graphene sheets are usually deposited on GCE via the drop casting method

Hybridization with chitosan

154-156

, NPs

157-158

and a combination of both

159-160

162-166

.

to form graphene

nanocomposites for improved performance characteristics have been widely studied has also been incorporated in FET devices

150-153

161

. Graphene

and proven to be more sensitive than conventional

metallic electrodes and similar to SiNW-based FETs 167. Zhou et al. constructed a chemically reduced GO-modified GCE by drop casting to detect DNA via direct oxidation

168

. Here, the GO-modified GCE was incubated with single-stranded DNA (ssDNA) or

double stranded DNA (dsDNA) and the oxidation currents of DNA nucleobases (adenine, thymine,



Page 22 of 59

cytosine and guanine) were measured by chronoamperometric measurements. Discrimination was achieved due to the different oxidation potentials of each nucleobase and the large electrochemical potential window of graphene. A GO-chitosan nanocomposite was developed to detect S. typhi specific DNA

156

. The GO-chitosan

mixture was first spin-coated on indium tin oxide (ITO) electrodes and the amine groups of the chitosan were reacted with glutaraldehyde to covalently immobilize the amino-modified ssDNA probe. The DNA hybridization event was evaluated monitoring the oxidation of the redox probe methylene blue (MB), used as an intercalator, by differential pulse voltammetry (DPV). When the biosensor was incubated with the complementary target DNA, the hybridization event occurred and thus MB could not intercalate in the ssDNA probe, while exposure to a non-complementary DNA allowed MB binding to the ssDNA and hence MB oxidation peak could be electrochemically measured. Biosensor response for a wide range of complementary target DNA from 10 fM to 50 nM. The MB oxidation peak decreased with increasing concentration of target DNA due to the hybridization event. The developed biosensors exhibited a LOD of 10 fM and 100 fM in buffer and spiked serum samples, respectively. Furthermore, this biosensor was thermally regenerated in aqueous solution and an activity loss of 20% was observed after the 6th use. Stability was also demonstrated to be up to 15 days at 4˚C. Tiwari et al. modified GO with chitosan, as well, as iron oxide NPs to detect E. coli O157:H7 specific DNA by means of EIS

169

. The detection principle was based on DNA hybridization that caused an

increase in Rct and showed a good LOD of 10 fM over a wide linear range of 10 fM to 10 µM. Mohanty and Berry developed a FET-based graphene DNA sensor that was able to detect a single bacterium of Bacillus cereus, a Gram-positive bacterium commonly found as a contaminant in food 170

. Here, ssDNAs were physically adsorbed on the graphene sheets by π-interactions between bases

of the DNA and graphene. Conductance increased upon hybridization of the target bacterial ssDNA. Although this work demonstrated the ability to detect small concentrations of bacteria, it should be noted that electrical measurements were conducted under a dry nitrogen atmosphere, which is not practical in real-world applications. Huang et al. fabricated a graphene-based FET by CVD to detect E. coli bacteria

171

. Graphene was

non-covalently functionalized with 1-pyrenebutanoic acid succinimidyl ester by π-π interaction, followed by the covalent attachment of the antibody via succinimidyl ester groups. Ethanolamine was


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used to quench the remaining succinimidyl esters. The sensing mechanism was based on the field effect produced by the negatively-charged bacterial walls that induced an increase in the hole density and thus in conductance. This immunosensor could detect concentrations of E. coli of 10 cfu/mL in 5

phosphate buffer saline (PBS), while 10 cfu/mL P. aeruginosa did not cause any significant change in signal, demonstrating the high specificity of the biosensor.

