Biomacromolecular nanostructures-based interfacial engineering: from precise assembly to precision biosensing

Biomacromolecular nanostructures-based interfacial engineering: from precise assembly to... Abstract Biosensors are a type of important biodevice that integrate biological recognition elements, such as enzyme, antibody and DNA, and physical or chemical transducers, which have revolutionized clinical diagnosis especially under the context of point-of-care tests. Since the performance of a biosensor depends largely on the bio–solid interface, design and engineering of the interface play a pivotal role in developing quality biosensors. Along this line, a number of strategies have been developed to improve the homogeneity of the interface or the precision in regulating the interactions between biomolecules and the interface. Especially, intense efforts have been devoted to controlling the surface chemistry, orientation of immobilization, molecular conformation and packing density of surface-confined biomolecular probes (proteins and nucleic acids). By finely tuning these surface properties, through either gene manipulation or self-assembly, one may reduce the heterogeneity of self-assembled monolayers, increase the accessibility of target molecules and decrease the binding energy barrier to realize high sensitivity and specificity. In this review, we summarize recent progress in interfacial engineering of biosensors with particular focus on the use of protein and DNA nanostructures. These biomacromolecular nanostructures with atomistic precision lead to highly regulated interfacial assemblies at the nanoscale. We further describe the potential use of the high-performance biosensors for precision diagnostics. biosensor, interface engineering, homogeneity, orientation control, fused proteins, DNA nanostructures INTRODUCTION The biosensing interface between the solid–liquid phases plays an important role for mass transfer, electron mobility, energy exchange and signal transduction. By tailoring the interface with functional molecules or materials, surface properties can be effectively modulated to ensure the homogeneity of the biosensor surface and thus to improve biomolecular recognition, such as nucleic acid hybridization or antigen-antibody binding for efficient optical or electronic signal transduction [1–5]. Given that this transduction process generally occurs at the solid–liquid interface, in-depth understanding of the interactions between biomolecules and the supporting interface is indispensable. For example, (i) what is the main difference between surface-confined and solution-phase biomolecular recognition [6]; (ii) how does interfacial architecture affect biomolecular adsorption, assembly, folding and diffusion [7,8]; and (iii) how do surface properties influence the folding free energy and structural dynamics of biomolecules [9,10]? Along the lines to address these questions, a number of strategies have been proposed to develop high-performance biosensing interfaces with improved biomolecular recognition efficiency and reduced non-specific adsorption [11]. Great efforts have been made to increase the homogeneity and orderliness of interfacial assembly of enzyme proteins, antibodies or DNA probes [12]. Especially, rational interfacial design has proven to effectively realize the orientation control of the biorecognition molecules or reduce non-specific adsorption and avoid false positives. Among these approaches, biomimetic interfacial engineering has shown great promise for in-vitro biodetection [13]. In nature, biorecognition processes are generally highly efficient and specific, which are often based on synergic functions of ‘molecular machinaries’ that comprise multiple biomolecules in molecule-crowding intracellular settings. Construction of biomacromolecular nanostructures represents a promising means to mimic their intracellular counterpart and engineer biosensing interfaces in a programmable way [14]. In this review, we aim to summarize recent progress in using biomacromolecule nanostructures for interfacial engineering of biosensors primarily on gold surfaces and with particular focus on structured protein and DNA probes. Biosensors based on other surfaces such as graphene oxide, metal oxides and 2D transition-metal dichalogenides nanomaterials have also been explored and could be found in several excellent reviews [15–18]. We also provide perspectives on their applications in precision diagnosis. PROTEIN-BASED INTERFACIAL ENGINEERING Bioactivities of protein molecules (i.e. antibodies and enzymes) are responsible for high efficient antigen recognition or direct electron transfer at the sensor interface. Nevertheless, it is challenging to retain their natural bioactivities and conformations when proteins are immobilized on the sensing interface via physical adsorption or covalent conjugation [19]. Since proteins can interact with the interface via their various surface groups, they tend to adsorb on the interface with random orientations that often do not favor their bioactivities [20]. Moreover, surface attachment might also alter the native conformation of proteins [21], thereby degrading their functions. Hence, a range of static and dynamic interfacial engineering approaches have been proposed to assemble proteins on the interface with better orientation and freedom for conformational change and functional modulation. STATIC ENGINEERING BY CONTROLLED IMMOBILIZAITON OF PROTEINS Orientation control of proteins at the interface is highly required to retain their bioactivities. However, proteins often adopt random orientations at the biosensing interface due to their multi-site interactions with the substrate. For example, Immunoglobulin Gs (IgGs) (Fig. 1a) consist of one constant fragment (Fc) and two variable antigen-binding fragments (Fab), resulting in random orientations of interfacial proteins: head-on (both Fab on surface), tail-on (Fc on surface), side-on (Fc and one Fab on surface) or flat-on (all fragments on surface) (Fig. 1b). To achieve optimal antigen-binding activity, antibodies should be immobilized with favorable orientation of tail-on (Fig. 1c). Efforts to address this problem have been focused on precisely defined protein adorsption (non-covalent) and anchoring strategies (covalent) [19]. Figure 1. View largeDownload slide Structure and interfacial orientation of natural antibodies. (a) Schematic depiction of the structure of natural antibody with two Fab and one Fc domain, as well as multiple reactive groups, such as amine, sugar and disulfide group. Antibody can be immobilized in either a random (b) with four typical orientations or oriented (c) fashion. Figure 1. View largeDownload slide Structure and interfacial orientation of natural antibodies. (a) Schematic depiction of the structure of natural antibody with two Fab and one Fc domain, as well as multiple reactive groups, such as amine, sugar and