nanoscale biosensors and biochips

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CHAPTER 1

NANOSCALE BIOSENSORS AND BIOCHIPS

Wayne R. Leifert1,*, Richard V. Glatz2, Kelly Bailey3,4, Tamara Cooper3,4, Marta Bally5, Brigitte Maria Stadler6, Erik Reimhult7 and Joseph G. Shapter8

1Commonwealth Scientific and Industrial Research Organization (CSIRO),

Division of Human Nutrition, Adelaide, Australia; 2South Australian Research

and Development Institute (SARDI), Department of Entomology, Adelaide,

Australia; 3Commonwealth Scientific and Industrial Research Organization

(CSIRO), Division of Molecular and Health Technologies, Adelaide, Australia; 4School of Molecular and Biomedical Science, The University of Adelaide,

Adelaide, Australia; 5Laboratory of Biosensors and Bioelectronics, Institute

for Biomedical Engineering, ETH Zurich, Zurich, Switzerland; 6Centre for

Nanoscience and Nanotechnology, Department of Chemical and Biomolecular

Engineering, The University of Melbourne, Melbourne, Australia; 7Laboratory

for Surface Science and Technology, Department of Materials, ETH Zurich,

Zurich, Switzerland; 8School of Chemistry, Physics and Earth Sciences,

Flinders University, Adelaide, Australia

*Corresponding author, Email: [email protected]

1. General Introduction

Recent advances in molecular biology, surface chemistry, protein purification, signal transduction/amplification, lipid chemistry and nanofabrication technologies have converged in a relatively new field of science, that of molecular biosensing. The growing level of interest in biosensing research has seen the formation of a range of specific journals reporting on advances in the field. Biosensors and biochips have many potential (and some current) applications including provision of point-of-care diagnostic tools, high-throughput drug discovery tools and in-field sensing tools for a variety of compounds including toxins and/or contaminants. Biosensors are generally accepted as being analytical

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devices based on a biologically active or biologically mimetic compound (a detector) coupled to a physical signal transduction mechanism (a transducer/reporter). The interaction of analyte with its biomolecular detector is thus exploited to produce an effect that can be measured through the transducer. Biochips generally refer to an array of individual biosensors, and can also be referred to as nucleotide or protein microarrays. This review explores examples of the various approaches to biosensor and biochip fabrication, including examination of components important to each system. We discuss advantages and disadvantages of biosensors that exploit whole cells as detectors and those which make use of one or several specific molecules, the most widely utilized being membrane-associated proteins, such as G-protein coupled receptors (GPCRs) and ion-channels, due to their diversity and current importance for sensing and screening technologies.

There are two major challenges to establishing functionally active biological components within a sensor or chip design, these being the controlled capture and positioning of the biological detector onto a surface and maintenance of the detector’s functional/structural integrity. In the case of membrane proteins, it is vital that an appropriate hydrophobic environment is available to maintain protein structure and function. The use of lipid supports for this purpose, is being widely studied, and is discussed in this review. In combination with appropriate surface compositions, various techniques of nanopatterning technologies are important in the fabrication of biosensors in order to control the location, distribution, amount and orientation of the biomolecules on the surface. There are also a range of physical and biological substrates on which sensor and chip platforms are built and we pay particular attention to nanotubes as physical substrates due to their promising application in electrochemical biosensing.

In order to detect changes occurring at the sensor or chip surface, a variety of transduction techniques exist and we discuss a range of nanosized reporter labels for their potential for biosensing applications. Finally, we touch on some of the applications and specific analytes commonly monitored using biosensing today.

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2. Biological Detectors Used in Biosensing and Biochips

Since the first description of the biosensor concept by Clark Jr. and Lyons in 1962 involving glucose oxidase and the oxygen electrode [1], a variety of front-end biological detectors have been investigated for use in biosensing. These bio-recognition elements include enzymes, membrane proteins such as cell-receptors and ion-channels, antibodies, nucleic acids, virus particles and intact cells.

We begin the discussion with an investigation of membrane proteins, which as diverse biomolecules crucial in cellular signaling and communication as well as regulation of transport into and out of the cell are of significant interest in the fabrication of biosensors and biochips. Because of their diversity, biological importance and level of characterization, the two key classes of membrane proteins utilized are the GPCR family and ion-channels.

2.1. G-Protein Coupled Receptor Biosensors (GPCRs)

2.1.1. Importance of GPCRs

Many disease processes involve aberrant or altered GPCR signaling dynamics and GPCRs represent the most significant target class for medicinal pharmaceuticals (≈50% of marketed drugs, see Table 1) [2]. GPCRs are associated with almost every major therapeutic category or disease class, including pain, asthma, inflammation, obesity, cancer, as well as cardiovascular, metabolic, gastrointestinal and central nervous system diseases [3]. It is this vitally important function of these cell-surface receptors combined with the huge diversity of specific ligands they bind, which makes GPCRs so physiologically significant and attractive for biosensing applications. Therefore, there is a need to develop sophisticated and appropriate GPCR biosensors for the detection of a variety of ligands (Figure 1). This section focuses on some of the available cell-free GPCR assay nanotechnologies [4] and describes some of the more sophisticated functional GPCR biosensors. Cell-free biosensors have the potential advantage of being applicable to a range of assay environments in which cells may be damaged.

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Table 1. Examples of some pharmaceuticals which target GPCRs for the indicated condition or disease state.

Brand Name (generic) G-protein coupled receptor(s)

Disease/Indication

Zyprexa (Olanzapine) Serotonin 5-HT2 and Dopamine

Schizophrenia, Antipsychotic

Risperdal (Risperidone) Serotonin 5-HT2 Schizophrenia

Claritin (Loratidine) Histamine H1 Rhinitis, Allergies

Imigran (Sumatriptan) Serotonin 5-HT1B/1D Migraine

Cardura (Doxazosin) α-adrenoceptor Prostate hypertrophy

Tenormin (Atenolol) β1-adrenoceptor Coronary heart disease

Serevent (Salmeterol) β2-adrenoceptor Asthma

Duragesic (Fentanyl) Opioid Pain

Imodium (Loperamide) Opioid Diarrhea

Cozaar (Losartan) Angiotensin II Hypertension

Zantac (Ranitidine) Histamine H2 Peptic ulcer

Cytotec (Misoprostol) Prostaglandin PGE1 Ulcer

Zoladex (Goserelin) Gonadotrophin-releasing factor

Prostate cancer

Requip (Ropinirole) Dopamine Parkinson’s disease

Atrovent (Ipratropium) Muscarinic Chronic obstructive pulmonary disease (COPD)

GPCR activation can be initiated by a wide variety of stimuli such as light, odorants, neurotransmitters and hormones (Figure 1). In cells, the extracellular ligand is specifically and sensitively detected by a cell surface GPCR. Once binding/recognition takes place, the GPCR triggers the activation of a cellular heterotrimeric G-protein (guanine nuceleotide-binding protein) complex consisting of Gα, Gβ and Gγ subunits (Figure 1). Finally, the “signal transduction” cascade (in whole cells at


least) involves the activated G-proteins altering the activity of downstream “effector” protein(s) to yield a response that can be used to detect the binding event. This is the basis of many existing screening assays and some biosensor designs have exploited this approach, combined with different surface-attachment and transduction technologies.

In cell-free GPCR assays discussed in this section, host cells which have been transfected with DNA encoding a particular GPCR of interest allow the cellular expression of the GPCR. Subsequently, the cells are treated in such a way as to allow a partial purification of the GPCRs (in their cell membranes) to obtain an ongoing supply of GPCRs. The purification process can result in small (nanometer scale) crude membrane fragments containing the GPCRs, which are suitable as detector molecules for biosensor applications [5].

2.1.2. Surface Capture of GPCRs

The arraying of membrane GPCRs has required appropriate surface chemistry for the immobilization of the lipid phase containing the GPCR of interest [6-8]. Surface modification with γ-aminopropylsilane (an amine presenting surface) provided the best combination of properties to allow surface capture of the GPCR-G-protein complex from crude membrane preparations, resulting in microspots of approximately 100 µm diameter. Atomic force microscopy (AFM) demonstrated that the height of the supported lipid bilayer was approximately 5 nm, corresponding to GPCRs confined in a single, supported lipid layer scaffold [9]. Using these chemically-derivatized surfaces, it was possible to demonstrate capture of fluorescently labeled β1, β2, and α2A subtypes of the adrenergic receptor, as well as neurotensin-1 receptors and D1-dopamine receptors. Dose-response curves using the fluorescently-labeled ligands gave IC50 values in the nM range suggesting that the GPCR-G-protein complex was largely preserved and biologically intact in the microspot. Furthermore, good long-term stability was achieved.

Waller et al. [10] conjugated dextran beads with dihydroalprenolol, an antagonist of β2-adrenergic receptors (β-AR). This allowed the capture of solubilized β-AR to this immobilized surface ligand. The β- AR was expressed as a fusion protein with green fluorescent protein

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Figure 1. A list of some of the known endogenous and exogenous GPCR ligands and a schematic depicting the transmembrane topology of a typical “serpentine” G-protein coupled receptor (GPCR) with its associated heterotrimeric G-protein complex. The membrane patch containing the GPCR with associated G-proteins is schematically shown attached to a theoretical solid support matrix. The receptor polypeptide chain traverses the plane of the membrane phospholipid bilayer seven times. The hydrophobic transmembrane segments of the GPCR are indicated by spirals. The ligand can bind to the receptor from the “extracellular” (outer) surface or depending on the receptor type, to a site deep within the receptor, surrounded by the transmembrane regions of the receptor protein. In this way, the receptor can act as a detector of its ligands. The G-proteins (Gα and Gβγ) are shown to interact with the “cytoplasmic” side of the receptor.

γγγγ

ligand

GPCR

γγγγββββ

αααα

Ni2+

SUPPORT MATRIX

lipid bilayermembrane

Acetylcholine

Adenosine

Adrenaline

Adrenocorticotropic hormone

Angiotensin II

Bradykinin

Calcitonin

Chemokines

Cholecystokinin

Corticotropin releasing factor

Dopamine

Endorphins

Endothelin

Enkephalins

Fatty acids

Follitropin

GABA

Galanin

Gastric inhibitory peptide

Gastrin

Ghrelin

Glucagon

Glutamate

Gonadotropin-releasing hormone

Growth hormone-releasing factor

Growth-hormone secretagogue

Histamine

Luteinising hormone

Lymphotactin

Lysophospholipids

Melanocortin

Melanocyte-stimulating hormone

Melatonin

Neuromedin-K

Neuromedin-U

Neuropeptide-FF

Neuropeptide-Y

Neurotensin

Noradrenaline

Odorants

Opioids

Orexin

Oxytocin

Parathyroid hormone

Photons (light)

Platelet activating factor

Prolactin releasing peptide

Prostaglandins

Secretin

Serotonin

Somatostatin

Substances P, K

Thrombin

Thromboxanes

Thyrotropin

Thyrotropin releasing hormone

Tyramine

Urotensin

Vasoactive intestinal peptide

Vasopressin

GPCR ligands

γγγγ

ligand

GPCR

γγγγββββ

αααα

Ni2+

SUPPORT MATRIX


γγγγ

ligand

GPCR

γγγγββββ

αααα

Ni2+

SUPPORT MATRIX


Acetylcholine

Adenosine

Adrenaline

Adrenocorticotropic hormone

Angiotensin II

Bradykinin

Calcitonin

Chemokines

Cholecystokinin

Corticotropin releasing factor

Dopamine

Endorphins

Endothelin

Enkephalins

Fatty acids

Follitropin

GABA

Galanin

Gastric inhibitory peptide

Gastrin

Ghrelin

Glucagon

Glutamate

Gonadotropin-releasing hormone

Growth hormone-releasing factor

Growth-hormone secretagogue

Histamine

Luteinising hormone

Lymphotactin

Lysophospholipids

Melanocortin

Melanocyte-stimulating hormone

Melatonin

Neuromedin-K

Neuromedin-U

Neuropeptide-FF

Neuropeptide-Y

Neurotensin

Noradrenaline

Odorants

Opioids

Orexin

Oxytocin

Parathyroid hormone

Photons (light)

Platelet activating factor

Prolactin releasing peptide

Prostaglandins

Secretin

Serotonin

Somatostatin

Substances P, K

Thrombin

Thromboxanes

Thyrotropin

Thyrotropin releasing hormone

Tyramine

Urotensin

Vasoactive intestinal peptide

Vasopressin

GPCR ligands


(GFP), thus allowing fluorescent measurement of bound receptor possible. Thus, it was possible to screen for ligands of the β-AR through the reduction in fluorescence as receptor was removed from the beads due to competition of free (added) ligand. Another successful bead-based approach used paramagnetic beads to capture CCR5 receptors from a cell lysate held within a lipid bilayer [11]. More recently, site directed immobilization of membrane extracts containing either the M2-muscarinic receptor or H1-histamine receptor, using complementary oligonucleotides has been investigated (unpublished data). Sequence-directed immobilization of oligo-tagged vesicles carrying GPCRs could potentially lead to the development of a self-sorting array platform for a large number of different receptor sub-types, through the inherent selectivity of complementary strands of oligonucleotides.