3 Electrochemical biosensors based on porous nanostructured materials Porous nanostructured materials including nanoporous anodic alumina (NAA) (pSi) 181-182

174-175

, mesoporous metal oxides

176-178

, mesoporous organosilicas

179-180

172-173

, porous silicon

and porous polymers

have proven their value as high-performing sensor platforms due to high specific surface-to-

area ratio, as well as, versatile chemistry. The large surface area of these materials can enhance the sensitivity of the device due to the increase in the number of immobilized bioreceptors and thus available binding sites. Due to these properties, porous nanomaterials have emerged as attractive platforms to detect a wide range of analytes. Some porous nanostructured materials also offer the possibility to exploit a promising sensing strategy based on nanochannel blockage (NB) among other sensing strategies. This strategy affords the possibility to engineer molecular recognition capabilities by simply tuning the morphology of the nanochannels. It consists of measuring the blockage caused by the analyte when it binds to the immobilized capture probes in the channel walls as shown in Scheme 2. This blockage could be due to steric or electrostatic effects or a combination of both. Voltammetric and impedimetric techniques are commonly employed to electrochemically measure NB, leading to highly sensitive analysis platforms. In the example shown in Scheme 2, the blockage caused by the antigen (blue ball in the scheme) binding in the channels produces a decrease in the voltammetric signal of the oxidation of K4[Fe(CN)6 to K3[Fe(CN)6], which is proportional to the quantity of antigen captured in the channels. Moreover, this strategy can significantly minimize matrix effects when working with real samples by controlling the nanochannel diameter due to the filtering capabilities of these materials

183

. In other

words, species larger than the nanochannel diameter are not able to diffuse through the channel for



Page 24 of 59

steric reasons, and thus the effect of species that could interfere in the binding event is reduced. The simplicity of this label-free strategy combined with highly sensitive electrochemical techniques offers the possibility to develop high-performing platforms that could be adapted to point-of-care portable devices.

Scheme 2. Scheme of the sensing principle based on NB for label-free detection. (Above) Porous nanochannel-modified electrode and functionalized with capture probes (red diamonds). (Below) DPV traces prior and after analyte (blue balls) binding.

Porous nanostructured materials, exploiting NB, have been successfully incorporated on conventional electrode surfaces such as Au, Pt and C. Of these, NAA has been most commonly employed, followed by MPS and then pSi. Recently, nanochannels formed by nanosphere assembly have also taken advantage of this sensing approach. The following section will give key examples of the electrode preparation and applications using these materials.


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3.1.1

Nanoporous anodic alumina (NAA)

NAA membranes (NAAMs) integrated onto different transducers have been extensively used to detect and quantify analytes based on NB strategy

173

. These membranes are typically prepared by

electrochemical anodization of high purity Al in acid electrolytes consisting of aqueous solutions of sulphuric, oxalic or phosphoric acid, followed by the dissolution of the oxide layer at the bottom of the pores

184

. They can also be fabricated directly on an ITO substrate by depositing an Al layer via CVD

as recently reported

185

. This alternative is particularly exciting due to the conductive properties of

ITO, which can therefore be employed as an electrotransducer. Moreover, NAA is rich in hydroxyl groups, which can be easily functionalized. Silanes or organic acids are generally used to modify the channels, followed by the desired bioreceptor immobilization

186-187

. NAA biosensors using the NB

approach and combined with voltammetric or impedimetric techniques have been reported for bacteriophage 188, virus 189-191, DNA 192-196, and protein 183, 197-200 detection. Smirnov’s group pioneered the use of NAAMs exploiting NB for electrochemical biosensors

201

. Here,

NAAMs with two channel diameters (20 and 200 nm) were covalently modified with ssDNA to detect target DNA by monitoring impedance changes upon DNA hybridization. This resulted in the blockage of the pore, increasing the pore resistance for the 20 nm channel diameter NAA biosensor, while no significant changes were observed for the 200 nm channel diameter one. This was attributed to the short length of dsDNA compared to channel diameter. The same group has recently reported NAA immunosensors to detect MS2 bacteriophage using EIS

188

. Sensor response to different MS2

concentrations (10 – 1870 pfu/mL) was monitored using two nanochannel diameters (73 and 97 nm). It was found the larger channel diameter exhibited higher sensitivity, given by the slope of the fitting curve, and LOD of 7 pfu/mL. The specificity of the immunosensors was also tested by incubating MS2 and Qβ bacteriophage mixtures in 1:1 ratio and proved to be unaffected. Toh’s group has also employed NAAMs for the fabrication of electrochemical biosensors based on NB

189, 198

. A NAAM was

grown on a Pt disk electrode, by sputtering a thick Al film (300 – 500 nm), followed by anodization

202

.