disulfide group. Antibody can be immobilized in either a random (b) with four typical orientations or oriented (c) fashion. Intermediate proteins (e.g. protein A or protein G) (Fig. 2a) with multiple binding domains specific to the Fc part of the antibody were employed to immobilize the antibody with favorable tail-on orientation. As compared to their randomly immobilized counterparts, the oriented antibodies improved the biosensing sensitivity and specificity [22,23]. However, even partially oriented intermediate proteins may induce random orientations of antibodies, which still poses a limit. To further improve the protein orientation, a highly organized IgG-protein with 3D structures and exposed antigen-binding sites was designed [24]. However, this oriented immobilization strategy is heavily dependent on solution pH that negatively affects the protein G orientation on gold and subsequent antibody binding [25]. Cysteine-functionalized protein G (Cys-protein G) has proven successful for controlled orientation of protein A or protein G at the interface [26]. A recombinant Cys-protein G trimer synthesized by repeated linking of protein G via flexible linkers can enhance the bioactivities of antibodies immobilized on magnetic silica nanoparticles. Compared to a Cys-protein G monomer probe, the trimer improved the sensitivity by 10-fold. By genetically fusing protein A domain to a Cys-exposing variant of Escherichia coli protein ompA, favorable binding orientation and antibody presentation were realized [27]. Similarly, a ‘super-oriented IgG’ was constructed to attach to the oriented protein A at the interface via enzyme conjugation, leading to enhanced affinity up to ∼100 fold over the partially oriented IgG physical immobilized at interface [28]. Figure 2. View largeDownload slide Protein-based interface modulation. (a) Protein A/G mediated non-covalent orientation of antibodies. (b) Covalent orientation of antibodies by either reactive groups distributed on the surface of antibody or fusion protein technology. Adapted from [34]. (c) Protein-DNA conjugates mediated orientation of antibodies. (d) Proteins reversibly modulate the sensing interface. Adapted from [36]. (e) MBP dynamically modulates the electrochemical interface. Adapted from [40]. Figure 2. View largeDownload slide Protein-based interface modulation. (a) Protein A/G mediated non-covalent orientation of antibodies. (b) Covalent orientation of antibodies by either reactive groups distributed on the surface of antibody or fusion protein technology. Adapted from [34]. (c) Protein-DNA conjugates mediated orientation of antibodies. (d) Proteins reversibly modulate the sensing interface. Adapted from [36]. (e) MBP dynamically modulates the electrochemical interface. Adapted from [40]. Although the interfacial orientation control with intermediate proteins can enhance biosensing performance, this non-covalent interaction between intermediate protein and interface is generally weak and sensitive to environmental changes, such as pH, ionic strength and temperature [25]. In response, direct covalent immobilization of proteins provides a feasible way to engineer the biosensing interface. Classic reactions of inherent reactive groups and tailor-made modifications on protein molecules have been used to improve the orientation and homogeneity of proteins at the interface (Fig. 2b). Amine (NH2) groups in the lysine side-chain on the antibody surface are usually used for relatively random covalent immobilization, although it generally lacks orientation control [29,30]. To improve the orientation control, in contrast, the unique carbohydrate moiety at the Fc portion of an antibody is more amenable to orient proteins via covalent immobilization. In the presence of periodate sodium [31] or boronic acid [32], the carbohydrate vicinal hydroxyl groups would be oxidized to aldehydes and cyclic boronate esters, respectively, which provide oriented covalent antibody immobilization (Fig. 2b). Analogously, the disulfide group in the hinge region of antibody was also exploited for oriented immobilization of antibody via reduction into thiol groups (Fig. 2b) [33]. These methods of covalent coupling of antibody at the interface has proven widely useful for antibodies with multiple activatable sites on antibodies. In addition, DNA can also be used to immobilize antibody, serving as both spacer and binding domain. Heath et al. [34] established a DNA-encoded antibody library for spatially multiplexed detection of nucleic acids, proteins and cells (Fig. 2c). Importantly, the hybridized duplex probe is rigid enough to maintain a favorable orientation (upright) of the antibody at the interface [35]. More importantly, this DNA-antibody conjugate renders the biosensing interface regenerable via facile water rinsing and robust to resist the denaturation induced by surface effect. Scaffolded protein nanostructures provide a powerful means to oriented assemble proteins at the interface. For example, Zhang et al. fused a coding sequence of streptavidin recognition peptide (streptag) to a specific site (3΄ end) of the phoA gene that codes for E. coli alkaline phosphatase (EAP) [1]. This fused protein was used for oriented immobilization of EAP enzyme on microtiter plates via streptag–streptavidin binding. With the insertion of a flexible linker peptide coding sequence between the streptag and EAP sequences, the enzyme activity of immobilized EAP increased by ∼8.4-fold over the fused protein without linker sequence. The linker peptide provides a spacer that minimizes the steric hindrance between EAP and streptavidin, enhancing the orientation effect. In a further step, by incorporating a linker peptide with a cysteine residue at the C-terminal of glucose oxidase (GOx), they engineered a fusion structure (GOx-linker-cysteine) that can be immobilized on gold surfaces with Au–S bond or on a silanized glass surface via disulfide bond (Fig. 2b) [3]. Importantly, this fused cysteine enables the GOx to be immobilized at the interface with well-controlled orientation, thus forming a homogeneous biosensing interface. With the synergistic effect of the linker spacer and interfacial homogeneity, a higher and more stable electrochemical current response was obtained as compared to the GOx without fused cysteine and linker spacer. Although the covalent orientation of proteins is effective, the high-affinity interaction (e.g. biotin–streptavidin binding) is usually irreversible, which makes the biotin–streptavidin based assembly non-regenerable. To realize a regenerable biosensing interface (e.g. surface plasmon resonance, SPR), two streptavidin affinity tags, nano-tag and streptavidin-binding peptide (SBP-tag), were employed (Fig. 2d) [36]. Both of them can specifically interact with streptavidin while the binding affinity (KD ∼4-17 nM) is weaker compared to the biotin–streptavidin binding (KD ∼1 fM), thus enabling the easily controlled