2.1.3. Ligand-Binding at GPCRs

Ligand binding to a GPCR attached to a surface has been reported for the chemokine CCR5 receptor using surface plasmon resonance (SPR) [12]. For such GPCR surface display, purification of the GPCR has not always been necessary and crude membrane preparations have either been fused with an alkylthiol monolayer (approximately 3 nm thickness) formed on a gold-coated glass surface, or onto a carboxymethyl modified dextran sensor surface [13]. One problem of surface based assays is the difficulty in obtaining the correct orientation of the receptor once attached to the surface. This problem was overcome by using conformationally-dependent antibodies [14]. In this biosensor application, SPR has a distinct advantage as a screening tool since this technique can detect the cognate ligand without requiring fluorescent or radio-labeling. This allows SPR to be used in complex fluids of natural origin thus simplifying the development of assay technologies.

Martinez et al. [15] used total internal reflection fluorescence (TIRF) to demonstrate ligand binding to the neurokinin-1 GPCR by surface immobilization of membrane fragments containing this receptor. The GPCR was expressed as a biotinylated protein using mammalian cells and could be selectively immobilized on a quartz sensor surface coated

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with streptavidin (streptavidin binds biotin with extremely high affinity). Using this approach, it was not necessary to detergent-solubilize and reconstitute the neurokinin-1 receptors, thus avoiding the deleterious effect(s) associated with such processes. The preparation of the biotinylated receptors allowed for a high-affinity interaction between biotin and streptavidin and thus a template-directed and uniform orientation of the neurokinin-1 receptor on the support matrix. TIRF measurements were made using a fluorescent-labeled agonist (i.e., the cognate agonist substance-P labeled with fluorescein). The highly sensitive TIRF fluorescence detection methodology was able to resolve the binding of fluorescently-tagged ligand (agonist) to as little as one attomol of receptor molecules [15]. This sensitivity far exceeds that of current physical approaches to biosensing (e.g. gas chromatography, mass spectrometry) and is a major reason why biosensing has become an important research area.

2.1.4. Detecting GPCR Conformational Changes

The detection of intrinsic conformational changes in the GPCR following ligand (agonist) activation generally involves the use of fluorescence-based techniques and has been limited to date. One study demonstrated the immobilization of β2-adrenergic receptors onto glass and gold surfaces [16]. The receptors were site-specifically labeled with the fluorophore tetramethyl-rhodamine-maleimide at Cysteine 265 (Cys265) and the agonist-induced signal was large enough to detect using a simple intensified charge-coupled device (ICCD) camera image. Therefore, it was suggested that the technique may be useful for drug screening with GPCR arrays.

In a recent study, ligand binding to the β2-adrenergic receptor has been demonstrated using plasmon waveguide resonance (PWR) [17]. Using this technique, changes in the refractive index upon ligand binding to surface-immobilized receptor results in a shift in the PWR spectra. Previously, PWR technology was used for detection of conformational changes in a proteolipid membrane containing the human δ-opioid receptor following binding of several types of ligands [18]. Although the


ligands in that study [18] were of similar molecular weight, there were distinctly different refractive index changes induced by ligand binding and these were too large to be accounted for by differences in the mass alone. The inference from this finding was that a ligand-specific conformation change in the receptor protein may have been detected, a phenomenon that further increases the specificity of ligand-mediated PWR spectral alterations.

2.1.5. GTP Binding at G-Protein Subunits

GPCR biosensors can also utilize the use of non-hydrolyzable GTP-analogs such as radiolabeled [35S]GTPγS or fluorescent-tagged Europium-GTP, which bind to the receptor-activated form of the Gα subunit targeting the site of guanine nucleotide exchange (GDP for GTP on the Gα subunit of the Gαβγ heterotrimer).

Guanine nucleotide exchange is a very early, generic event in the signal transduction process of GPCR activation which can be measured without the need for intact cells and is, therefore, an attractive event to monitor. The radiolabeled [35S]GTPγS or fluorescent Europium-GTP binding assays measure the accumulative level of G-protein activation following agonist activation of a GPCR by determining the binding of these non-hydrolyzable analogs of GTP to the Gα subunit. Therefore, they are defined as “functional” assays of GPCR activation because GTP-binding indicates that a cellular response will occur, not just that a binding event was detected at the receptor (which may or may not lead to a cellular response). This is important for screening applications to detect novel compounds which activate or block GPCRs. Ligand regulation of the binding of [35S]GTPγS is one of the most widely used assay methods to measure receptor activation of heterotrimeric G-proteins, as discussed elsewhere in detail [19,20]. The move toward a fluorescent based Europium-GTP assay partly overcomes some of the limitations of radioactive-based assays and has already been successfully used with the following GPCRs, motilin, neurotensin, M1-muscarinic and α2A-adrenergic receptors [21,22].

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2.1.6. G-Protein Dissociation

GPCR biosensors can involve the detection of the final stage of activation of the G-protein heterotrimeric complex, that being the putative dissociation or rearrangement of the subunits following GPCR-induced G-protein activation [23,24]. This level of GPCR activation has currently not been investigated in great detail and may prove to be extremely valuable in future functional biosensor applications as G-protein signalling always involves alterations to the heterotrimer structure. An advantage of cell-free GPCR assays involving G-proteins is that GDP can be used to “reset” all GPCRs to the inactive state, thereby allowing for greater resolution of receptor activation by effectively removing background signalling.

Bieri et al. [25] used carbohydrate-specific biotinylation chemistry to achieve appropriate orientation and functional immobilization of the solubilized bovine rhodopsin receptor with high contrast micropatterns of the receptor being used to spatially separate protein regions. This reconstituted GPCR:G-protein system provided relatively stable results (over hours) with the added advantage of obtaining repeated activation/deactivation cycles of the GPCR:G-protein system, as occurs in vivo. Measurements were made using SPR detection of G-protein dissociation from the receptor surface following the positioning of the biotinylated form of the rhodopsin receptor onto a self-assembled monolayer (SAM) containing streptavidin. Although SPR is useful for the study of G-protein interactions, it may not be well suited to detect binding of small ligand molecules directly due to its reliance on changes in mass concentration. An advantage of repeated activation/deactivation cycles of GPCRs is that different compounds may be tested serially with the same receptor preparation, allowing for discernment of differential activation of the same receptors by different ligands. The above approach appears promising for future applications of chip-based technologies in the area of GPCR biosensor applications.

The well known and highly utilized interaction between Ni2+ and histidine residues (most often used for purification oligohistidixne-tagged proteins) may be a useful means of attachment for GPCRs and/or


G-proteins. To this end, different surface chemistries are being investigated to optimize the affinity interaction [26,27]. Modifying the surface of epoxy-activated dextran beads by forming a Ni2+-NTA conjugate was shown to produce beads with a surface capable of binding hexahistidine (his)-tagged β1γ2 subunits [28]. Tethered β1γ2 subunits were then used to capture Gαs subunits which in turn were capable of binding membrane preparations containing a β2-adrenergic receptor-GFP fusion protein. Alternatively, a fluorescent labeled ligand binding to the tethered β2-adrenergic receptor could be detected; the whole complex being measured using flow cytometry. Flow cytometry’s greatest advantage is its ability to be multiplexed, where different molecular assemblies can be made in one sample and then be discriminated by their unique spectral characteristics [10,28-30]. Indeed, particle-based nanotechnologies, e.g. quantum-dots [5,31,32], constitute another emerging enabling technology for GPCR biosensor applications.

2.1.7. GPCRs as Biological Detectors of Volatiles

Many organisms, from nematodes to mammals, use GPCRs to sense volatile compounds and the neural signal transduction due to GPCR-volatile interactions is the basis of smell [33-36]. Vertebrates are known to utilize olfactory receptors (ORs) that are similar to other metabolic GPCRs [35]. Invertebrate ORs were discovered relatively recently and it appears that these ORs are atypical GPCRs in that they have reversed membrane topology and that they apparently each form dimers with the same highly conserved OR-like receptor which can function as an ion channel, independent of ligand-mediated activation of the OR with which it has dimerized [37-39]. Due to the inherent OR-volatile specificity and high affinity of the OR-volatile interaction, ORs are obvious candidates as biomolecules that could be adapted to detect specific volatiles. Biosensors of volatiles would have a multitude of potential applications, particularly for sensing hidden entities (e.g. explosive screening) and in the agrifood industry (e.g. quality control, fermentation monitoring). The current attempts at volatile biosensing are not only limited by generic issues such as instability of membrane proteins but also by the poor level

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of functional characterization of the different ORs [40]. Therefore, current attempts to produce a sensor for volatile compounds are limited to the most highly characterized receptors, these being human OR17-40 [41], I7 from rat [42], and Odr10 from the nematode Caenorhabditis

elegans [43,44]. Hou et al. [45] reported an attempt to produce a biosensor that

utilized the membrane fraction of recombinant yeast expressing rat I7 receptor, which is known to be activated by heptanal and octanal. This design used a gold electrode that was functionalized with biotinylated I7-specific antibody, to which the membrane fractions were applied. Specific odorants were applied to the I7-presenting surface and interactions were monitored through variation in polarization resistance. Ligand binding was discernable although a specific response was difficult to resolve at ligand concentrations below 10-12 M, which is nevertheless more sensitive than current detection technologies. A common difficulty of presenting such ligands to the biosensor surface is the need for an organic solvent to carry the ligand. In the case discussed, dimethyl sulphoxide was used for this purpose and was found to only alter polarization resistance by 10%, even at the highest concentration tested (0.1 mM). Whole yeast cells have also been utilized to produce a I7-based biosensor [42]. Yeast cells were engineered to express I7 and a mammalian G-protein capable of linking ligand-mediated receptor stimulation to activate a MAP kinase pathway that induced synthesis of luciferase. Thus, in the presence of the luciferin substrate, ligand binding was detected as a dose-dependent fluorescent response. Importantly, sensitivity of the response was altered by the type of G-protein used to couple the OR to downstream elements.

Another design utilized the human OR1740 protein expressed in yeast and receptor-containing nanosomes were produced by sonication of recombinant yeast membranes [46]. The nanosomes were then captured on a SAM functionalized with biotinyl groups that were used for attachment of neutravidin and then a biotinylated monoclonal antibody specific to the receptor. Interestingly, the myc-tagged receptors were functional when immobilized via a C-terminal tag but not when attached by the N-terminus.


A crude Odr10-based olfactory biosensor was produced by expressing the OR in bacteria, obtaining a membrane fraction and coating a quartz crystal microbalance (QCM) chip with the membranes [47]. The authors of this study reported that ligand application could be detected as a change in QCM frequency. Another attempt utilized mammalian (HEK-293) cells to express Odr10, and SPR was then used to detect binding of the applied ligand (diacetyl) [48].

2.1.8. The Future of GPCR Biosensors

With growing interest and commercial investment in GPCRs in areas such as drug targets, orphan receptors, high throughput screening of drugs etc., greater attention will focus on biosensor development to allow for miniaturization, ultra-high throughput, and, eventually, microarray/biochip assay formats that will require nanotechnology-based approaches. The production of stable, robust, cell-free signaling assemblie’s comprising receptor and appropriate molecular switching components will form the basis of future GPCR/G-protein platforms which should be adaptable for laboratory- and field-based applications as microarrays and biosensors.

2.2. Pore-Forming Proteins

Ion-channels are transmembrane proteins that regulate the transport of ions and/or small molecules across the lipid membrane. They can exist in the open or closed state, which can be regulated by a range of stimuli, including a change in membrane potential, mechanical stress or the binding of a ligand [49]. Some of the most commonly investigated pore-forming peptides, are bacterial porins (e.g. OmpF) [50], α-hemolysin from the human pathogen Staphylococcus [51] and Gramicidin, a polypeptide antibiotic [52]. These have all been studied for their adaptability to a sensor platform [53-57]. Lipid supports, which are essential for the functional capture of the channels, are discussed in section 3.

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Figure 2. Pore-based Sensing. (a) Schematic of an ion-channel switch (sandwich) developed by Cornell et al. [59] (i) The lipid bilayer is composed of archaebacterial membrane spanning lipids (MSL) and half membrane spanning tethered lipids (DLP). MAAD are spacer molecules attached to the gold surface via a sulfur-gold bond. The mobile lipids (DPEPC/GDPE) and ion channels (Gα) attached to antibodies (Fab) via a streptavidin linker (SA), can move throughout the bilayer, unlike the immobilized ion channels (GT). (ii) Mobile channels (Gα) become cross-linked to tethered antibodies (Fab) on the membrane spanning lipids (MSLα) in the presence of analyte (A), preventing formation of complete channels and therefore decreasing measured current. Reproduced with permission [59]. (b) Braha et al. [57] demonstrated simultaneous stochastic sensing for zinc, cobalt and cadmium ions with an engineered protein pore. (i) Schematic representation of the pore with the single metal binding site (metals are represented by different sized, or filled, balls) in the lumen of the channel. Each time metal ions bind to the pore, the current is modulated, as illustrated in the trace which reflects the currents flowing through the pores that were recorded during the application of a +40 mV membrane potential. Arrows indicate the current through the fully opened pore (ii). Reproduced with permission [57].

a)

b)

(i) (ii)

(i)

(ii)


Native or engineered membrane-bound ion-channels are a promising class of receptors for biosensing applications as they allow the sensitive detection of analytes and produce an output (electrical current) which is inherently suitable for digitization [58]. Ion-channel switches, first reported by Cornell et al. [59] (Figure 2a) and single-channel stochastic sensors [58] (Figure 2b), both demonstrate mechanisms by which analyte detection and quantitation can be determined by measured changes in current due to ion-channel activity. An important thing to consider when utilizing ion-channels is that the current, produced by the movement of ions through a pore in a lipid membrane, is dependent on the accessibility of a clear passage through the pore and therefore, can fluctuate when ion-channels do not traverse the membrane or when the pores are partially or fully blocked.