Nanochannels were functionalized with antibodies via physical adsorption to detect West Nile virus (WNV) (Figure 6), and in particular, WNV protein domain III (WNV-DIII) and West Nile virus particles using alternating current voltammetry (ACV) in the presence of ferrocenemethanol as a redox probe 189

. The effect of the antibody concentration used for membrane modification, and the pH and ionic



Page 26 of 59

strength of the incubation solution, were investigated to determine the optimal sensing conditions, and found to be 0.2 µg/mL of antibody, pH 7.6 and 0.1 M NaCl. This platform showed a low LOD for both WNV-DIII (4 pg/mL) and WNV particles (2 particles/100 mL), similar to PCR-based techniques. Moreover, the performance of the immunosensor in WNV spiked blood serum was successfully demonstrated. This platform and sensing approach was also applied to the ultrasensitive detection of Dengue virus particles pneumophila DNA

190

and real time cDNA PCR sample of Dengue virus

194

, as well as, Legionella

193

, being able to detect as low as 1 cfu/mL, 9.55 x 10-12 M and 3.1 x 10-13 M, -12

respectively. The same group improved the LOD of Dengue virus DNA down to 1 x 10 both sides of the NAA membrane with Pt (50 – 100 nm thickness)

203

M by coating

. By coating both sides of the

membrane, the resistance of the solution on the nanochannel entry or “mouth” was eliminated.

Figure 6. (A) SEM image of the NAAM surface (above) and thickness (below) and (B) immunosensor scheme for the detection of WNV-DIII and WNV particles. Reproduced with permission from ref 189 (Copyright 2009 ACS Publications) and ref 190 (Copyright 2012 Elsevier).

Merkoçi’s laboratory has also extensively explored and reviewed this sensing approach for electrochemical sensors based on NAAMs

173, 204-205

, with special focus on protein and DNA detection


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for clinical diagnostics

183, 192, 197, 199-200

.

Here, commercial NAAMs with 20 or 200 nm channel

diameter were physically fixed onto commercial SPCEs and placed in a methacrylate cell as shown in Figure 7. These electrodes were functionalized with different capture probes including antibodies, ssDNA and aptamers. NB was measured by means of DPV. They firstly reported a label-free immunosensor to detect human immunoglobulin G (IgG) in blood samples

197

. Channel walls were

first silanized with APTES followed by the covalent attachment of anti-human IgG antibodies. The main parameters influencing the immunoassay, such as antibody concentration, incubation and reaction time, were optimized as shown in Figure 7C. The optimized conditions were used to investigate the effect of the channel diameter (20 vs. 200 nm) in the detection of human IgG. It was found that the smallest channel diameter could detect 2.5 times lower concentrations than the largesized one. The sensitivity of the large-sized channel diameter (200 nm) was further improved using two approaches. The first consisted of using labels in a sandwich immunoassay-based strategy 183

(Figure 7D)

. AuNPs tags of two different sizes (i.e. 20 and 80 nm) were employed to amplify the

signal, as well as, enhance the NB effect by using them as blocking agents inside the channel. The biosensor using 80 nm-sized AuNP-modified detection antibody allowed the detection of 2 µg/mL human IgG, a ten-fold enhancement compared to the detection using 20 nm AuNP-modified detection antibody, and 250 times lower than the label-free assay (500 µg/mL). Moreover, protein LOD decreased to 50 ng/mL by increasing the AuNPs size with Ag metal deposition after binding inside of the NAA surface, which further enhanced the blockage of the channels upon the biorecognition event and the redox probe diffusion (Figure 7E). The second strategy was based on using a larger redox indicator

200

. Prussian Blue NPs coated by polyvinylpyrrolidone, that were 4 nm in size, were chosen

as an alternative to the small [Fe(CN)6]

4-

ions employed in the label-free assay. The larger redox

indicator was expected to increase the steric effects along the channel and thus hinder their diffusion towards the transducer. This approach remarkably improved the performance of the biosensor, showing an excellent LOD of 34 pg/mL and allowing the detection of proteins at levels encountered in clinical samples.