association and dissociation. With this tunable binding mechanism, Zhang and coworkers developed a SPR biosensor chip coated with streptavidin for reversible, site-directed protein immobilization mediated by the nano-tag [36]. The streptavidin surface could be regenerated repeatedly without loss of activity even by injection of 50 mM of NaOH solution. This reversible biosensing interface could be readily generalized to build other SPR biosensors, permitting anchoring of various proteins on the streptavidin surface in a stable, site-directed and reversible fashion. DYNAMIC ENGINEERING BY CONFORMATIONAL CHANGE OF PROTEINS In nature, biomolecular recognition is generally accompanied by conformational changes of proteins, such as ligand-induced protein folding or unfolding. Such a special conformation-switchable mechanism can be used to engineer the biosensing interface in a dynamic manner [37]. It can also be extended to design a biosensor exclusively based on the conformational switch of proteins [38]. Nevertheless, most proteins, unlike single-stranded (ss)DNA, do not undergo significant conformational changes upon ligand binding. Hence, a highly sensitive method that can precisely probe the weak conformational alterations is desirable. Extensive studies have shown that electrochemical sensors are capable of tracing the relatively small conformational changes of a surface-confined, redox-labeled macromolecule. Since heterogeneous electron transfer between an electrode and a surface-confined redox molecule exhibits an exponential dependence on both distance and the Marcus coupling factor [39], even small conformational changes in proteins can produce large variations in electron-transfer rates, which in turn translate into measurable changes in electrochemical signals. For example, Benson et al. [40] exploited the ligand-induced hinge-bending motions in the electrode surface-confined maltose binding protein (MBP) that they used to engineer the biosensing interface in a dynamic fashion (Fig. 2e). In this report, a gold electrode was first assembled with a functional monolayer, serving as a binding interface for site-specific immobilization of proteins [40]. Subsequently, the MBP protein was immobilized on the electrode surface with a specific orientation such that a redox reporter group (ruthenium complex) was fixed on the electrode for electrochemical readout. As the target ligand of maltose binds, the target-induced hinge-bending motion in the protein proceeds, which moves the Ru(II) reporter away from the electrode. This target-responsive dynamic modulator in turn triggers a concentration-dependent decrease in the observed electrochemical response, which provides the electronic detection of maltose at mM concentrations. It should be noted that the relatively low sensitivity (at the mM level) of this dynamic sensor is limited by the low affinity of MBP against maltose (KD = 4 mM) but does not pose a fundamental restriction of this strategy. Apparently, this method can be generalized to detection of other related natural or designed MBP-like proteins that would undergo a similar ligand-induced conformational change [41]. Allosteric protein switches are ubiquitous in the biological signal transduction system, which enable cells to sense and respond to specific molecular events. Inspired by nature, tailor-engineering of protein switches with custom input and output functions is significant in molecular diagnostics and physiological decoding [42]. To demonstrate this concept, RG13, a fusion protein of MBP and TEM1β-lactamase, was engineered as a model protein switch with electrochemically activated switching behavior by reducing the disulfide bonds in the switches [43]. In this approach, an electrochemical signal can be used as an exogenous input to control the on/off state of protein switches by modulating the oxidation state of an introduced disulfide bond on the electrode surface. The presence of maltose is the key to activating the enzyme activity due to the induced large hinge-bending conformational change in the MBP domain of RG13. This strategy allows the allosteric protein switch to dynamically regulate the interface signaling. Although the protein switch or folding is effective to engineer the biosensing interface, the ligand-induced conformational change in protein switches is relatively small. In contrast, a more flexible peptide, like aptamer, can trigger a large structure change by binding to its corresponding antibody. Lai and coworkers [44] reported an electrochemical peptide-based dynamic biosensor for the sensitive and specific detection of HIV anti-p24 antibodies where a highly antigenic epitope from the HIV-1 capsid protein, p24, was used as the recognition peptide by modifying with MB (methylene blue) and thiol at two ends, respectively. Of note, this epitope is a short linear peptide lacking defined secondary structure, and adopts predominantly an α-helical conformation in native state. Thus, the binding of the target antibody facilitates the formation of a more rigid complex that induces a large conformation change. This dynamic peptide-based biosensor achieved a detection limit of 10 nM, and had a dynamic range that is broader than the typical concentration range commonly observed with HIV-infected patients. DNA-BASED INTERFACIAL ENGINEERING Single-stranded (ss-)DNA can hybridize to its complementary DNA strictly obeying the Watson-Crick base-pairing rules, which allows DNA recognition and structural assembly in a highly predictable and finely programmable manner. These advantages open opportunities for biosensing interfacial engineering. It was reported that the large inter-probe distances and upright orientation of surface-tethered DNA probes is imperative to realize efficient hybridization. However, engineering an upright and accessible DNA recognition layer is challenging because of the unexpected surface adsorptions, disordered conformations and inhomogeneity of grafting density of DNA bioprobes at the biosensing interface. Static and dynamic interface engineering using DNA and DNA structures have been developed to overcome this challenge. STATIC ENGINEERING BY CONTROLLED IMMOBILIZATION OF DNA PROBES In a typical ssDNA probe-based DNA biosensor, an efficient probe–target hybridization process is fundamental for improving the biosensing performance, particularly in sensitivity and specificity. Nevertheless, high hybridization efficiency depends on a favorable orientation (upright) of DNA probes and rational inter-probe distance (Fig. 3a) [45]. It is expected that the ssDNA probes can adopt an upright orientation at the biosensing interface via a single-point attachment. By taking thiolated DNA (SH-DNA) as an example, however, significant interactions exist between the DNA bases and the gold surface via multiple nitrogen atoms [46], which allows ssDNA molecules to lie down on the Au surface, resulting in a largely limited accessibility