2.2.1. Ion-Channel Switch

The ion-channel switch described by Cornell et al. [59] comprises a gold electrode, to which a lipid membrane carrying gramicidin ion-channels bound to antibodies, are tethered. A current is produced (turned on) when ions flow through the channel (in the presence of an applied potential). It is subsequently switched off when mobile channels diffusing in the outer half of the membrane, become cross-linked to antibodies immobilized at the membrane surface [59] (Figure 2). The ion-channel switch has since demonstrated specific signals derived from interactions with a range of analytes including bacteria, DNA, proteins and drugs [60]. Recently, Oh et al. [61] reported the detection of influenza A virus in clinical samples using the ion-channel switch biosensor.

2.2.2. Stochastic Sensing

Stochastic sensing involves monitoring current that flows through a single pore and the alterations of this flow in response to analyte binding events (Figure 2). Each time an analyte binds to a binding site within the pore, the current flow is altered and during the on/off equilibrium of the binding analyte, a characteristic flow pattern is produced and monitored.

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The frequency of the current fluctuations is related to concentration of the analyte, whereas the current signature (revealed by fluctuation duration and amplitude) is related to the analyte’s identity (reviewed by Bayley and Cremer [58] as well as Schmidt [62]). Some reported desirable attributes of stochastic sensing include fast detection, as well as the sensitivity and reversibility of sensor elements [57]. Current fluctuations monitored through a single pore, have demonstrated the capability for a variety of analytes, including metals [57], organic molecules [63] and oligonucleotides [64]. Shim and Gu [65] and Kang et al. [66] recently demonstrated increased stability of a single pore chip by encapsulation of the associated lipid bilayer within an agarose gel. This resulted in a more robust chip that can be stored for longer periods than bilayers alone. Pore-based transduction systems such as these have great potential in the field of biosensing.

2.3. Cell- and Viral-Based Sensing

The biological component of a biosensor is currently most often an enzyme, antibody or other sub cellular component (such as the receptors and ion-channels mentioned in previous sections). The purification of these proteins can be labor intensive, expensive and the resulting product incompletely purified or unstable. Whole-cell sensors preserve the localization and temporal control of protein function and can utilize reporting processes that may involve multiple enzymes and signaling cascades. These types of reporting system are advantageous when biological/metabolic relevance of an analyte is important rather than simply its detection. Thus, unlike purified enzymes and antibodies, cell biosensors can report on bioavailability, metabolic regulation, toxicity, genotoxicity (DNA damage) etc. The key challenge for cell-based biosensors is the maintenance of cell viability under assay conditions. This requirement may lessen the utility of cell-based systems in field applications or environments that are deleterious to cell viability.

Cellular responses can be specific to a substance (e.g. GPCR-ligand interactions) or a general response to adverse environmental conditions (e.g. regulated apoptosis) and each of these could be monitored depending on the specific biosensor design. Often cells are genetically


modified with a reporter protein such as luciferase, GFP or β-galactosidase. The genes for these reporters are placed under the control of a promoter that responds to the analyte of interest resulting in the expression of the reporter protein, which can be detected. This approach has been exploited in identifying many environmental pollutants that induce particular promoters and for measuring stress responses. With advances in nanotechnology, cell-based nanobiosensors are now emerging with increasing sophistication, sensitivity and detection methods that are available, although most are still in the proof-of-concept stage. Different cell types can be more applicable to certain applications and/or measurements. In the following sections we discuss nanobiosensors utilizing bacterial, yeast, fungal, algal and mammalian cells. Several examples of cellular GPCR-based (olfactory) biosensor designs are discussed earlier (see 2.1.7)

2.3.1. Bacterial Biosensors

Bacteria are particularly exploited in biosensors since they contain many defence mechanisms against analytes of interest such as environmental pollutants, including mercury and arsenic [67]. Bacteria can be easily produced at low cost and genetically manipulated and as mentioned previously. Reporter genes are often fused to DNA elements that respond to the presence of these analytes to produce a signal [67]. For example, bacteria can be used to detect genotoxic agents since increased DNA damage (due to the presence of a genotoxin) results in degradation of the endogenous repressor of SOS genes and subsequent expression of genes associated with DNA repair [68]. Other stress responses such as oxidation, nutrient starvation, membrane damage, heat shock and apoptotic responses can also be measured [67,68].

In a “lab-on-a-chip” format, E. coli were genetically engineered such that the activation of the fabA, dnaK or grpE promoters produced the enzyme β-galactosidase [69]. These bacteria were applied to electrochemical cells (100 nL capacities) on a silicon chip as a broth or immobilized within agar. The cells contained embedded electrodes for electrochemical measurements. Upon exposure to the representative toxicant phenol, and in the presence of the substrate

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p-aminophenyl-β-D-galactopyranoside, β-galactosidase was expressed because of activation of the relevant promoters. β-galactosidase cleaved p-aminophenyl-β-D-galactopyranoside into p-aminophenol and β-D-galactopyranoside. p-aminophenol is electrochemically active and the application of a 220 mV potential produces oxidation of p-aminophenol molecules, which is converted to a current that can be monitored to detect phenol concentrations as low as 1.6 ppm. The fabA promoter gave the largest response to phenol exposure which corresponded to the increased sensitivity of this promoter towards membrane damage inflicted by phenol. It is hoped that advantages such as the small sampling requirement, high signal to noise ratio, potential for high-throughput and high degree of robustness will combine to make this platform suitable for field applications.

Carbon nanotubes are also being exploited in bacterial nanobiosensors. Pseudomonas putida have been coated onto an osmium redox polymer on a carbon nanotube-modified electrode and covered by a dialysis membrane to form an amperometric biosensor with increased electron transfer efficiency [70]. The respiratory activity of the cells was correlated with the oxidation of glucose, measured via the osmium redox polymer that acted as an electron acceptor. The system was then modified to measure phenol, which could be detected in an artificial wastewater sample, using phenol adapted P. putida.

2.3.2. Fungal and Algae Cell Biosensors

The use of fungal cells (such as yeasts) in biosensors can provide the advantages of using bacteria but being eukaryotic cells they may provide information that is more relevant to higher eukaryotic organisms. This is a particularly important attribute for toxicity and drug screening. Fungal cells remain relatively easy to culture and genetically manipulate, and can be more robust with regard to pH, ionic strength and temperature than mammalian cells [71] due to their resistant cell walls. Wild-type cells can be used as biological oxygen demand sensors or to detect catabolic substrates. Oxygen consumption can be correlated to many physiologically relevant processes such as cell viability, protein synthesis and mitochondria function. Saccharomyces cerevisiae cells have been


immobilized onto amine functionalized polystyrene nanobeads that were loaded with the oxygen-sensitive fluorescent ruthenium (II) [72]. This allowed optical detection of oxygen consumption since in the presence of molecular oxygen, the fluorescence of ruthenium (II) was quenched. In the presence of high concentrations of glucose, there was increased fluorescence indicating that cellular respiration and oxygen consumption was increased compared to when glucose was absent.

Genetically modified cells can report on gene regulation in response to environmental factors or can be engineered to express and monitor the activation of other receptors such as those involved with olfaction or disease processes. An olfactory receptor (a GPCR) that responds to 2,4-dinitrotoluene (DNT), a mimic for the explosive trinitrotoluene (TNT), has been identified and yeast were engineered to express the receptor and its associated signaling components including the G-proteins, adenylyl cyclase and a cAMP responsive DNA element that promoted the expression of GFP upon stimulation of the receptor [73] (for more discussion of the application of explosives detection see section 7.3). S. cerevisiae has also been engineered to express the human olfactory receptor, OR17-40. These cells could be immobilized on interdigitated gold microelectrodes coated with poly-L-lysine and the conductance on the surface of the electrode was shown to be modified by receptor-ligand interactions [74].

Algae could potentially be exploited for environmental biosensing since the inhibition of algal photosynthesis can be correlated to toxic effects of pollutants such as herbicides. (for further discussion on environmental monitoring with biosensors, see section 7.4). Inhibition of photosynthesis by photosystem II (PSII)-inhibiting herbicides (e.g. atrazine) can be measured as a change in chlorophyll fluorescence, caused by these compounds blocking the PSII quinone-binding site thereby inhibiting photosynthetic electron flow. This approach was demonstrated by using Chlorell vulgaris entrapped on quartz microfiber filters, and using a fibre-optic bundle to monitor chlorophyll fluorescence, which increased in the presence of atrazine [75]. Changes in chlorophyll fluorescence in response to exposure to formaldehyde and methanol vapor have also been monitored in a biochip platform that

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could be used to simultaneously test many different algal strains with different sensitivities to toxicants [76].

2.3.3. Mammalian Cell Biosensors

Mammalian cell biosensors are finding particular usefulness within the pharmaceutical industry since using whole cells can maximize the information content of the assay and allow examination of a compound’s action on cells close to the intended target, within the context of all cellular machinery. This is particularly useful in the pharmaceutical industry that requires technologies that can provide reliable predictive information on lead compounds early in their development to reduce unwarranted development costs [77].

Impedance-based technologies (reviewed by McGuiness in 2007 [78]) can be used to detect cell death or proliferation, as well as smaller changes caused by receptor signaling and resulting in cytoskeletal rearrangements or changes in cell-cell interactions and adherence. Sensor chip-based impedance spectroscopy has been applied to measure the activation of GPCRs binding with neuropeptide Y (these receptors are implicated in human breast carcinoma [79]). Adenylyl cyclase activity in MCF-7 mamma carcinoma cells adhered to a microelectrode array, was stimulated with forskolin resulting in reduced impedance [79]. This effect could be blocked by pre-treatment with neuropeptide Y, which is known to inhibit adenylyl cyclase activity. Ligand-binding to various GPCRs in adherent cells has also been characterized using mass redistribution cell assay technologies (MRCAT) and resonant waveguide grating (RWG) [80].

Another example of a mammalian cell biosensor used malignant cells taken from a specific patient and exposed to chemotherapeutics to predict if that patient’s response will be favorable. Live metastatic human mammary cancer cells have been adhered to the gold surface of a QCM to sense for disruption of microtubules within the tumor cells in response to anti-tumor agents taxol and nocodazole [81]. To detect the excretion of interleukin-2 from mouse T-cells, silica nanoparticles were used to form a nanoparticle layer between two layers of gold, and antibodies specific to interleukin-2 were immobilized onto the sensor


surface. Concanavalin-A was used to trigger secretion of interleukin-2 from the cells which was detected by the antibodies with a limit of 10 pg/ml using localized SPR [82]. The authors suggest that this technique has potential to be applied to high-throughput cell analysis systems for reporting on various cell activities and functions.

These examples do not require cell engineering and are label-free, making them less invasive and therefore possibly more physiologically relevant. Other label-free technologies for whole cells have recently been more thoroughly reviewed [77]. However, most mammalian cell biosensors do rely on engineering cells to produce signals such as fluorescence, or over expression of a target that can produce an observable change in cellular physiology. These genetically encodable fluorescent biosensors have recently been reviewed [83] and will not be covered in detail here. Nanotechnologies are also being developed to sense changes within whole cells. Plasmonic biosensors or gold nanoparticles (20 nm) have been functionalized with an anti-actin antibody and a TAT-HA2 peptide, which mediated the endocytotic uptake of the nanoparticles into the cell, and their subsequent release from endosomes into the cytoplasm [84]. Binding of the nanosensors to actin could be measured since this brought the probes into such close proximity that the plasmon resonance became red-shifted, which could be detected by darkfield reflectance imaging or confocal microscopy, with detection limited to between 623-643 nm. It is hoped that further advances will enable monitoring of cytoskeletal rearrangements made as a biological response. Other optical sensing components can be entrapped within inert nanosized polymer particles termed PEBBLEs (probes encapsulated by biologically localized embedding) [85]. Possible probes include calcium sensitive dyes, pH sensitive dyes or enzymes such as horseradish peroxidase which can be used to detect reactive oxygen species [86] leading to applications such as analysis of effects of drugs, toxins or environment, on cell physiology. The inert polymer is permeable allowing the encapsulated probe to interact with analytes, and protects both the dye from interference by biological conditions, and the cell from any dye-associated toxicity. PEBBLEs are introduced into cells through surface modification with peptides that mediate cellular uptake, transfection using lipid reagents, picoinjection or gene gun

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bombardment. PEBBLEs containing the BME-44 ionophore, ETH 5350 chromophore, and ionic exchanger KTFPB, have been used to measure increases in potassium in rat C6 Glioma cells treated with kainic acid to open potassium channels [87]. Calcium-sensing PEBBLEs have also been use in SY5Y neuroblastoma cells to report on calcium released from mitochondria in response to exposure to the neurotoxin m-dinitrobenzene [88].