Page 28 of 59

Figure 7. (A) Set up of the NAAM electrode. (B) SEM images of the top surface (left) and crosssection (right) of the membrane. (C) Optimization of the antibody concentration (a), antibody adsorption time (b), and immunological reaction time (c), to obtain the best sensing performance using 200 nm NAA (the optimal values circled in red). (D) Illustration of the different immunoassay strategies to detect human IgG by label-free assay (1) and sandwich assay with AuNP-modified antibody (2) and Ag coated AuNPs-modified antibody (3). (E) DPV voltammograms obtained after incubating 5 µg/mL of IgG directly (a), or with 20 nm-sized AuNPs-modified antibody (b), 80 nm-sized AuNPs-modified antibody (c) and 80 nm-sized AuNPs-modified antibody after Ag deposition (d).


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Reproduced with permission from ref 173 (Copyright 2016 Elsevier) and ref 197 (Copyright 2010 Elsevier). Kant et al.

206

recently investigated the influence of the NAAM geometry (i.e. nanochannel diameter

and length) using a model analyte-bioreceptor interaction: biotin-streptavidin. The sensing performance was evaluated by measuring changes in the channel resistance and conductance by non-faradaic EIS for different sized nanochannels. NAAMs were first functionalized with 2carboxyethyl phosphonic acid to introduce carboxylic groups. Then, the top and bottom surface functionalization was removed with air plasma for 30 s in order to just modify the inner channel walls. The inner channels were modified with streptavidin via EDC/NHS chemistry. The performance of the biosensor with three channel diameters (25, 45 and 65 nm) was first studied by incubating in different biotin concentrations. The results showed that biotin-streptavidin binding caused significant changes of surface charge and conductivity and thus in the channel resistivity, which was found to be a key parameter to study the performance of the biosensor. Changes in the channel resistivity were larger in magnitude for the smallest nanochannel (25 nm) due to charged streptavidin-biotin complexes being closer to each other and more sensitive to oscillation of ions inside the channel compared to the larger nanochannels (45 and 65 nm). Since the biosensors with the smallest nanochannel diameter exhibited greater sensitivity, NAA platforms with 25 nm channel diameter with different nanochannel lengths (4.5, 9, 13.5 and 18 µm) were fabricated to investigate the influence of the channel length in the sensing performance. Long nanochannels (> 10 µm) were observed to hinder the diffusion of the analyte due to the higher nanochannel resistance, thus small changes were observed inside the channel. It was concluded that NAAMs with nanochannel diameter of 25 nm and length up to 10 µm exhibited the best performance. This work shows that nanochannel dimensions play a crucial role in achieving the desired sensing capabilities.

3.1.2

Mesoporous silica (MPS)

MPS features highly ordered structures with tunable channel diameter (2 – 50 nm) and shape, and 2

high surface area (up to 1500 m /g)

27

. This material is easy to functionalize, due to its hydroxyl-

terminated surface. Fabrication strategies are based on nanoparticles and surfactant self-assembly as



Page 30 of 59

templates, followed by the subsequent removal of these templates, resulting in the final porous structure

207

. Highly ordered MPS films have started receiving attention for biosensing

207-209

.