of the target sequences with reduced hybridization efficiency (Fig. 3b). Such non-specific adsorption onto the Au via Au–N interactions has been confirmed by Tarlov and coworkers [47]. Further characterization, such as X-ray photoelectron spectroscopy and Fourier transform infrared spectroscopy, revealed that the nonspecifically adsorbed ssDNA could not be removed, even with extensive rinsing or heating to 75°C, and the ‘specifically’ anchored ssDNA monolayer is not oriented perpendicularly to the surface, especially at low densities. Short DNA probes tended to orient parallel to the surface, whereas the relative long strands preferred to form disordered film due probably to adjacent entanglement [48]. Although the densely packed short DNA probes may assume an upright conformation on the surface, the restricted target accessibility would offset this advantage of favorable probe orientation (Fig. 3c). Figure 3. View largeDownload slide ssDNA-based interface modulation. (a) Schematic depiction of an ideal DNA recognition layer with a large intermolecular distance and linear upright orientation. (b) Sparse probes are prone to lying flat-on the surface due to non-specific adsorption. (c) Densely packed short DNA probes adopt an upright conformation, but yield poor hybridization efficiency because of reduced accessibility. (d) Diluents, such as MCH and OEG, co-assembled with ssDNA to favor the upright orientation on the surface and reduce non-specific adsorptions. Figure 3. View largeDownload slide ssDNA-based interface modulation. (a) Schematic depiction of an ideal DNA recognition layer with a large intermolecular distance and linear upright orientation. (b) Sparse probes are prone to lying flat-on the surface due to non-specific adsorption. (c) Densely packed short DNA probes adopt an upright conformation, but yield poor hybridization efficiency because of reduced accessibility. (d) Diluents, such as MCH and OEG, co-assembled with ssDNA to favor the upright orientation on the surface and reduce non-specific adsorptions. To help address this dilemma, Tarlov et al. [49] introduced a coassembling small molecule (mercaptohexanol, MCH), synergizing with SH-DNA to engineer the recognition ability of bioprobes at the interface (Fig. 3d). This ‘helper’ molecule is able to largely remove the nonspecifically adsorbed DNA and meanwhile protrudes the surface-attached DNA probes into solution phase via the repulsion between the net negative dipole of alcoholic terminus and the negatively charged DNA backbones. These mixed ssDNA/alkylthiol monolayers have been extensively investigated by varied surface-measuring techniques and confirmed with a favorable upright orientation of ssDNA bioprobes [50]. This synergic static modulation strategy was widely employed in engineering DNA biosensors and biochips [51–53]. However, the co-assembled small molecule diluents (MCH) cannot resist the non-specific protein adsorption effectively. In response, a ssDNA/oligo-ethylene glycol (OEG) mixed monolayer was used to improve the protein resistance of the biosensing interface, particularly in complex matrices, such as blood [54,55]. We also note that relatively rigid double-stranded (ds) DNA molecules represent another route to improving interfacial probe arrangement with enhanced sensitivity and specificity, especially for toehold design [56]. Although helix structures or the co-assembled diluents can aid DNA probes to adopt a favorable orientation at the interface to some extent, the probe density is still a critical factor controlling the kinetics of target/probe hybridization [57]. Interestingly, a simple optimization of the assembly concentration of DNA probes can accurately modulate the interfacial probe average density. One concern in this approach is that local lateral interactions inevitably exist in DNA films, particularly for long sequences [58,59]. This enables the prediction of the most favorable hybridization in the ‘Langmuir’ (L) regime hard to reach, because it only exists in the limits of sparse films where probes are so far apart that they do not interact with each other [60]. From a single-molecule view [61], the aggregation patches on the Au surface can significantly reduce target accessibility, which, nevertheless, could not be easily eliminated using MCH as the dilution molecule. To better address this challenge, these empirically ‘static’ modulation methods dependent on the scheduled probe/diluent ratio may need to couple with nanostructured surfaces or conceptually new probe-design strategies. DNA nanostructures can be used to construct a scaffolded biosensing interface to tune the sensitivity of biosensors in a programmable fashion [14,62]. So far, researchers have engineered a variety of DNA nanostructures with well-defined dimension, topography and precisely controlled functions by using DNA nanotechnology. These functional DNA nanostructures have been actively exploited to develop in-vivo or in-vitro biocomputing and biosensing devices [63–65]. Because the sensitivity of biomolecular detection depends not only on the affinity between biomolecules, but also on the interfacial properties of the biosensors [66], the size reduction of biosensors, particularly to the nanoscale, usually accelerates the mass transport rate and improves the sensitivity [67,68]. However, the limited space available in nanosensors restricts the effective probe numbers and biorecognition events. To address this challenge of size reduction, a trans-scale biosensor that incorporates nano-architectures into macroscopic surfaces is necessary [5]. Nevertheless, reproducible engineering of nanostructured surfaces with well-defined topography remains technically difficult, though high-cost photolithography potentially offers a route to the fabrication of nanostructures at the wafer-scale [69]. Recently, Fan et al. [14] developed a conceptually new ‘soft lithographic’ strategy to reproducibly engineer and programmably modulate a biosensing interface using well-defined 3D DNA nanostructures (Fig. 4a). By patterning the macroscopic gold electrode with tetrahedral DNA nanostructures (TDNs) varying in sizes, the detection limit of DNA sensors can be programmably tuned over four orders of magnitude. Figure 4. View largeDownload slide DNA nanostructure-based programmable modulation on the sensing interface. (a) Assembling DNA nanostructures with varying sizes for programmable modulation on the sensing interface [14]. (b) A universal biosensing platform based on tetrahedron-structured DNA probes (TSPs) [70]. Reprinted with permission from [14,70]. Copyright 2015 and 2010 by WILEY-VCH, Weinheim, respectively. (c) TSPs-based E-DNA sensor for microRNA detection [85]. (d) TSP-conjugated antibody for sensitive PSA detection amplified with HRP-AuNP [79]. (e) TSP-conjugated aptamer for sensitive exosome detection [82]. (f) TSP-conjugated aptamer-HCR for sensitive detection of cancer cells [83]. (g) TSP-mediated capillary microarray for multiplexed bioassays, achieving detection limits of 1 μM and 0.1 nM for small molecules (ATP and cocaine), respectively [93]. Reprinted with permission from [79,82,83,93]. Copyright 2014 and 2017 by the American Chemical Society. Figure 4. View largeDownload slide DNA nanostructure-based programmable modulation on the sensing interface. (a) Assembling DNA nanostructures with varying sizes for programmable modulation on the sensing interface [14]. (b) A universal biosensing platform based on tetrahedron-structured DNA probes (TSPs) [70]. Reprinted with permission from [14,70]. Copyright 2015 and 2010 by WILEY-VCH, Weinheim, respectively. (c) TSPs-based E-DNA sensor for microRNA detection [85]. (d) TSP-conjugated antibody for sensitive PSA detection amplified with HRP-AuNP [79]. (e) TSP-conjugated aptamer for sensitive exosome detection [82]. (f) TSP-conjugated aptamer-HCR for sensitive detection of cancer cells [83]. (g) TSP-mediated capillary microarray for multiplexed bioassays, achieving detection limits of 1 μM and 0.1 nM for small molecules (ATP and cocaine), respectively [93]. Reprinted with permission from [79,82,83,93]. Copyright 2014 and 2017 by the American Chemical Society. A typical DNA tetrahedron-structured probe (TSP) was self-assembled with a relatively long, probe-bearing oligonucleotide and three thiolated oligonucleotides, which carries a pendant DNA probe (hybridizable domain) at one vertex and three thiol groups (-SH) at the other three vertices (Fig. 4b), allowing it to firmly anchor on the gold surface via Au–S bonds [70]. Of note, this 3D-structured probe demonstrates several advantages over its single-stranded or duplex counterparts in the engineering of biosensing interfaces. First, the high mechanical rigidity of TSP allows it to adopt a highly ordered, upright orientation at the Au surface, which was stabilized by the three thiol ‘legs’ that greatly increase the stability of the surface-confined probes by ∼5000 times as compared to the mono-thiolated DNA strands [71]. Second, the TSP structures avoid inter-probe entanglement (in ssDNA probes) via spatially segregated pendant probes, and also reduce the surface effects by positioning the probes in solution-phase-like settings [72]. Importantly, it is able to precisely anchor a single DNA probe on a single TDN, which is hard to achieve with inorganic nanostructures. Third, the TSP-decorated surfaces are protein-resistant, and the ‘helper’ molecule, MCH, is not necessary in this case, both of which allow the TSP-based sensors to be deployed directly in complex matrices, such as serum, for practical applications. In addition, diffusion and convection at the nanostructured interfaces are also expected to be higher than those at macroscopic ones [68]. In TSP-based sensors, the lateral distance between probes is dominated by the TDN scaffold, and thus this distance can be precisely tuned at the nanometer scale by varying the size of TDNs [73]. To this end, five types of TDNs (TDN-7, TDN-13, TDN-17, TDN-26 and TDN-37) with different sizes have been rationally designed [14]. The numbers (7, 13, 17, 26 and 37) designate the base pairs of each edge in the TDNs. Because each base pair was separated by 0.34 nm in a double helix, the edge lengths of the TDNs were able to be calculated with precision, equal to 2.4, 4.4, 5.8, 8.8 and 12.6 nm, respectively (Fig. 4a). It was observed that the electrode surface density of TDN probes was inversely proportional to the size of the TDN, which was further confirmed by statistical analysis that revealed the nearly linear relationship between the lateral distance and TDN size. Such an ability to precisely regulate the lateral distance can also tune the surface hybridization capabilities, including hybridization kinetics and efficiency, which depends heavily on the probe distance. By regulating the probe distance via the deployment of TDNs with varied size, the detection limit of the biosensor was programmed [14]. This result is consistent with that reported by the Howorka group [74], in which the molecular recognition of receptors on TDN-scaffolded surface was improved by an order of magnitude. With the use of TDNs, a variety of ultrasensitive biodetections, such as nucleic acids [70,75–78], proteins [79–81], exosomes [82] and cells [83], have been conducted. By using TSP-patterned gold electrodes, Pei et al. [70] engineered a biosensor with a typical sandwich hybridization strategy, in which the upright probe can selectively capture its complementary sequence with high discrimination ability towards non-cognate sequences (Fig. 4b). The free domain was subsequently flanked by the biotinylated reporter probe that induced the production of electrochemical signals via the substrate catalysis of horseradish peroxidase (HRP). In combination with hybridization chain reaction amplification, this tetrahedral platform improved the detection limit to 100 aM [76]. For microRNA detection, a synergic TSP-based electrochemical sensor was constructed by integrating with rolling circle amplification (RCA) and silver nanoparticles that were attached to the RCA products in tandem [84]. This TSP-RCA sensor achieved a limit of detection as low as 50 aM. To further improve the sensitivity, Wen et al. [85] increased the inter-probe distance using interfacial engineering and multi-enzyme amplification, which yields an extremely low detection limit of 10 aM (∼600 microRNA molecules in 100 μL) (Fig. 4c). In addition to direct use as a capture probe, the TDN can also serve as a probe scaffold to build immuno-sensing layers by conjugating with antibodies. For example, antibodies of tumor-necrosis-factor alpha (TNF-a) can be easily and reversibly anchored onto DNA nanostructured surfaces via a DNA-bridged antibody linking with EDC-NHS chemistry, which converts the DNA recognition layer into a protein recognition one [80]. Next, Zuo et al. [79] reported another antibody immobilization strategy that uses highly efficient click chemistry instead of a carbodiimide reaction to conjugate the PSA antibody with a DNA-bridged TDN scaffold (Fig. 4d). By precisely controlling the nanospacing of anchored antibodies with a nanostructured scaffold, they achieved an extremely high sensitivity (1 pg/mL) in PSA detection. Instead, a more direct, DNA-bridge-free approach was introduced to anchor anti-IgG on the TDN scaffold for ultrasensitive IgG and bacteria quantification [81,86,87]. By replacing the pendant probe with an expanded nucleotide-containing aptamer, Tan et al. [82] developed a TSP-assisted aptasensor for direct capture and detection of hepatocellular exosomes (Fig. 4e). This TSP-assisted aptasensor demonstrated improved accessibility and detected the exosome with 100-fold higher sensitivity as compared to the ssDNA-functionalized aptasensor. In combination with a multibranched hybridization chain reaction that offers multiple biotins