Currently, cell-based approaches usually utilize a large population of cells within which different responses are occurring, and the responses are averaged. However, the response of individual cells can be different in an environment that is free of influences from neighboring cells. Optical nanosensors are being developed to allow intracellular measurement of biological processes within single live cells. Tapered optical fibres with nanosized tips (30-50 nm) have been applied to a wide range of applications (reviewed in Leung et al. [89]) and can be used to probe conditions inside a cell. The tip of the nanofibre is approximately 10-fold smaller than the wavelength of excitation light transmitted along the fibre. Photons travel as far along the fibre as possible but cannot escape from the tip although an evanescent field continues to travel a short distance through the remainder of tip providing excitation light for molecules that are in close proximity (<100 nm) to the nanoprobe [90]. Optical fibres have been derivatized with a substrate of caspase-9, a biomarker of apoptosis that once cleaved, yields a fluorescent product. Manipulation of the fibre into MCF-7 cells exposed to an inducer of apoptosis (ALA-PDT) allowed the monitoring of caspase-9 activity detected by illumination of the fibre optic probe. The level of activity was indicative of the amount of fluorescent product formed in the presence of active caspase-9 [90].

2.3.4. Cell Immobilization and Arrays

Cells can also be immobilized to form arrays for advanced chip-based applications in medical diagnostics, or the detection of environmental pollutants in the field. Cell immobilization can facilitate the display of a variety of targets or enable measurement of physiological responses. Amphibian tumor cells (FT cells) have been cultured on a


Cr/Au microelectrode array chip [91]. Oxidation of norepinephrine (released by the cells upon stimulation with ATP via an endogenous receptor) and an increase in intracellular calcium were measured. In this example ATP was detected using an endogenous receptor, an approach that may produce results with increased physiological relevance. However, cells could also be engineered to express recombinant receptors.

Recombinant-cell microarrays have also been produced using suitable surfaces such as glass, silicon or tissue culture polystyrene that can undergo surface modifications (reviewed by Hook et al. [92]). Nucleic acid microarrays have also been produced and generally utilize DNA for expression of a desired protein or interfering RNA (RNAi). The relevant nucleic acid is firstly spotted and DNA-gelatin mixtures are often used to ensure spatial confinement of the nucleic acid molecules. Cells are then applied and allowed to attach to the surface. The nucleic acid detaches (or desorbs) from the surface by reversal of the hydrophobic (or electrostatic) interactions between the nucleic acid and the surface. A transfection reagent generally facilitates the nucleic acid being taken up by cells where it is then expressed or used to silence genes within the cell. Often the coding sequence of a reporter protein such as GFP is fused to the nucleic acid sequence of interest, allowing determination of transfection and/or expression levels. Transfected-cell microarrays have been applied to express GPCRs in a format where greater than 3000 receptor:ligand interactions could be measured in a single 96-well plate. The cells were loaded with a Fluo-4 calcium indicator dye and agonist binding was detected as an increase in intracellular calcium measured using fluorescence microscopy [93]. This functional screen could be applied to the “deorphanization” of GPCRs with no known ligand or for drug screening for characterized GPCRs.

In recent times, several polymers have been investigated as substrates upon which cells can be cultured, with the aim of culturing cells directly on precoated surfaces which could then be used as biomolecular detectors. Lakard et al. [94] tested three polymer substrates for their ability to allow rat neuronal cells to adhere and proliferate, namely polyethyleneimine (PEI), polypropyleneimine (PPI) and polypyrrole. Data indicated that PEI and PPI were the best candidates

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and did not affect cell morphology. In addition these polymers had advantages such as strongly binding to electrode surfaces and were insoluble in most organic solvents. Subsequently, the same researchers were able to culture neuronal cells expressing ORs in defined positions on a silicon wafer with the aim of producing an olfactory biosensor [95].

Single cells can also be trapped within minute depressions on a CD-like chip containing microfluidic channels along with passive valves and chambers allowing sample loading and waste storage [96]. Centrifugal force was used to load HEK293 or Jurkat cells into the depressions and assays performed using paraformaldehyde or UV irradiation to detect cytotoxicity or apoptosis, respectively. Compared to testing a larger population of cells, this single cell assay showed a higher rate of survival when cytotoxicity was measured presumably due to the lack of proteases and other toxicants released from neighboring lysed cells.

2.3.5. Virus-Containing Biosensors

While virus particles (virions) are not cellular, they are generally composed of a relatively complex mixture of peptides, exceeding that of the receptor based system, and may often be membrane-enveloped. Therefore we discuss virion-based biosensors at this point in the review, after the more complex cellular biosensors. Being proteinacious and genetically encoded, virions can be simply produced and can display peptides in a biologically functional form, either through genetic engineering or chemical conjugation of peptides or small molecules. In addition, virions can also be conjugated to fluorescent moieties such as a quantum dots or fluorophores for optical monitoring [97]. An important goal in the development of sensitive imaging sensors is the ability to specifically target cells and tissues of interest to allow sensitive imaaging or delivery of therapeutics. The abnormal characteristics of tumor cells produce cell-surface or extracellular matrix proteins that can be used as markers to distinguish tumor cells from normal tissue. Often the higher metabolic activity of tumors gives rise to over expression of a number of receptors such as the folic acid or transferrin receptors [97,98]. Virions have been utilized from Cowpea mosaic virus (CPMV), bacteriophages and other viruses that are not typically human pathogens and are


therefore unlikely to cause infection in humans, reducing costs associated with minimization of biohazards (e.g. quarantine laboratories and protective equipment) [97]. High yields of CPMV can be obtained from infected plants and virions are simple and inexpensive to purify. The virions are relatively thermally stable and can tolerate a wide pH range and a variety of solvents, making them useful candidates for field applications. By modifying virions for surface display of folate and blocking the remainder of the virion surface with PEG, the viruses could be targeted to tumors that over express folic acid receptors [98]. Philamentous bacteriophages also tolerate a wide range of conditions particularly high salt concentrations, low pH, presence of chaotropic agents and prolonged storage. Similarly to eukaryotic viruses, bacteriophages can be engineered for surface display of a desired peptide that could detect a specific cellular target or specifically modify the behaviour of a target cell, in a measurable way. Networks of phages and gold nanoparticles (44 nm in diameter) were observed to form spontaneously and phage particles displaying the peptide CDCRGDCFC, have been use to recognize αv integrins present in high levels on the surface of melanoma cells, where binding resulted in receptor-mediated phage internalization [99]. This was detected using dark field microscopy utilizing the large degree of scattering resulting from gold nanoparticles incorporated into cells (after washing), which are ideal contrast agents. Additionally, surface enhanced Raman scattering (SERS) spectra also correlated with the level of cell binding and internalization and could therefore also be used to monitor these processes.

3. Lipid Supports for Biosensor and Biochip Fabrication

3.1. Why Functionalize Biosensors with Lipid Membranes?

Lipid membranes are versatile and convenient alternatives to study the properties of natural cell membranes (see sections 2.1 and 2.2). Due to a combination of factors such as ease of formation, control over complexity, stability and the applicability of a large range of analytical techniques, artificial lipid membranes now have a central role in

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membrane research. Research on membrane sensor platforms has, in particular, been stimulated by the possibility of studying membrane proteins in a near-native environment, but it is also emerging as a nanoscale surface functionalization platform in its own right for controlled bioresponse. With more than 50% of all drug targets being membrane proteins, which require a lipid membrane to remain functional [100], target applications for commercial biosensors are dominated by the quest for high-throughput membrane protein drug screening assays and more predictive in vitro admetox platforms. Additionally, the development of artificial tongues and noses built on biological principles is also being sought.

Traditionally, lipid membranes have been regarded as having two important biological functions: (i) acting as an electrochemical barrier between cells and the environment and between different cellular compartments; and (ii) as scaffolds for membrane proteins. In recent years, it has been increasingly realized that the dynamically rearranging lipid membrane, containing a multitude of different molecules, could on its own perform important messenger and switching functions [101,102]. Its role in controlling protein and cell function might be much greater than previously thought [102]. This has led to an interest in more sophisticated membrane sensor configurations which offer the possibility of obtaining richer information about the biophysical state of the membrane, in real time. Desired features include ability to measure/characterize changes in density, thickness and ordering of the lipids, sub-micron domain chemical composition, membrane asymmetry, surface charge, mechanical properties and morphology. Obtaining this kind of information with the necessary level of detail requires sensing at the nanolevel and in many of the approaches now being developed, nanostructured sensors play a key role.

This section primarily describes lipid membrane architectures, which can be used to study membrane properties at, or in close proximity to, a solid interface. Interest in such systems is high due to their stability and large range of highly sensitive analytical label-free methods available to characterize the membranes and their associated interactions or alterations. In particular, the integration of membranes with electrochemical and optical readout schemes allows for simultaneous


measurements of binding, conformational changes and functional response of membrane and incorporated membrane proteins.

Several different surface-based membrane platforms have been developed over the years. They can be divided into hybrid lipid bilayers (Figure 3a), solid-supported lipid bilayers (SLB, Figure 3b), tethered lipid bilayers (tSLB, Figure 3c), polymer-cushioned lipid bilayers (pSLB, Figure 3d), pore-spanning lipid membranes (nano- or micro-BLM, Figure 3e) and tethered liposomes (Figure 3f).

Figure 3. Classification of different membrane functionalized sensor platforms. (a) Hybrid lipid bilayer with a lipid monolayer formed onto an alkane self-assembled monolayer (SAM) functionalized sensor substrate; (b) supported lipid bilayer (SLB) self-assembled on hydrophilic support; (c) tethered lipid bilayer (tSLB) self-assembled on covalently attached hydrophobic molecules, often derived from lipids, with a hydrophilic spacer layer attached to the substrate; (d) polymer-supported lipid bilayer (pSLB) assembled either (i) directly on a polymer cushion with adjusted wetting behavior or (ii) using hydrophobic molecules as anchors that are incorporated into the polymer matrix; (e) pore-spanning lipid membrane assembled across a nano- or micro-sized aperture in a support.; (f) tethered liposomes.

3.2. Methods to Assemble Supported Lipid Membranes

With the exception of tethered liposomes (hollow phospholipid bilayer vesicles self-assembled in water from amphiphilic molecules [103]), which only require adequate substrate attachment, the formation of a lipid membrane on a sensor surface requires detailed control of the

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interaction of the physisorbed lipids (and incorporated proteins) with the substrate. Several methods have been developed to self-assemble sensor-supported lipid membranes, and are schematically represented in Figure 4: self-assembly from vesicles in solution (Figure 4a) [104,105], Langmuir-Blodgett type deposition (Figure 4b) [106] and self-assembly from lipid dispersion directly on the surface by painting and detergent dialysis (Figure 4c) [107]. Additional methods such as stamping of a membrane (including native cell membranes) onto a surface have also been developed but do not yield high and homogenous coverage beyond that which allows qualitative binding studies [108].

Figure 4. Methods used to assemble lipid membranes on substrates. (a) Self-assembly from vesicles on (i) hydrophilic substrates, (ii) hydrophobic substrates and (iii) hydrophobic tethers with hydrophilic spacers (tSLB); (b) assembly from Langmuir films to (i) first monolayer deposited on hydrophilic substrate, (ii) bilayer completed by Langmuir-Blodgett deposition, or (iii) bilayer completed by Langmuir-Schäfer deposition; and (ci) detergent dialysis, (cii) painting and solvent extraction.

The original and possibly still most used methods are based on spreading the membrane from a solvent onto a surface (or across an aperture in a support). However, these methods are becoming increasingly replaced by those relying on assembly from pre-formed small (20-200 nm in diameter) unilamellar liposomes, which are fused into an SLB (for example) [109,110]. Also, Langmuir-Blodgett deposition techniques are less used, because they require more complex setups and controlled environments and thus typically have lower


reproducibility despite the advantage of control over membrane tension during deposition. The main reason for the emergence of self-assembly from vesicles (compared to solvent spreading methods) is that it allows for the formed membrane to remain free of solvents, which render many transmembrane proteins non-functional. Additionally, liposome composition and other properties can be easily controlled, including by the use of vesicles harvested directly from cells, and thus provides a simpler and more versatile way of controlling the composition and lateral distribution of lipid and protein material on the sensor surface.