Jin et al. constructed a MPS-based electrochemical immunosensor harnessing the NB sensing strategy

210

. MPS was synthesized following the protocol described in

211-213

and selectively modified

so that the biorecognition event could only take place in the channels. This was achieved by modifying the internal channel walls with antibodies and the external ones with trimethylsilyl chloride (TMC). Briefly, the external surface was modified with TMC using the hydroxyl groups on the surface prior to template removal and hence mesoporous channel formation. After template removal, the hydroxyl groups in the internal walls were modified with APTES, followed by antibody immobilization. The antibody-modified MPS was then attached to the ITO surface for the simultaneous detection of proteins. Biosensors performance was improved by incorporating AuNPs

214

and SWCNTs

215

in the

MPS channel to promote electron transfer. Other MPS platforms integrated on screen-printed graphite electrodes (SPGEs)

216

and GCE

213

were also successfully developed. However, these techniques

relied on sandwich assay strategies, increasing the analysis time and cost due to the use of additional immunoreagents. Although MPS NB-based biosensors were successfully fabricated, most of the approaches relied on the use of signal amplification (AuNPs, CNTs) or labelled strategies, making the analysis expensive and time-consuming. Further research, based on label-free detection and without the use of decorations for electron transfer enhancement would be required determining the feasibility of MPS for a simple, cost-effective biosensor based on NB.

3.1.3

Porous silicon (pSi)

pSi provides an attractive means to detect a broad range of target analytes due to its large specific 2

surface area (up to 800 m /g) and ease of functionalization

217

. The average pore size and thickness

can be easily adjusted to allow the penetration of different sized species. pSi is commonly fabricated by electrochemical anodization of single crystalline Si in an aqueous HF solution, resulting in a hydride-terminated surface. This hydride-terminated surface is modified by oxidation, silanization, hydrosilylation and thermal carbonization techniques. Most of the pSi electrochemical biosensors


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developed take advantage of the large surface area provided 174-175, 218-222, while pSi electrodes relying on NB to transduce analyte binding have only been reported by our group

223

. This work showed the

potential of pSi membranes to exploit NB effect as a promising sensing strategy. Figure 8A shows the pSi membrane-modified Au electrode fabrication that consisted of the electrochemical anodization of Si, followed by the detachment from the Si substrate by applying a series of high current pulses and transferred to the gold-coated slide. pSi membranes, consisting of Si nanochannels arrays, were functionalized with anti-MS2 antibodies for the direct detection of MS2 bacteriophage in water samples. This was achieved by thermal hydrosilylation with undecylenic acid to obtain a carboxylterminated surface, followed by the EDC/NHS activation to immobilize the anti-MS2 antibodies (Figure 8B). The performance of the immunosensor was investigated using DPV for three different sized channel diameters (85, 57 and 40 nm) for a wide range of MS2 concentrations (1 – 1010 pfu/mL) in PBS buffer as shown in Figure 8C. The oxidation current decreased for the increasing concentration of MS2 for all the immunosensors and the immunosensor with the largest nanochannel diameter (85 nm) showed the highest sensitivity with an excellent LOD of 6 pfu/mL. Moreover, the feasibility to detect MS2 in spiked reservoir water samples was also successfully demonstrated and was unaffected by interfering species that could be present in water samples. Detection limits at levels encountered in wastewater and sewage impacted wetlands were achieved (Figure 8E). The excellent performance of this platform highlights its potential to be used in water quality measurements for in-the-field applications.



Page 32 of 59

Figure 8. (A) Schematic of the pSi membrane-modified Au electrode fabrication. (B) Functionalization of the pSi membrane-based electrode. (C) (Left to right) DPV voltammograms for the anti-MS2 immunosensors with average nanochannel diameter of 85, 57 and 40 nm for the increasing concentration of MS2 bacteriophage in buffer. (D) Dose response curves for all immunosensors modified with either MS2-specific antibodies or non-specific antibodies (control) in buffer. (E) Dose response curves of MS2 in river water (orange) and buffer (blue) for the 85 nm nanochannel diameter immunosensor. Reproduced with permission from ref 223. Copyright 2016 Elsevier.