and branched arms, this TSP-functionalized interface has been engineered to realize multivalent capture and highly sensitive detection of cancer cells (Fig. 4f) [83]. The TSP-based biosensing interface has also proven to be an effective platform for the detection of small molecules, ions and even multiplex targets, because the suspended probe can be easily replaced with various functional nucleic acids. As a proof of concept, a split aptamer for a small molecule, cocaine, was incorporated into TDN to develop a performance-enhanced electrochemical sensor for cocaine [88]. To realize higher sensitivity, another electrochemical cocaine sensor was presented based on structure conversion from a frustum pyramid to an equilateral triangle [89]. In this approach, the presence of cocaine triggered the aptamer-bearing DNA nanostructure change from ‘Close’ to ‘Open’, achieving a detection limit of 0.21 nM. Likewise, an anti-ATP aptamer was immobilized at the top of the tetrahedron, thereby forming an ATP sensor that could detect a concentration as low as 0.2 nM [90]. In conjunction with mercury-specific oligonucleotides, Yin et al. [91] constructed a turn-on Hg2+ sensor with a detection limit of 100 pM, which was 100-fold lower than that obtained with a ssDNA-based sensor. As a versatile probe scaffold, TSP can easily evolve into a microarray platform by replacing the pendant probe with the corresponding target-responsive probes. Recently, Li and coworkers fabricated a high-throughput addressable microarray by covalently coupling TSP onto glass substrates for multiplexed analysis with improved sensitivity and specificity [92]. The detection limits of this sensor array for miRNA (let-7a), PSA and small molecule (cocaine) were 10 fM, 40 pg/mL and 100 nM, respectively. To accelerate the target binding, Qu et al. [93] reported a rapid (<5 min) arrayed DNA-nanostructure-supported aptamer pull-down (DNaPull) assay under convective flux in a glass capillary (Fig. 4g). This DNaPull assay allowed multiplexed analysis of the contents in droplets with nano- or picoliter volumes, achieving detection limits of 1 μM, and 0.1 nM for small molecules (ATP and cocaine) and a biomarker (thrombin), respectively. DYNAMIC ENGINEERING BY CONFORMATIONAL CHANGE OF DNA Dynamic engineering is probably more attractive due mainly to the target-induced conformational change. These conformation-switchable DNA molecules, such as hairpin (quadruplex and pseudoknot with intramolecular secondary structures), aptamer and DNAzyme, possess relatively rigid and ordered structures that may, in principle, prevent inter-strand entanglement at the interface and/or increase the orderliness of DNA bioprobes. In 2003, Heeger and coworkers [94,95] developed a conceptually new electronic equivalent of the molecular beacon, the electrochemical DNA (E-DNA) sensor, which exploits target-induced conformational changes of hairpin (stem-loop) probes on Au electrode surfaces (Fig. 5a). This surface-confined ‘dynamic’ probe was labeled with an electroactive ferrocene (Fc) at the 5΄ end and a thiol at the 3΄ terminal. Initially, the hairpin localizes Fc proximal to the electrode surface, thus enabling a rapid electron transfer. After hybridization, the stem opened and extended to a rigid duplex, forcing Fc away from the electrode surface. This dynamic response allows separation between Fc and the electrode surface reaching several nanometers that yields a large and measurable change in electrochemical signal because of the exponential decay of electron transfer with distance at the interface. This sensor achieved a selective detection of 10 pM target DNA featuring in single-step and electronic (electrochemical) detection. More importantly, this design utilizes only the intrinsic conformational change of surface-confined DNA probes to modulate interface signaling, avoiding the introduction of exogenous reagents. By using such a dynamic, one-step response strategy, Lubin et al. [96] successfully detected DNA authentication tags that are associated with paper or drugs. Later, Lai and coworkers [97] demonstrated a sequence-specific detection of unpurified amplification products of the gyrB gene of Salmonella typhimurium, signifying a promising way towards rapid, sample-to-answer pathogen detection. Figure 5. View largeDownload slide DNA-based dynamic interface modulation. (a) The stem-loop-structured probe design for a reagentless electrochemical DNA (E-DNA) sensor [94]. (b) A signal-on stem-loop-structured E-DNA sensor with HRP-catalyzed signal amplification [98]. (c) Stand displacement reaction-based signal-on electronic detection [101]. Reprinted with permission from [94,101]. Copyright 2003 and 2006 by National Academy of Sciences, USA. (d) Two-stem-loop-structured E-DNA sensor [102]. (e) Three-stem-loop-structured E-DNA sensor [103]. (f) A conformation-switchable triplex E-DNA sensor [104]. (g) DNAzyme-based dynamic probe as an ion (Pb2+) sensor [105]. (h) Dynamic DNA aptamer as an ATP sensor [107]. Reprinted with permission from [98,102–105,107]. Copyright 2007, 2008, 2009 and 2014 by the American Chemical Society. Figure 5. View largeDownload slide DNA-based dynamic interface modulation. (a) The stem-loop-structured probe design for a reagentless electrochemical DNA (E-DNA) sensor [94]. (b) A signal-on stem-loop-structured E-DNA sensor with HRP-catalyzed signal amplification [98]. (c) Stand displacement reaction-based signal-on electronic detection [101]. Reprinted with permission from [94,101]. Copyright 2003 and 2006 by National Academy of Sciences, USA. (d) Two-stem-loop-structured E-DNA sensor [102]. (e) Three-stem-loop-structured E-DNA sensor [103]. (f) A conformation-switchable triplex E-DNA sensor [104]. (g) DNAzyme-based dynamic probe as an ion (Pb2+) sensor [105]. (h) Dynamic DNA aptamer as an ATP sensor [107]. Reprinted with permission from [98,102–105,107]. Copyright 2007, 2008, 2009 and 2014 by the American Chemical Society. The hairpin structure of DNA bioprobes at the interface can be employed to sterically hinder the accessibility of enzymes to bioactive sites at one end of the hairpin DNA, and thus a target-responsive ‘signal-on’ signaling would be generated. Liu et al. [98] developed a highly sensitive enzyme-based E-DNA sensor that pushed the detection limit down to low femtomolar concentrations (Fig. 5b), nearly three orders of magnitude lower than that of the initial E-DNA setup [94], where the sensitivity is limited to picomolar concentrations, partially because one Fc label can only transfer one electron from/to the Au electrode. In this work, however, one HRP can convert ∼104 reactions of hydrogen peroxide to water; the target binding-induced signal change could be greatly amplified. Wei et al. [99] developed a similar sensor for electrochemical detection of salivary mRNA targets with 0.4 fM sensitivity. This hairpin capture probe can also be designed with an external toehold and