3.3. Supported Lipid Membrane Platforms

The literature on sensing using supported lipid membranes is vast and only the wider context and a few selected examples are discussed here. For an in-depth review of the many suggested membrane sensor platforms, we refer the reader to a recent review by Janshoff and Steinem [111] or to topical reviews on specific configurations [112,113]. The first developed and most widespread in vitro method for studying ion-channel function is the bilayer lipid membrane (BLM), spanning across apertures between two aqueous compartments [114]. While a successful approach under laboratory conditions, this platform has a few major drawbacks for general application: (i) instability manifested in the collapse of the membrane within a few hours of preparation; (ii) preparation requiring solvents which stay in the membrane, making it incompatible with many proteins which alter conformation and function in the presence of solvent; and (iii) limited to only electrochemical and fluorescence based sensing techniques. As a result, there has been an increasing emphasis on solid-supported membrane platforms in recent years.

Solid-supported platforms offer inherent stability thanks to the underlying support. In their original form, these solid-supported bilayers [115-118] (hybrid bilayer Figure 3a; SLB Figure 3b) had a severe drawback for studying membrane protein function in that they offer very little space between the lipid membrane and the solid support to accommodate hydrophilic domains of the integrated protein and ions transported across the membrane [111,119], although their assembly on electrodes [115,120-122] and use for characterization of inserted

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membrane proteins [123-128] have been demonstrated. However, due to their simple geometry and proximity to the surface they offer an excellent model system for studying the self-assembly of lipid membranes and to probe membrane physicochemical properties by a wide range of techniques [105,109,111,113,129-131]. Further, decoupling of the membrane from the surface by forming a tethered supported lipid bilayer (Figure 4c), first demonstrated by Vogel and co-workers [132], has been achieved primarily by forming a supported lipid bilayer onto a hydrophobic anchor layer separated from the surface by a short hydrophilic polymer spacer [59,111,133-144], typically oligo(ethylene glycol) with a lipid anchor [111,132,134-137,145,146]. The additional space of a few nanometers allows for integration of small transmembrane proteins without undesired surface interaction, but the reservoir is still too restricted to monitor continuous ion transport and the preparation can be difficult for solvent-free membranes [145]. Knoll and coworkers have shown how to reproducibly achieve high-insulating tSLBs by creating membranes on ultra-flat substrates and improving stability by using tethers derived from phytanoyl thiolipids inspired by those found in members of the Archaea [145,146]. A competing platform using similar tethers based on cholesteryl anchors has been shown [136,147]. Tethering membranes directly to pre-immobilized transmembrane proteins, which has the advantage of yielding 100% orientation of proteins, has also been shown [133,148]. Recently, the fusion of native membrane fragments to form the upper leaflet of tSLBs has also been reported with high sealing resistance [149]. This method avoids the needs for reconstitution of sensitive proteins.

Membranes can also be formed on top of hydrogels and other thicker (∼10 nm) polymer cushions (Figure 3d) [113,150-153]. However their application to biosensing is so far limited. Tanaka et al. [113,154,155] have demonstrated the formation of lipid membranes on hydrogels (mainly cellulose derivatives), but cushions such as PEI, PEG and pH-responsive polymers have also been used [153,156-159]. These membranes, also formed directly from erythrocyte ghosts, have been shown to retain the mobility, density, function and orientation of transmembrane proteins containing large hydrophilic domains, like cell-adhesion receptors [155,160]. Thick, fully decoupling, hydrophilic


spacer layers at the sensor substrate can also be obtained by self-assembling the polymer with the SLB, without covalent binding to the substrate. Methods based on self-assembly from PEG-lipsomes and stabilization with trehalose to create robust membranes, have been demonstrated by Albertorio et al. [161,162].

Due to advances in sensor nanofabrication technology and in lipid self-assembly, free-spanning membranes are currently experiencing a renaissance as nano-BLMs (Figure 3e). Fabricating apertures in the sub-micron range, allows for greatly enhanced membrane temporal stability [112,163,164]. Furthermore, decreased size and advances in nanoscale control of the surface properties make it possible to self-assemble nano-BLMs from small unilamellar vesicles. Several publications report the development of platforms based on nanoporous solid substrates [152,165,166]. The best results have been achieved by forming the membrane after rendering the top surface of anodized alumina foils with dense sub-100 nm orifices made hydrophobic by an alkanethiol SAM. The system thus comprises spanning membranes sealed and stabilized to the solid support by forming a hybrid bilayer. Nano-BLMs are stable for tens of hours and activity of single ion-channels inserted into the membrane can be measured [164,166-168].

A recently proposed method to completely decouple the membrane from the substrate but still keep it in proximity of the surface is to tether intact liposomes to the sensor substrate through several hydrophilic linkers (Figure 3f) [169-172]. While early demonstrations utilized tight binding by for example lipid-biotin-avidin linking, the major linking strategy has become the use of complementary DNA-anchors that allow for tagging and addressing libraries of different vesicles [169,173,174]. Importantly, the size of tethers and vesicles is suitable to capture the membrane system within the evanescent sensing zone of optical and acoustic sensors [175,176]. When tethered to an SLB on the sensor, liposomes retain high lateral fluidity allowing liposome-liposome interactions (and therefore potentially those between inserted protein) to be studied [177,178] while immobilization onto micro- or nanopatterned surfaces creates stable arrays. Stamou and others demonstrated addressing of single liposomes per spot onto large areas with applications in, for example, affinity studies for GPCRs [170,179] (also see section

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2.1.2), and recently similar addressing plus sorting of functionality was also shown [174,176]. The main drawback of the tethered vesicle system is that it does not allow for detailed measurements of membrane ordering or transport of material across the membrane.

Micropatterning of planar lipid membranes on solid and polymer supports has also been demonstrated, which although not yet developed with the same ability to label and target a certain functionality to an array spot (as allowed by tethered vesicles), nevertheless shows promise as a potential strategy to selectively assemble planar membranes and confine them by materials contrast [180], microcontact printing of membrane or diffusion barriers [154,181-184], photopolymerization [185,186], polymer wettability contrast [154], “nanoshaving” by AFM [187,188], polymer lift-off [189], microspotting onto tethers [147] or microfluidics [189-191]. At least the two latter methods allowed for arraying of membrane function and subsequent array bioaffinity sensing.

3.4. Advanced Sensors Functionalized with Lipid Membranes

The discussion of membrane platforms in earlier sections concentrated on the study of transmembrane proteins and charge translocation across the membrane. Despite the many advantages for functional biosensing conferred by electrochemical sensing methods, not least of which is the ability to characterize ion-channel and ion-transporter function [192] (see also section 2.2), the bulk of work on lipid bilayers has been on characterizing the assembly and properties of assembled membranes using optical and acoustic sensing methods. Typical methods to characterize the formation of planar lipid membranes include fluorescence recovery after photobleaching (FRAP) [193], QCM with dissipation monitoring (QCM-D) [194] and electrochemical impedance spectroscopy (EIS) [140]. These methods facilitate probing for completeness (FRAP, EIS), lateral fluidity (FRAP) or formation kinetics and mass (QCM-D). Sub-micron miniaturization of the membrane combined with the realization of the importance of lipid organization (especially for complex biological mixtures) makes for example, optical and acoustic evanescent probing of lipid distribution and membrane conformation increasingly interesting. Recent data from


self-assembly of supported lipid bilayers demonstrate that combining different biosensing techniques also provides the potential to perform detail characterization of morphological transitions from vesicular layers to planar bilayers [105] and of morphological changes in a supported lipid bilayer itself [130]. Furthermore, advances in waveguide spectroscopy have made it possible to probe transitions in supported membrane structures in terms of thickness, density and lipid alignment [129,131,195,196]. This information can be used for biosensor applications involving influence of ligand binding, drugs and ions on membrane properties (see GPCRs and ion channels; section 2). As mentioned earlier, this methodology could differentiate between different ligands binding to GPCRs reconstituted into waveguide-supported lipid bilayers by the specific induced conformational changes to the GPCR and surrounding membrane [197]. With the advent of nanooptical (in particular nanoplasmonic) sensors, it has been demonstrated that single liposomes and supported lipid bilayer islands can be confined to a single nanoplasmonic sensing element [198,199]. The high local sensitivity of such sensors combined with approaches such as the liposome arraying technology described earlier makes it possible to envision ultra-dense membrane affinity or more advanced array sensor platforms in the future.

3.5. Future Perspectives

The increasing number of publications in recent years demonstrates the interest for lipid membrane based biosensors. A recent trend in the field is the utilization of nanoscale optical and electrochemical sensor architectures [168,198] as well as microfluidics [200,201]. Future developments will address the miniaturization of sensor elements and production of membrane arrays. For integration into commercial biosensors, especially for field applications, an important goal is to improve the robustness of membrane functionalized sensors by increasing their stability [66,147,202,203]. An additional area of further development is the creation of more sophisticated sensor integrated membranes which mimic a greater range of biological membranes and their properties. Such functionalization could greatly enhance our understanding of the molecular basis of membrane function in biology.

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Combined with sensor platforms that probe not only the affinity of ligands but also the biophysical properties and organization within the membrane, direct probing of spatial and temporal dynamics of membrane signaling at the nanolevel could be another area of investigation [80].

4. Nanopatterning for Biosensing and Biochip Fabrication

A number of nanopatterning technologies are being developed in order to control the location, distribution, amount or conformation and orientation of biomolecules in the nanorange [204]. The sensitivity and selectivity of a biosensor relies on the specific interaction between biomolecules, hence uncontrolled, non-specific interactions have to be suppressed in order to avoid false responses [205-216]. Parallel nanofabrication approaches enable fast production of a large number of samples that can be applied over a large surface area. Alternatively, serial nanofabrication methods, although slower, usually offer better control over size and composition. In addition, the combination of top-down approaches with self-assembly (bottom-up) concepts is further increasing the flexibility to position biomolecules onto a pre-defined area. The most prominent nanofabrication methods will be addressed in this section.

4.1. Parallel Nanopatterning Methods

The ability to produce high-quality, large scale nanopatterns quickly and cheaply is challenging. A large number of different nanopatterning methods are being considered for novel biosensing platforms including the next generation of photolithography (see section 4.1.1 e.g., Extreme Ultraviolet Interference Lithography [217]), soft lithography (see section 4.1.2 e.g., replica molding and microcontact printing) [218,219], nanoimprint lithography (see section 4.1.3) [220,221] and nanosphere lithography (NSL, see 4.1.4 e.g., colloid lithography [222] or colloidal block-copolymer micelle lithography [223]).


4.1.1. Photolithography

Photolithography is a well-established patterning method enabling fast and highly reproducible creation of micron- and sub-100 nm structures by illuminating a photosensitive polymer. In particular, while illumination through a mask is successful in the micron-range, maskless approaches (so called interferometric lithography) were found to be valuable for creation of periodic large-scale nanofeatures over a large surface area [217,224-227]. Subsequently, biological contrasts (i.e. active nanopatches) embedded in a non-interacting PEGylated background could subsequently be incorporated into such a pre-pattern [228]. Alternatively, photolithography can be used to create nanowire- [229,230] or carbon nanotube- [231] field effect transistor biosensors (see section 5 for more detailed discussion of nanotubes).

4.1.2. Soft Lithography

Microcontact printing is a popular parallel micron- and nanopatterning approach, which was first introduced in 1993 by Whitesides and coworkers who created alkanethiol patterns on gold [232]. This three-step patterning method consists of: (1) production of the re-usable master, (2) formation of the elastomeric stamp from the master and (3) the “inking” of the stamp and printing of its features onto a substrate [219,233]) which enables patterning of a large variety of biomolecules (proteins [234-238], DNA [239-241], supported lipid bilayers [182,183,242] or liposomes [171]) crucial for creating biosensing platforms in the micron- and nanorange. Recently, an approach to automate the microcontact printing process has been reported, making soft lithography even more competitive (inexpensive and higher throughput) when compared with photolithography [243]. Supramolecular nanostamping is an alternative printing approach introduced by Yu et al. [244], which enables high-resolution DNA nanopatterning [245-247].

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4.1.3. Nanoimprint Lithography

Nanoimprint lithography [248-250] and closely related techniques [251-253] facilitate imprinting of a rigid nanostructured mold into a thin layer of spin-coated resist. The excess resist is then removed in the deformed areas via reactive ion etching (pattern transfer). There are different ways that nanostructures created via imprint lithography can be used for biosensing platforms. They have been utilized in a “lab on a chip” device providing a nanoscaled channel network. The imprinted (polymer) contrast was used as an etching mask [254-256] or a step in a lift-off process [257,258] in order to provide access to a bio-active nanostructured surface for biosensing applications. Alternatively, PEG-based UV-curable polymers have been directly imprinted and further decorated with proteins [259-262].