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3.1.4

Polystyrene (PS) nanochannels

A very interesting strategy based on the self-assembly of polystyrene (PS) nanospheres to fabricate nanochannel arrays was recently proposed for electrochemical analysis (Figure 9)

224

. PS

nanochannels were fabricated on ITO electrodes, screen printed onto a polyethylene terephthalate substrate, by depositing a monolayer of PS nanospheres using a dip-coating technique. The spaces between the assembled nanospheres generate well-ordered inter-particle space or nanochannels. Two sizes of nanosphere were employed, 200 nm and 500 nm, with resulting inter-particle spacings (nanochannel diameter) of 24 nm and 65 nm, respectively. The formed nanochannel arrays were used for immunoglobulin G (IgG) detection (chosen as a model analyte) by functionalizing the carboxyl-terminated nanospheres with an anti-IgG antibody using EDC/NHS chemistry. The immobilized antibodies form an immunocomplex inside the nanochannels after capturing IgG, partially blocking the diffusion of the redox probe along the channel, resulting in a decrease in the DPV signal. The biosensor developed using 200 nm nanospheres exhibited the lowest LOD (580 ng/mL) with excellent selectivity for IgG against other proteins that could be present in real samples. Furthermore, performance in human urine samples was also better in terms of LOD than others previously reported by the same group using a NB approach based on NAA substrates

183, 197

. The improvement in

sensitivity can be attributed to the decrease of the length of the nanochannels from 60 µm (NAA) to 200 – 500 nm (PS). Moreover, the ease of fabrication and robustness of this platform may bring about new opportunities to develop sensing platforms based on this technology.



Page 34 of 59

Figure 9. Electrochemical biosensors based on PS nanosphere assembly to form well-ordered PS nanosphere channel arrays (based on inter-particle distance). (A) SEM images of PS monolayer on ITO electrode with 200 nm (up) and 500 nm (down) size nanospheres. (B) Schematic representation of the sensing strategy based on the channel blockage. (C) DPV spectra for the increasing concentrations of IgG from top to bottom: 1, 50, 100, 200 and 300 µg/mL in 1 mM K3[Fe(CN)6]/0.1 M NaNO3. (D) Dose response curves of IgG for the monolayers prepared using 200 nm sized (solid line) and 500 nm sized nanospheres (dashed line). Reproduced with permission from ref 224. Copyright 2015 Royal Society of Chemistry.

The most significant nanostructured biosensors employed for the detection of pathogens and toxins described in this review are summarized in Table 2. This table includes the different target pathogen/toxin detected, the type of nanostructured platform employed and some of the most important characteristic of a biosensor such as detection technique, LOD, detection range and the final application.


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Table 2. Most significant electrochemical nanostructured biosensors for pathogen and toxins detection.

Pathogen/ toxin

Platform

Detection

LOD

Detection range

Application

Ref.

EIS

100 cfu/mL

102 – 105 cfu/mL

----

79

EIS

150 cfu/mL

1.5 x 102 – 1.5 x 107

Beef

cfu/mL

cucumber

technique E. coli

MWCNTs-modified electrode AuNPs-graphene electrode

and

130

samples Graphene-FET

I

vs

V

10 cfu/mL

102 – 105 cfu/mL

----

171

10 fM

10 fM – 10 µM

----

169

100 cfu/mL

100 – 500 cfu/mL

----

86

1 cfu/mL

0.2 – 106 cfu/mL

Fruit juice and

78

measurements

E. coli specific

GO- iron oxide NP

DNA

and

EIS

chitosan

composite S. infantis

CNT-FET

I

vs

V

measurements SWCNTs-modified

EMF

electrode

(Potentiometric)

S. typhi specific

GO-chitosan

DPV

DNA

composite

S. typhi

vs

time

milk samples 10 fM

10 fM – 50 nM

Serum samples

156

35



Pathogen/ toxin

Platform

Detection

LOD

Detection range

EIS

100 cfu/mL

CV

Page 36 of 59

Application

Ref.

10 – 10 cfu/mL

6

----

79

3 ng/mL

10 – 1480 ng/mL

----

80

Conductance vs

1

virus

----

----

99

time

particle

Conductance vs.

1

virus

----

----

99

time

particle

Conductance vs.