immobilized on the gold electrode surface of quartz crystal microbalance (QCM) [100]. Through the toehold-mediated strand displacement reaction, this QCM biosensor achieved a highly selective and sensitive detection of single-nucleotide polymorphism (SNP) in the p53 tumor suppressor gene. In addition to typical hairpin DNA probes, it is also possible to design other dynamic probes, such as partial duplex [101], pseudoknot [102], triple-stem [103] and triplex structures [104], to engineer biosensing interfaces. For example, a partially complementary duplex probe was designed for a signal-on, label-free electronic DNA sensor that achieved a sub-picomolar concentration detection limit with a strand displacement mechanism (Fig. 5c) [101]. This improved detection limit is believed to be attributed to the rapid shift from rigidity (duplex) to flexibility (ssDNA) of the sensing probe. Next, a special pseudoknot probe that consists of two stem-loop structures sharing one strand as the stem or loop was deployed directly in complex matrices, such as blood serum for highly sensitive DNA detection (Fig. 5d) [102]. Furthermore, Xiao et al. [103] used a triple-stem structured probe to perform SNP assay even in the complex medium of blood serum (Fig. 5e). A thermodynamic analysis suggested that this triple-stem probe takes advantage of the large thermodynamic changes in enthalpy and entropy resulting from major conformational rearrangements in target-responsive binding, leading to exquisite sensitivity to single-base mismatches [103]. Another highly specific E-DNA sensor was also developed by using a triplex probe (Fig. 5f) [104]. In this approach, a clamp-like DNA probe can bind with its complementary target sequence through two distinct hybridization modes, namely Watson-Crick recognition and Hoogsteen interaction. These bindings lead to the formation of a triplex that can improve both the affinity and specificity of molecular recognition. Functional nucleic acids, such as DNAzymes and aptamers (RNA or DNA), can also be used to dynamically engineer the biosensing interface. A typical Pb2+-responsive DNAzyme probe complex was employed for electrochemical detection of Pb2+ [105]. In the presence of Pb2+, the rigid structure of the complexes would be cleaved, which produced specific electrochemical signals. Similarly to ion-dependent DNAzyme, aptamers change their conformations in response to ions (Fig. 5g). The presence of K+ can induce a contractile DNA duplex probe (K+ aptamer), which incorporated two short separated motifs of G/G mismatches, which transited from the duplex DNA to the G-quadruplex, and thus changed the charge-conducting properties of the DNA structure [106]. Beyond ions, aptamer can also specifically bind with small molecules and biomacromolecules. Zuo et al. [107] designed an adenosine triphosphate (ATP) sensor using an anti-ATP aptamer (Fig. 5h). Their study verified that ATP could stabilize the tertiary aptamer structure at the surface and could responsively denature the initial aptamer duplex. By incorporating a DNA aptamer into the three-way junction elements, a rationally designed probe architecture for the detection and quantification of thrombin (a plasma protein) was reported [108]. Compared to static engineering, dynamic engineering has shown superior properties in several aspects. First, the enhanced rigidity over the linear probes offers better-ordered orientation of DNA probes at the biosensing interface. Second, because of their internal hybridized nature, these dynamic probes are inherently resistant to inter-probe interactions at the interface. Third, thermodynamic studies with foldable probes (single- or triple-stem structure) have demonstrated their capability to sensitively discriminate single-base mismatched DNA. However, it should be noted that foldable probes are not sufficiently rigid to survive on highly crowded surfaces; hence, appropriate modulation of the surface density is still critically important for high-performance sensors. In addition, MCH is still required to help stem-loop structures to stand upright at the surface, which makes the surface inevitably a heterogeneous one. PERSPECTIVES In this review, we summarize the biomolecular nanostructures-mediated interfacial engineering, in which proteins and nucleic acids are capable of engineering the biosensing interface in static or dynamic manners to improve sensitivity and specificity. In particular, the precise gene manipulation or DNA nanotechnology allows the design and engineering of proteins with site-specific modification and DNA nanostructures, respectively, which considerably improves the capacity to control the probe density, orientation, homogeneity and accessibility at the interface. Based on these interface engineering strategies, a variety of biosensors have been tailored to probe biochemical substances, antigen-antibody binding or nucleic acid hybridization for highly sensitive detection of health-related enzyme substrates and molecular biomarkers, such as microRNA, DNA and protein, as well as tumor-derived exosomes and cancer cells. Despite the progress, several formidable challenges exist. First, in protein-based interface engineering, the indirect use of protein A or G reduces the efficiency. Thus, a more direct approach for anchoring oriented protein at the interface should be explored in the future, such as direct coupling of the Fc portion of the antibody with a rigid spacer (e.g. dsDNA) for covalent immobilization, which permits proteins with favorable spatial orientation and homogeneity in a rational density. Designing one-step conformation-switch protein with large distance change is another way to enhance the target response of the protein interface in future dynamic sensors [38]. Second, in DNA-based interface engineering, the rigidity- and flexibility-tunable DNA probe is highly desirable in designing surface-confined single-step DNA sensors by controlling the upright orientation and conformation switch. DNA tetrahedron is a versatile rigid scaffold that allows the incorporation of different soft functional nucleic acids into one or two edges of the 3D structure, thereby forming dynamic structures in response to different target molecules [109,110]. DNA-conjugated protein probes should be explored, which can combine the advantages of each other (e.g. high stability of DNA and rapid binding rate of antigen antibody), and meanwhile offset shortcomings in part, like the instability of proteins in practical use and regeneration [111]. Of note, the synergistic complex structure composed of DNA nanostructure and antibody (TDN-antibody) may provide additional advantages: (i) the framework of TDN serves as a rigid spacer that endows the antibody with an orderly upright orientation and solution-phase setting; (ii) the comparative size between TDN and antibody (∼5 nm) enables the protein probes to be immobilized with a controlled lateral distance and