4.1.4. Nanosphere Lithography

NSL is a bottom-up approach enabling cheap, fast and large-scale production of nanopatterns via colloidal self-assembly [222]. The concept was introduced in the early 1980s and is simple and straightforward and utilizes a particle monolayer as a mask for contact imaging [263], etching or material deposition [264-267]. In recent years, impressive advances have been made to facilitate defect-free particle monolayers and subsequently, high fidelity bio-active nanopatterns [268-275]. NSL is not only interesting for the creation of high-density nanoarrays of biomolecules, but also for nanopatterned metal colloids that were used as read-out systems for biosensing applications based on localized surface plasmon resonance (LSPR) [276-278], as demonstrated by Van Duyne and coworkers [279-282] or Frederix et al. [283,284]. By monitoring changes in the UV/visual absorption band of the nanoparticles adsorption of chemical or biological species could be detected. For instance, amyloid-β derived diffusible ligands (ADDLs) could be detected in the cerebrospinal fluid of an Alzheimer’s disease patient using a nanoscale optical biosensor [279] (Figure 5). The “sandwich assay” was used as a biosensor where the signal is generated by a tagged reporter molecule binding to a biological detector that is


Figure 5. Biomarker detection using a nanoscale optical biosensor. (a) Illustration of the biosensing design for the detection of amyloid-β-derived diffusible ligands (ADDLs) using a sandwich assay (schematic in the top right corner). The changes of the optical properties (LSPR) of silver nano-triangles created via NSL (atomic force microscopy image on the bottom right) upon the adsorption of the biomolecules were monitored. (b) A sandwich assay and a LSPR nanosensor (i) were used to analyze human cerebrospinal fluid (CSF) from an aging person (ii) and an Alzheimer’s disease patient (iii). The LSPR spectra of each adsorption step were monitored and the presence of ADDL was detected only in the sample of the Alzheimer’s disease patient (shift in the LSPR spectra in (iii). Reprinted with permission [279].

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attached to a surface with linker molecules (often oligonucletoides or proteins). Alternatively, holes in a conductive gold film created via NSL were utilized in LSPR and were shown to be a sensitive detection option for several bio-recognition events [199,285].

4.2. Serial Nanopatterning Methods

Direct-writing techniques such as e-beam lithography [286], focused ion-beam lithography [287] or dip pen nanolithography (DPN) [288,289] can virtually produce any type of nanostructure with resolution between 5-50 nm. Although e-beam lithography and focused ion-beam lithography are able to create a large variety of nanopatterns [290-293], the cost and time to write them are disadvantageous when compared to the parallel nanopatterning methods described in section 4.1. These important disadvantages have seen the use of these two nanopatterning techniques limited. DPN, on the other hand, has become a very popular and versatile nanopatterning approach since its invention in 1999 [294]. DPN is a direct-writing, scanning probe based technique where an inked AFM tip is used to transfer biomolecules to a pre-defined location on a surface. Using this approach, nanoarrays of DNA [295,296], peptides [297,298], proteins [299,300], lipids [301], viruses [302-304] and bacteria [305] have been created. In addition, enzymes patterned via DPN have been used to perform localized reactions on a surface [306-308]. In order to produce nanoarrays for parallel high-throughput screening, the inherent slowness of the process when using a single AFM tip is being overcome by either using an array of individually actuated tips [309-311] or of passive tips. While the former concept enables the creation of complex chemical patterns, the latter is much simpler and was recently used to create 450,000 sub-100 nm features in less then 30 minutes (Figures 6a and 6b) [312]. Figure 6c shows an example whereby massive parallel DPN was used to create a large-scale nanoarray of supported phospholipid bilayers. The recent advances in parallelization of the patterning approaches together with the fact that virtually any biomolecules can be arranged into any nanoshape, offers unique opportunities for designing multi-component biosensing platforms.


Figure 6. Dip Pen Nanolithography: (a) Schematic illustration of massive parallel dip pen nanolithography using a 2D cantilever array. (b) A micrograph and a SEM (inset) image of such a cantilever array. Reprinted with permission [312]. (c) Fluorescent images of supported phospholipid bilayer patterns; (i) lower magnification image, (ii) close-up in the array, (iii) a two-component bilayer nanoarray using two different inks (phospholipids doped with two different fluorescent dyes) is shown. Scale bars are 5 µm. Reprinted with permission [301].

All of the parallel and serial nanopatterning methods discussed in

section 4 have the potential (in combination with an appropriate surface chemistry) to be implemented as a biochip especially when integrated with other bio-analytical components into a small portable device, a “lab on a chip” [313-317]. Such nanoarrays are expected to significantly impact future biosensing applications. However, the reliable and large-scale production of a heterogeneous nanoarray still needs to be achieved. Without this, a meaningful nanoarray based biosensing platform cannot be established since commercial drug discovery or diagnostic screening

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applications require sensitive and selective parallel analysis of a large number of biomolecules and their interactions in a short time.

5. Sensing Substrates: A Closer Look at Nanotubes

The term nanotube is now used to describe a variety of hollow structures made from a wide range of materials including carbon [318,319], boron nitride [320], titanium dioxide [321], silica [322] and even “soft” matter such as peptides [323]. Many of these structures have unique properties that have been exploited for biosensing [322-326]. This discussion will concentrate on carbon nanotubes as they are easily the most utilized structure for development of biosensors, and many carbon nanotube-based approaches have been applied to detect a range of analytes [319]. We focus on electrochemical approaches as this represents the most active (and best developed) area of biosensing with carbon nanotubes.

The great promise of nanotubes as biosensing elements is the potential to develop systems where direct electron transfer between enzymes and electrodes is possible. This innovation is key to the development of mediatorless (third-generation) enzyme biosensors, where no co-substrate is required in the recycling of the enzyme back to its active form. The mediatorless enzyme biosensor using nanotubes is most obviously applicable to the oxidoreductase enzymes where redox reactions cause electron flow and the extremely high conductivity of the nanotubes is used to detect this flow.

5.1. Carbon Nanotube Electrodes for Communicating with Redox

Proteins

Carbon nanotubes consist of graphene sheets wrapped into a hollow cylinder with the ends capped or open [318,324]. In the case of multiwalled carbon nanotubes (MWNTs) the concentric graphite tubules are in the range of 2 to 25 nm in diameter with 0.34 nm between tubule sheets. With single-walled carbon nanotubes (SWNTs) a single graphene sheet is rolled seamlessly into individual cylinders (typically of 1-2 nm) with capped ends containing carbon atoms which are all sp2. SWNTs


can be metallic conductors, semiconductors or small-band gap semiconductors depending on their diameter and chirality [318]. Closed nanotubes can be opened in oxidizing environments such as nitric acid [327]. Open-ended nanotubes have been shown to have excellent electron transfer properties [324] compared with closed nanotubes. The open ends of the carbon nanotubes typically contain carboxylate and quinone functionalities in common with edge-planes of pyrolytic graphite, allowing linking of functionalized nanotubes with edge-planes of pyrolytic graphite, while the nanotube walls have similar electron transfer properties to the basal planes of pyrolytic graphite.

Electrodes have been made using either MWNTs or SWNTs. In many nanotube electrodes thus far presented in the literature, the electrode is prepared by forming a paste with a filler compound and packing this into an electrode body, or simply by dispersing the tubes in a solvent and drop-coating onto the electrode to leave a bed of nanotubes on the electrode surface [328,329].

The first example of achieving electron transfer to proteins using carbon nanotube-modified electrodes was by Davis et al. [330] where an electrode of MWNTs was first opened in nitric acid and then mixed with nujol, bromoform, mineral oil or water. Cytochrome c and azurin were subsequently adsorbed onto and/or within the tubes with retained activity. The electrodes were shown to have an excellent ability to probe the redox sites of these proteins which was superior to that provided by edge-plane pyrolytic graphite. Similar results have been obtained by others who probed redox proteins with their active sites close to the protein surface, such as cytochrome c [331,332] and horseradish peroxidase (HRP) [333,334].

A more recent example [328] of this approach initially coated SWNTs in the biocompatible polymer chitosan, which has the dual effect of making the nanotubes more dispersible in water in addition to removing the hydrophobic character of the outer nanotube surface allowing adsorption of a biological molecule, in this case glucose oxidase. Using this approach, the detection limit of glucose was 0.01 mM with a response time of 10-15 seconds. The electron transfer rate of this electrode type was also shown to be superior to similar MWNTs or nanoparticle electrodes highlighting, the importance of the reduced size

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of the SWNTs which likely allows a closer approach to the redox active centre of the probed enzyme. This is advantageous as most redox active biological molecules have their redox centers embedded deep within the protein’s quaternary structure [335]. For example, in the case of glucose oxidase, the smallest distance between the protein exterior and its redox active center, Flavin adenine dinucleotide (FAD), is 13 Å [336]. Consequently, electrons cannot be efficiently transferred between the enzyme and the electrode and hence, mediators or redox relays are required. The use of nanotubes overcomes this obstacle by essentially “plugging into” the enzyme and getting close to the active centre which facilitates efficient electron transfer.

Boron-doped nanotubes [337] have been used to detect glucose with high sensitivity and selectivity, and importantly, in blood plasma with little sample preparation. The low potential at which the glucose is observed allows its detection with only minor disruption from common interferents such as ascorbic acid, acetaminophen and uric acid, giving improved resolution of the electrochemical signal.

The approach to attachment of biological species to carbon nanotubes is generic and should work for most biomolecules, which means the range of potential applications is enormous. Recent work utilized organophosphorus hydrolase (OPH) non-specifically bound to horizontally-aligned SWNTs on a SiO2 substrate [338]. This electrode hydrolyses organophosphates (OPs) many of which are used as insecticides and other pesticides [339]. Changes to OPH upon exposure to OPs cause changes to nanotube conductance, which is monitored to determine a response. This type of electrode gives a real-time response and has been used for multiple analyses.

Currently, a key challenge is to produce multiplexed electrodes that allow monitoring of two or more active species (for example, by co-adsorbing enzymes [340]) and facilitating simultaneous detection of a number of analytes. Ultimately, it could be possible to co-adsorb enzymes and their mediators, or perhaps enzymes and cofactors, to build smart electrodes which only are activated in certain situations that cause release of the activating element.

Studies involving direct electron transfer to enzymes, illustrate the potential advantages of carbon nanotubes modified electrodes. However,


these studies employed randomly entangled nanotubes which give a poorly defined electrode surface and poorly defined protein immobilization. Aligned nanotubes electrodes will provide a more controlled surface upon which to immobilize thus improving communication with redox proteins [324,341]. Additionally, covalent attachment of the bioentity to the nanotubes promises to significantly increase efficiency of electron transfer, especially for the cases where distances between the redox centre and the probe are relatively large, for example, when biomolecules are attached to the apex of nanotubes. One approach is to vertically-align nanotubes using chemical vapor deposition (CVD) to produce vast arrays of long nanotubes [342]. Biosensors based on this design have been demonstrated to for example, detect glucose [343]. The main drawback of this approach is that high temperatures are required to make the nanotubes and adhesion of the nanotubes to the substrates is often weak, meaning that their use in long-term or field applications is questionable. Chemically producing the vertically-aligned nanotube arrays has the advantage of using a strong covalent bond for nanotube attachment, making further modification of the nanotubes straightforward.

5.2. Aligned Carbon Nanotube Electrodes for Direct Electron Transfer

to Enzymes

One the earliest examples of the fabrication of aligned carbon nanotubes electrodes used a short thiol to make a amine terminated SAM on a gold surface [341] which could be further reacted to attach acid-functionalized SWNTs [327] to the substrate. The free end of the nanotubes, which still had available active groups, was then covalently bound to the enzyme microperoxidase MP-11 [341]. Attaching MP-11 to the aligned SWNT modified gold electrodes and subsequent electrochemical interrogation showed peaks characteristic of the heme redox active center of MP-11. Further to this, Willner et al. [344] attached the active centre of glucose oxidase (FAD) to the end of a vertically-aligned nanotube and then reconstituted the bound enzyme by wrapping apo-glucose oxidase around the FAD. This construct was able to detect glucose. Other early work detected hydrogen peroxide with

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myoglobin- or horseradish peroxidase-modified nanotubes attached to pyrolytic graphite electrodes [345].

A disadvantage of using thiol-gold based SAMs is that despite the strong interaction previously observed between a thiol and gold [346], studies of alkanethiols on gold have revealed that they are susceptible to thermal instability [347,348], UV photoxidation [349] and adsorbate-solution interchange leading to poor long-term stability [350]. Therefore, the ability to use a different substrate for thiol attachment would be desirable. In view of the importance of silicon as the primary semiconductor material in modern microelectronic devices, efforts to control its electronic properties and tailor the chemical and physical characteristics of its surface are of major importance. Early work in this area reported the preparation of well-aligned carbon nanotube arrays on silicon (100) surfaces by reaction of hydride-terminated silicon (100) with ethyl undecylenate, producing SAMs that were linked by stable silicon–carbon covalent bonds [351]. However, the presence of a SAM of organic material hinders electron transport between carbon nanotubes and the underlying silicon substrate.

Figure 7. Nanotube sensor substrates. Schematic of general approach to attachment of a biomolecule to aligned carbon nanotubes anchored to a silicon substrate. Initial attachment of the nanotube to the silicon is done via a condensation reaction between surface –OH groups and the carboxylic acid groups of the oxided carbon nanotubes. Subsequently, unreacted acid groups on the nanotube are available for further modification to directly attach the biomolecule or as shown in the figure condensation reactions allow attachment of an intermediate which has a high affinity for the bioentity to be attached.