29

virus

2.85 – 2.85 x 103

Clinical exhaled

100

time

particle /µL

virus particle /µL

breath samples

10 fM of the

1

technique K. pneumoniae

MWCNTs-modified

2

electrode

Sterigmatocystin

SWCNT-chitosan composite-modified electrode

Adenovirus

Influenza A

SiNW-FET

SiNW-FET

SiNW-FET

DEN-2 virus DNA

SiNW-FET combined

I

sequences

with RT-PCR

measurements

amplicons

amplicons

L.

TiO2

EIS

500 cfu/mL

monocytogenes

modified electrode

Pathogen/ toxin

Platform

Detection

LOD

nanowire-

vs

V

–

100

fm

of

----

101

10 – 107 cfu/mL

----

102

Detection range

Application

Ref.

36


Page 37 of 59 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


technique Hepatitis B virus

AuNP-SWCNT

DNA

composite

EIS

1

aM

of

1 – 106 aM

----

130

1 – 106 aM

----

130

20 – 120 aM

----

140

complement ary 21-base DNA

Papilloma

virus

DNA

AuNP-SWCNT

EIS

composite

1

aM

of

complement ary 24-base DNA

MNV

AuNP-modified

SWV

SPCEs

MS2

pSiM-modified

bacteriophage

electrode NAAM-modified

180

virus

particles 10

DPV

6 pfu/mL

1 – 10 pfu/mL

Reservoir water

223

EIS

7 pfu/mL

10 – 1870 pfu/mL

----

188

Detection

LOD

Detection range

Application

Ref.

electrode

Pathogen/ toxin

Platform

technique

37



WNV

NAAM-modified electrode

ACV

4 pg/mL of

5 – 55 pg/mL of

WNV-DIII

WNV-DII and 0.03 –

and 2

0.6 cfu/mL

Page 38 of 59

Blood serum

189

particles/100 mL of WNV particles

38


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4 Summary and outlook Research on point-of-care devices is underpinned by electrochemical sensing platforms that allow direct, label-free detection, due to their feasibility to perform near real-time measurements and ease of handling. The simplicity in the measurement is sometimes achieved by a more sophisticated design of the sensing platform, often involving the use of nanomaterials that own excellent electrical conductivity and mechanical strength, large surface area, easy functionalization, and unique morphology. The use of nanostructured materials is rapidly expanding in the field of electrochemical biosensors. The unique properties of the emerging nanomaterials have paved the way for remarkable improvements in the sensing capabilities. These nanomaterials fall squarely into two main categories: nanoscale and porous nanomaterials. The use of nanoscale materials to prepare electrodes includes nanotubes, nanowires, nanoparticles and nanosheets; with carbon nanotubes, silicon nanowires, metal nanoparticles and graphene being the most employed. Amongst them, 2D materials are gaining interest for the development of FET devices. Porous nanostructured materials such as nanoporous anodic alumina, mesoporous silicon and porous silicon have also had a major impact on the field. Amongst them, porous nanostructured materials exploiting NB, as sensing strategy, are promising platforms for a direct and simple detection system that could be adapted to point-of-care devices. These materials are expected to play a key role in the development of devices for single molecule detection. New advances in nanofabrication will make nanoscale materials accessible to the broad research community and enable the development of highly performing analysis devices. Despite this progress, there are still several challenges that need to be overcome to make these devices commercially available for environmental monitoring and food quality control. This includes addressing the ability to handle real-world sample matrices and integration of the transducers into easy-to-use devices. Therefore, as we transition into the next generation of biosensors, the aim will be to better engineer the surface of the electrode to avoid non-specific adsorption of interfering species and combine the sensors with microfluidics solutions and electronic components for signal read-out to achieve integrated devices that can be reduced to practice.

Acknowledgment Author acknowledges the top-up scholarship funding from the National Centre of Excellence in Desalination Australia (NCEDA). This work was performed in part at the Melbourne Centre for



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Nanofabrication (MCN) in the Victorian Node of the Australian National Fabrication Facility (ANFF). The authors thank Marc Cirera for his help with the schematics (http://marccirera.com/).


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