A new approach to covalently attach carbon nanotubes to silicon

(without the use of intermediate molecules) has been developed using hydroxyl terminated silicon as the substrate [350]. This approach yielded


vertically-aligned, shortened carbon nanotube architectures on a silicon (100) substrate. Compared to older techniques, the new approach has several advantages including the lower temperatures involved in preparation and the possibility for subsequent modifications. Electrochemical analysis of this interface demonstrated excellent conductivity to the substrate, a factor that is likely to see this approach adopted for numerous potential applications [352,353]. The attachment of SWCNTs directly to the silicon surface provides a simple and novel avenue for the fabrication and development of silicon-based electrochemical and bio-electrochemical sensors. As outlined in Figure 7 and earlier in this section, the approaches to attach biomolecules to functionalized nanotubes are well established [324,341] and future research will be aimed at further developing this biosensor platform.

6. Reporter Technologies: Nano-Sized Labels for Biosensing

Applications

Various transduction techniques including electrochemical, optical, piezoelectric, and thermometric have been applied to development of biosensing platforms. Many of these have been mentioned throughout previous sections in this review. Reviews on these different groups of biosensor transduction systems, including optical [354] (e.g. SPR [355]), piezoelectric [356], electrochemical [192,357,358], and thermal [359] can be found in the scientific literature.

The demand for unsophisticated, low cost, portable yet sensitive biosensing devices, has directed biosensor research towards development of novel biosensing platforms with improved signal generation and transduction mechanisms. To this end, nanoparticle labels conjugated to reporter molecules show great potential for generation of strong optical or electrical signals upon biomolecular recognition. They have therefore been explored as promising alternatives to conventional labels that commonly include enzymes, fluorescent dyes or radioactive conjugates. There are a range of general reviews available on the use of nanoparticles as reporters [360-368]. In this section, we highlight properties of nanometer-sized labels that provide unique means for signal amplification, while overcoming limitations of traditional labels. We

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focus on optical and electrical biosensors, currently the two most common transduction approaches. A selection of biosensors using nanosized signal generation elements, such as metallic and semiconductor nanoparticles or vesicles formed from amphiphilic molecules, will be discussed.

6.1. Biosensors Utilizing Optical Reporting

Currently, biosensors relying on transduction of an optical signal are by far the most widespread [369]. Generally, optical read-out relies on generation of a colorimentric, fluorescent or chemiluminescent signal which can be visualised by eye, using e.g. CCD cameras or reflectometers as well as confocal and flatbed scanners. A large range of fluorescent (or fluorescence quenching) molecules are available for tagging of reporter elements. These include GFP (and variants such as yellow- and cyan-fluorescent proteins; YFP and CFP) [370,371], anthozoan GFP-like proteins [372-377], anthozoan non-fluorescent (quenching) chromoproteins [372], biarsenical ligands [378-380] and lanthanide probes [381,382], amongst others. Fluorescent components can be used as stand alone reporters (i.e. to report on the presence or absence of the fluorescent compound) or can be used with spectrally matched fluorescent/bioluminescent (or quenching) partners to produce resonance energy transfer (RET) assays that produce a fluorescent signal that is indicative of an interaction between labeled partners. RET assays may therefore be designed to potentially provide information about biological interactions or biosensor construction. Alternatively, label-free biosensors based on changes in optical properties (e.g. refractive index) of a thin film upon adsorption of biomolecules, have also been reported [383].

6.1.1. Metallic Nanoparticle Labels

Metallic nanoparticles exhibit unique optical and catalytic properties (see reviews [363,384]) that have seen them utilized for a variety of biological applications. A main advantage of gold colloid labels is that staining protocols involving the wet chemical deposition of metal on the


Figure 8. Optical biosensors using nanometer-sized reporter labels. (a) Scanometric detection of a DNA array using silver amplified gold colloids; (i) working principle: in the presence of tagged DNA, a sandwich complex is formed using an oligonucleotide reporter tagged with a gold colloid followed by silver amplification; (ii) visualization of the tagged nanoparticles using a commercial flatbed scanner. Reproduced with permission [388]. (b) Multiplexing of a scanometric assay using silver amplified gold colloids and Raman active dyes. (i) working principle: the spectrum of a Raman dye is used as a “barcode” to individualize different oligonucleotide reporters (i.e. multiplexing); (ii) scanometric detection of a protein microarray based on silver amplified gold. The spectrum of three different Raman dyes is used for multiplexing. Top array image: the three protein targets are present. Bottom array image: only the protein target associated to cy5 is present. Reproduced with permission [400]. (c) Test strip immunoassay for botulism toxin. (i) working principle: dye-coated liposomes carrying a receptor (GT1b) for the toxin is mixed with the sample. After migrating along the strip, liposomes carrying the toxin are immobilized in the detection zone coated with anti-BT antibodies, resulting in a colorimetric signal; (ii) dose curve response obtained from scanned images of the strips. Reproduced with permission [419].

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nanoparticles have long been available and can be used to increase the size of the colloid (and therefore the sensitivity of the assay) after biorecognition. Silver staining is most commonly used (Figure 8a). This feature makes it possible to detect low particle concentrations (to picomolar levels [385]) using the naked eye alone [385,386]. Quantification of the signal can be achieved using conventional flatbed scanners [387-389] or a CCD camera [390].

When conjugated to oligonucleotides, gold nanoparticles were also shown to confer very sharp melting profiles upon hybridised nucleotide strands [387]. This provides increased accuracy in discriminating single base mismatches, compared to standard assays performed using oligonucleotides labeled with fluorophores. Taton and coworkers [387] took advantage of this property and reported (after a silver amplification step) a femtomolar (50 fM) DNA detection limit on a microarray imaged with a conventional flatbed scanner. A commercially available device (Verigene® System by Nanosphere Inc., http: // www. nanosphere. Us / VerigeneSystem _ 4411. aspx ) uses side illumination to detect as little as 0.0025 probes/µm2 [388] (Figure 8a). This technology has been applied to a variety of biological assays [389,391-393].

Metallic nanoparticles also have the property of scattering light of a specific wavelength upon illumination with white light. Several groups have taken advantage of this phenomenon, commonly referred to as resonance light scattering (RLS), to develop new biosensing platforms. With this approach, a very intense signal is produced and no bleaching or quenching effects are observed (as for conventional fluorophores) [394]. Typical setups for the detection of RLS from nanoparticles in a microarray format are based on side illumination combined with TIRF [395-397] or on dark-field illumination [398]. With an RLS scanner, sensitivities were at least 50 times better than confocal scanners, with limits of detection in the femtomolar range (5-10 fM) being reported [396].

Multiplexing can potentially be achieved using metallic nanoparticles since the scattered wavelength depends on properties of the particles such as composition, size and shape [394,399] each of which could be varied within a single biosensor platform. Gold and silver nanoparticles were also shown to enhance the Raman scattering signal of


organic molecules; this effect is commonly referred to as Surface Enhanced Raman Scattering (SERS). Raman spectra are generated upon illumination of a sample with laser light that produces excitation of vibrational and rotational states of chemical bonds which translate as wavelength shifts. Several DNA [400] and protein [401-403] biosensor setups have taken advantage of SERS and have used specific Raman spectra of (fluorescent and non-fluorescent) Raman dyes in the vicinity of metal nanoparticles as individual barcodes for multiplexing and identification of various reporter molecules (Figure 8b). The level of specificity of resultant Raman spectra allowed parallel use of up to six distinguishable labels and a femtomolar limit of detection [401]. These studies illustrate the potential for multiplexed microarray platforms relying on visualization of metallic nanoparticles.

6.1.2. Quantum Dot Labels

Now that difficulties associated with their water solubility and functionalization are being overcome, quantum dots appear to be a promising alternative to conventional fluorescent labeling (with organic dyes) for a variety of biological applications. These semiconductor nanocrystals (with sizes usually ranging from 2 to 10 nm) exhibit unique luminescent characteristics; they were shown to be brighter and more stable against photobleaching than conventional fluorophores. Moreover, they exhibit a very broad adsorption spectrum and an emission spectrum that is narrow and size dependent (i.e. tunable) making them ideal candidates for multiplexed assays. Several groups have therefore started using quantum dot labels usually coupled to fluorescent microarray read-out systems [404-408]. Up to four different quantum dots could be detected simultaneously using one excitation light source [409]. We refer the reader to a range of general reviews on biological applications and properties of quantum dots [362,410-412].

6.1.3. Liposomes as Optical Labels

Liposomes have been investigated as labels for a variety of biosensing applications because signaling molecules (such as enzymes,

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fluorescent dyes and importantly, membrane proteins (see section 2)), can be encapsulated in their interior, bound to their surface or inserted within the bilayer (see reviews [368,413]). Signal amplification relies on the fact that a large number of signaling molecules can be associated to one binding event.

Several portable immuno- or DNA sensors based on lateral flow assays, have been developed for field applications such as rapid detection of food and waterborne cellular pathogens [414-417], toxins [418,419] or pathogenic spores [420-422]. Flow assays have been performed on membrane strips (dip-stick sensor), in microcapillaries or in microfluidic channels. For example, in a colorimetric dip-stick assay [414-416,418-422], reagents (sample and dye-loaded liposomes tagged with a reporter molecule) migrate on a membrane via capillary action until they reach the “capture zone” where a sandwich complex is formed. This results in a colorimetric signal that can be detected visually and quantified with a scanner [414,418,419,423] or hand-held reflectometer [416,420-422,424] (Figure 8c). Using the latter, nanomolar detection limits are commonly reported [424].

Fluorescence-based assays have utilized microcapillary columns [417,425-427] or microfluidic channels [415,428,429] coated with reporter molecules for target immobilization and formation of a sandwich complex with liposomes. Alternatively, complexes have been formed on the surface of a magnetic microparticle prior to immobilization in the channel using a magnet [430-432]. The signal was detected either directly using microscopy, or upon vesicle lysis and transport of the released dye to the detector situated at the end of the channel. This approach provided picomolar limits of detection (e.g. 5.5 pM for a cholera toxin assay [429]) after lysis of sensor liposomes.

6.2. Biosensors Utilising Electrochemical Reporting

In an electrochemical biosensor, a biological signal is translated into an electric signal that is used as the means of detection. Advances in micro- and nanotechnology, as well as in the semiconductor industry, have opened the way for development of novel electrochemical biosensing platforms [192]. Such platforms have been presented as a low


cost alternative to optical transduction systems for the production of small, hand-held (i.e. portable) sensors with simple read-out systems [365,433].

6.2.1. Metallic and Semiconductor Nanoparticles as Electrochemical

Reporters

Metallic and semiconductor nanoparticles have also attracted much interest in biosensor research due to their unique electrical properties (see reviews [360,365,384]). The first electrical biosensing devices used metallic particles to measure changes in electrical conductivity between two microelectrodes upon biorecognition. In the assays presented by Park et al. [434] and Velev et al., [435] the biological sensor molecules were immobilized between two microelectrodes. Upon binding of a target molecule, a sandwich complex was formed with a reporter molecule tagged with a gold colloid. The change in conductivity after a silver enhancement step could be related to the amount of target material present in solution and limits of detection down to 0.2 pM were reported [435]. Similar biosensors making use of capacitance changes have also been described [436].

Alternatively, nanoparticles can be detected by monitoring label dissolution after biorecognition and surface immobilization. This can be achieved both directly by voltammetry [433,437,438] or potentiometry [439] measurements upon particle oxidation, or indirectly by stripping voltammetry [440-449] or stripping potentiometry [450-452]. In the latter method, chemical dissolution of the label after biorecognition is followed by a concentration step involving electrodeposition of the metal ions on the electrode. This is followed by electrochemical dissolution that results in an electrical signal, which can be used for quantification (Figure 9a). A great variety of metallic and semiconductor labels have been utilized including gold and silver colloids [440,442,448,449], indium microrods [451] and nanoparticles of CdS [443,444,446] or PbS [445,447]. Using PbS nanoparticles, a detection limit of ~0.2 pM was reported. An interesting feature of semiconductor nanoparticles is that multiplexing can be achieved using the defined stripping profiles as a barcode for the identification of different labels [453-455] (Figure 9b), with one study measuring up to five targets simultaneously [453].

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Figure 9. Electrochemical biosensors using nanometer- sized reporter labels. (a) Detection of surface bound PbS nanoparticles using stripping voltammetry. Reproduced with permission [447]. (b) Assay multiplexing by use of three different semiconductor nanoparticle labels; (i) assay scheme; (ii) anodic stripping voltammetry analysis in the presence of the three lables (left) or only one label (right). Reproduced with permission [455]. (c) Amplified DNA detection using negatively charged liposomes and faradaic impedance spectroscopy. Reproduced with permission [459].

6.2.2. Liposomes as Electrochemical Reporters

Electrochemical biosensors based on the release of encapsulated redox markers from phospholipid vesicles after biorecognition have also been described. Quantification is usually achieved by amperometry [456,457] or voltammetry [458]. Liposome-based electrochemical detection was also applied in lateral flow biosensor formats [456,457].

Alternative electrochemical approaches were proposed by Willner and coworkers [366] who described a transduction mechanism whereby the biorecognition event induced an increase in electron-transfer resistance at the electrode-solution interface which could be detected


using faradaic impedance spectroscopy. In this assay format, negatively charged liposomes are used as labels for detection of a DNA target. Upon hybridization of the tagged reporter DNA, the interface between the sensor and the solution becomes highly negative and electrostatically repels positively-charged redox probes (which are present in the electrolyte) thereby increasing the electron transfer resistance (Figure 9c). Further amplification could be achieved by producing vesicle networks using biotin-streptavidin as a linker [459,460].

7. Biosensing Applications

The purpose of this section is to recognize current applications, discuss analytes of interest for currently utilized biomolecules and to give context to some of the immobilization transduction strategies discussed throughout this review. This is by no means a comprehensive review but more-so an overview of the biosensing field with some recent examples. Biosensors offer enormous potential to detect a wide range of analytes in a range of disciplines and we highlight applications in health care, the food industry, environmental monitoring, and defense/security. The primary goals when developing commercial biosensors for these applied uses are increased speed, heightened sensitivity, improved accuracy, portability and minimization of sample preparation. To facilitate adoption of next generation biosensors, these assay variables must be clearly superior to current standards (e.g. gas chromatography, mass spectrometry and high performance liquid chromatography).

7.1. Medical

As biomarker research continues to identify new protein and chemical markers associated with a given pathology, commercial interest grows in producing pre-symptomatic diagnostic tools that can detect subclinical concentrations of known markers. Medical diagnostics requires fast, accurate, inexpensive devices. The most common analytes targeted in medical biosensing include glucose, lactate, urea, creatinine, cholesterol, uric acid and DNA.

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Glucose is a common target analyte due to the prevalence of diseases characterized by altered sugar metabolism, such as types of diabetes (reviewed by Wang [461]). Glucose oxidase is commonly used as the biomolecular sensor element in glucose biosensors due to its affinity for glucose and the electrochemical signal produced by its interaction with glucose. Recently, work has become more focused on micro-type biosensors for applications in embedded glucose monitoring systems within the human body. Jia et al. recently reported on carbon nanotube-based needle-type glucose biosensors with this application in mind [462] (further discussion on nanotube based glucose sensing can be found in section 5). An alternative to the electrochemical approach for glucose biosensing is the use of optical transduction techniques (as reviewed by Pickup et al. [463]). One such approach was described by Wang et al. who reported the use of a fluorescent reporter system whereby oxygen consumed as a result of glucose oxidation caused a detectable change in fluorescence quenching [464] with the level of change being proportional to the glucose concentration of human serum samples being monitored.

Urea concentration estimation is important for monitoring kidney function and any related disorders. The enzyme urease, which catalyses the reaction of urea into ammonia and carbon dioxide is commonly used as the biorecognition element for urea sensor systems. Recent studies include those by Jha et al. who entrapped urease in polyvinyl alcohol and a polyacrilamide polymer membrane, and monitored levels of NH4

+ which is indicative of urease activity in the presence of urea [465]. Uric acid is used as an indicator of a wide range of conditions such as leukemia, pneumonia, kidney injury, hypertension and ischemia. The detection scheme for uric acid utilizes the enzymatic activity of uricase which produces a decrease in oxygen that is proportional to the concentration of uric acid. This system was recently used to determine the uric acid concentrations in human serum and urine [466].

Creatinine is another key medical analyte that is monitored to determine renal, muscular and thyroid dysfunction. Recent studies demonstrated an amperometric sensor based on creatinine amidinohydrolase and sarcosine oxidase [467] which are both used to detect the presence of this creatinine. Other analytes that have been


measured by biosensors include cholesterol [468], insulin [469], nitric oxide (e.g. detected by myoglobin [470] or peroxidase activity [471]) and cytochrome c [472] all of which have diagnostic uses.

In addition to the diagnostic applications mentioned here, biochips have potential in the drug discovery arena, a lucrative area of research that has utilized many of the existing cell-based biosensors. Biochips carrying biomolecules that are involved in certain diseases may provide a high-throughput method for drug screening of these targets (as discussed in more detail in section 2.1 of this review).

7.2. Food and Wine

Food monitoring is important for the detection/quantification of microbial content, freshness/quality, and toxic ingredients (including pesticides and allergens). Additionally, testing manufacturing processes, such as fermentation for wine and beer production are also important to ensure quality and consistency of the end product. The ability to monitor multiple analytes with increased sensitivity will lead to improved control of production processes and improved profitability. The commercial demand for fast and sensitive food analysis methods has paved the way for the applicability of biosensors in this field. A review of enzyme-based biosensors in food analysis has been published by Prodromidis and Karayannis [473] who discussed detection of glucose, fructose, sucrose, lactulose, lactose, lactic acid, malic acid, citric acid, glutamic acid, ascorbic acid (vitamin c), ethanol, and lysine, as well as the freshness indicators such as inosine which is used to monitor the progress of fermentation in wine-making since it is a fermentation byproduct that effects wine quality [474]. Additionally, inosine is an indicator of microorganism presence in food products [475].

Mycotoxins, which are toxic secondary metabolites produced by filamentous fungi, are dangerous contaminants which can be found in various food-stuffs, particularly grains. Mycotoxin-targeting biosenors based on enzyme and DNA bio-elements, have previously been reviewed by Prieto-Simon et al. [476]. More recently the same authors described an electrochemical immunosensor capable of detecting an example of

Leifert et al. 56

one such toxin, Ochratoxin A, in wine [477]. Some mycotoxins (e.g. aflatoxins) have been designated as biowarfare agents (discussed in section 7.3) due to their potential carcinogenicity [476].

7.3. Explosives and Biowarfare

Due to the nature of analytes targeted for explosives detection, a method known as “stand-off” detection is advantageous. This involves the detection of chemical vapors without necessarily seeing the explosives (e.g. if they are purposely concealed). Examples of biosensors developed for explosives detection have recently been reviewed by Smith et al. [478] TATP (triacetonetriperoxide) and TNT (2,4,6 – trinitrotoluene) are two explosives which have been investigated using immunosensors, enzymatic sensors, biologically inspired/biomimetic sensors and whole-cell biosensors (cell-based biosensing is discussed in more detail in section 2.2.2) [478]. Guan et al. also investigated the use of stochastic sensing (see section 2.2.2) to detect TNT using a genetically engineered pore-forming protein [479]. Interestingly, several mammalian ORs (see 2.1.7) have been isolated (Olfr226 from rat; Olfr2 and MOR226-1 from mouse) that are responsive to 2,4-dinitrotoluene [73] and therefore, could potentially be applied as olfactory sensors for certain explosives.

“Biologically inspired” (biomimetic) sensors utilize a more robust substitute to mimic the biological element within the system. This may be important in some field applications where conditions to maintain an active biomolecule may not be easily adapted to a hand-held device. Additionally, in an application such as landmine detection, the outcome of false negatives due to sensor protein denaturation (e.g. receptor/antibody) is not trivial [480]. Some of these biological mimics adapted to a variety of biosensing/biochip applications (not only security), include aptamers (single stranded DNA acting as receptor/protein mimics) [481] and carbon nanotubes (as ion-channel mimics) [482,483].

There are a number of biological agents that are tagged as a potential security threats. These include botulism toxin, Smallpox virus,


Hemorrhagic fever viruses (including Ebola and Yellow fever viruses), Fracisella tularensis (causing tularemia), Yesinis pestis (causing plague), and Bacillus anthracis (anthrax) [484]. Other toxins that pose a threat include ricin and diphtheria toxins. In the area of bio-defence, sensitive, early and accurate detection of agents such as these is of high importance. Huelseweh et al. have developed a simple and rapid modification to the ELISA technique that allows simultaneous detection of a number of different biowarfare agents on a protein chip using biotinylated antibody recognition and streptavidin-HRP signal amplification [485]. Investigations into detection of smallpox virions [486] and antibodies to the Ebola virus [487] were conducted using optical immunosensors as a faster, portable alternative to the ELISA assays commonly employed.

Halverson et al. [488] investigated the three proteins that comprise the active anthrax toxin, protective antigen (PA), lethal factor (LF) and edema factor (EF), and the effect of the pore made by a fragment of PA, in the presence and absence of LF and EF. The change in current through the pore was monitored as an indication of the presence of the anthrax proteins. Another study described construction of a hand-held SPR-based device capable of detecting ricin at 200 ng/ml in 10 minutes [489]. While this device is still in its prototype stage and researches admit to certain limitations of the system as it stands, it is still a step towards a generic hand-held sensing device. Additionally, the SPR technique can potentially be adapted to monitor a wide range of interactions and therefore is likely to find use in many applications.

7.4. Environmental

Detection of environmental pollutants in air, water and soil, is important for the health and well-being of humans and other biodiversity that rely on these resources. A recent review by Wanekaya et al. [490] covers many of the current biosensing developments in the environmental field. Other reviews in this area have focused on particular classes of biological detectors, such as enzymes [491] or whole cells [67], while others have focused on particular analytes such as heavy

Leifert et al. 58

metals [492-495], dioxins [496], pesticides [497], endocrine-disrupting compounds [498], or water-born pathogens [499]. These sensor technologies are moving from proof-of-concept to real-world applications where river water, wastewater, groundwater and soil have been tested for a number of the analytes mentioned above (for a review of literature involved see Rodriguez-Mozaz et al. [500]).

Some recent studies into development of biosensors targeting organophosphorous (OP) pesticides provide an example of some approaches to environmental sensing. OP pesticides have been widely utilized to control agricultural and household pests, however, they are also toxic to humans and other mammals. They constitute a range of chemical structures and exhibit a range of physicochemical properties, with their primary toxicological action arising from inhibition of the enzyme, acetylcholinesterase (AChE), which is important to nerve impulse responses [501]. It is this enzyme which is predominantly used as the sensor biomolecule for current OR biosensor designs. Recently, Istamboulie et al. [502] captured recombinant histidine-tagged AchE proteins on magnetic beads using Ni2+-histidine affinity. These beads can be attached to the surface of a working electrode for amperometric transduction, by the application of a magnetic field. An advantage of this system was the re-usability of the electrode. Another report recently described the use of screen-printed electrodes for production of an amperometric biosensor using immobilized AchE [503,504]. Vamvakaki et al. [505] presented a pH-sensitive fluorescent indicator to monitor the activity of AchE encapsulated within liposomes. The internal environment of the liposome has been reported to improve enzyme stability (for more discussion of liposomes see section 3), and liposomes can be adaptable to sol-gel matrices for biochip/array pesticide-screening applications.

Microcantilevers were also recently explored for their use in detecting OPs by Karnati et al. [506]. In addition, this group investigated an alternative enzyme (organophosphorous hydrolase, OPH) for its ability to hydrolyse OP. The detection was based on the deflection of the cantilever due to changes in enzyme conformation induced by the hydrolysis of OP by cantilever immobilized OPH. Nanotubes have also been used in conjunction with this OPH as discussed in section 5.1.


8. Conclusion

It is evident from the increasing number of biosensor-related publications and the array of associated approaches, that biosensing is seen as an important area of research worldwide. This interest is set to increase as more biologically active proteins are characterized and the first commercial biosensors are developed. Currently, molecular biosensing is still a relatively new field of research with few commercial examples. In this review, we have discussed some of the key classes of biological sensors being studied (membrane proteins, enzymes, whole cells and virions) and their current assay technologies, the use of lipid supports, nanopatterning approaches, use of carbon nanotubes and current biosensing applications and their target analytes.

Currently, the main factors limiting the wide-spread use of biosensors include (1) the need to overcome issues of functional integrity of sensor proteins/cells under harsh purification or assay conditions, (2) limited ability to correctly and consistently orientate sensor proteins, (3) limitations to use of lipid supports, particularly with regard to targeted insertion of membrane proteins and even coverage of biochips, (4) need for metabolic relevancy of detected events for medical applications, (5) lack of portability and (6) expense. Many of these limiting factors currently drive aspects of biosensor design, often restricting the approach that can be used for a given application, however, advances in biosensor-related technologies should see design flexibility improve.

There are some clear trends emerging and the next decade of biosensor research is likely to be characterized, for example, by increasing use of cell-free approaches. Cell-based approaches, which are in a sense traditional, are often utilized and may lend metabolic relevancy to a sensing event. To this end, it is likely that culturing surfaces will be developed which can be applied to a potential biosensor surface (e.g. chips and electrodes) and facilitate cell growth, proliferation and differentiation directly on targeted areas of the surface. However, cellular approaches are inherently limited by the need to maintain cells and their fragility under assay conditions, particularly in the field. In addition, cell-based sensors are largely confined to medical and other screening applications.

Leifert et al. 60

Other trends are likely to include miniaturization of sensing elements, increased use of biochips containing sensor arrays, multiplexing of sensing and transduction systems, advances in the use of lipid supports and membrane protein sensors, increased use of optical biosensors that utilize a range of novel fluorescent reporters, increased use of microfluidics and biosensing of volatiles. It is also likely that good progress will be made in attempts to increase efficiency of electron transfer for electrochemical approaches. In combination, these advances will see commercially produced biosensors become increasingly commonplace and the range of biosensing applications expand rapidly.

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