favorable accessibility; (iii) through simple thermal denaturation of bridged DNA, the antibody layer could be regenerated, guaranteeing a renewable sensing interface. The purity of the assembled DNA nanostructures or DNA-protein conjugates is another concern, which is limited by the separation techniques, such as electrophoresis and high-performance liquid chromatography. Hence, the approaches with higher separation efficiency, for example, single-step magnetic isolation, are required to be developed in future. Moreover, advanced synthetic biology [42], smart bioresponsive materials [112] and genome-editing technology (e.g. CRISPR-Cas9) [113] might be exploited to design and modify the functional protein structures with more favorable interfacial orientation and accessibility compared to the routine means. These artificial protein or DNA probes may also be synergized with the high-curvature nanostructured substrates [67] to enhance the interface-modulated capacity. When interfaced with a single-step, multiplexed microfluidic system, the functional structured bioprobes would be promising to enable precise quantification of clinically related multi-targets at different levels, particularly in resource-limited settings [114–117]. We anticipate such biomolecule-driven interface-modulated strategies will ultimately pave the way towards a new generation of molecular diagnosis and therapy. FUNDING This work was supported by the Ministry of Science and Technology of China (2013CB932803, 2013CB933802, 2016YFA0201200, 2015CB559100, 2016YFC0903402), National Natural Science Foundation of China (21775034, 21305034, 21390414, 21227804), Hubei Province Health and Family Planning Scientific Research Project (WJ2017Q032) and the Strategic Priority Research Program B (Pre-research) of the Chinese Academy of Sciences (CAS). REFERENCES 1. Shao W , Zhang X , Liu H et al. Anchor-chain molecular system for orientation control in enzyme immobilization . Bioconjugate Chem 2000 ; 11 : 822 – 6 . https://doi.org/10.1021/bc000029s Google Scholar Crossref Search ADS 2. Drummond TG , Hill MG , Barton JK . Electrochemical DNA sensors . Nat Biotechnol 2003 ; 21 : 1192 – 9 . https://doi.org/10.1038/nbt873 Google Scholar Crossref Search ADS PubMed 3. Shi J , Zhang X , Xie W et al. Improvement of homogeneity of analytical biodevices by gene manipulation . 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Nat Med 2011 ; 17 : 1015 – 9 . https://doi.org/10.1038/nm.2408 Google Scholar Crossref Search ADS PubMed © The Author(s) 2018. Published by Oxford University Press on behalf of China Science Publishing & Media Ltd. This article is published and distributed under the terms of the Oxford University Press, Standard Journals Publication Model (https://academic.oup.com/journals/pages/open_access/funder_policies/chorus/standard_publication_model) http://www.deepdyve.com/assets/images/DeepDyve-Logo-lg.png National Science Review Oxford University Press

Biomacromolecular nanostructures-based interfacial engineering: from precise assembly to precision biosensing

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Oxford University Press
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© The Author(s) 2018. Published by Oxford University Press on behalf of China Science Publishing & Media Ltd.
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2095-5138
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2053-714X
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10.1093/nsr/nwx134
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Abstract

Abstract Biosensors are a type of important biodevice that integrate biological recognition elements, such as enzyme, antibody and DNA, and physical or chemical transducers, which have revolutionized clinical diagnosis especially under the context of point-of-care tests. Since the performance of a biosensor depends largely on the bio–solid interface, design and engineering of the interface play a pivotal role in developing quality biosensors. Along this line, a number of strategies have been developed to improve the homogeneity of the interface or the precision in regulating the interactions between biomolecules and the interface. Especially, intense efforts have been devoted to controlling the surface chemistry, orientation of immobilization, molecular conformation and packing density of surface-confined biomolecular probes (proteins and nucleic acids). By finely tuning these surface properties, through either gene manipulation or self-assembly, one may reduce the heterogeneity of self-assembled monolayers, increase the accessibility of target molecules and decrease the binding energy barrier to realize high sensitivity and specificity. In this review, we summarize recent progress in interfacial engineering of biosensors with particular focus on the use of protein and DNA nanostructures. These biomacromolecular nanostructures with atomistic precision lead to highly regulated interfacial assemblies at the nanoscale. We further describe the potential use of the high-performance biosensors for precision diagnostics. biosensor, interface engineering, homogeneity, orientation control, fused proteins, DNA nanostructures INTRODUCTION The biosensing interface between the solid–liquid phases plays an important role for mass transfer, electron mobility, energy exchange and signal transduction. By tailoring the interface with functional molecules or materials, surface properties can be effectively modulated to ensure the homogeneity of the biosensor surface and thus to improve biomolecular recognition, such as nucleic acid hybridization or antigen-antibody binding for efficient optical or electronic signal transduction [1–5]. Given that this transduction process generally occurs at the solid–liquid interface, in-depth understanding of the interactions between biomolecules and the supporting interface is indispensable. For example, (i) what is the main difference between surface-confined and solution-phase biomolecular recognition [6]; (ii) how does interfacial architecture affect biomolecular adsorption, assembly, folding and diffusion [7,8]; and (iii) how do surface properties influence the folding free energy and structural dynamics of biomolecules [9,10]? Along the lines to address these questions, a number of strategies have been proposed to develop high-performance biosensing interfaces with improved biomolecular recognition efficiency and reduced non-specific adsorption [11]. Great efforts have been made to increase the homogeneity and orderliness of interfacial assembly of enzyme proteins, antibodies or DNA probes [12]. Especially, rational interfacial design has proven to effectively realize the orientation control of the biorecognition molecules or reduce non-specific adsorption and avoid false positives. Among these approaches, biomimetic interfacial engineering has shown great promise for in-vitro biodetection [13]. In nature, biorecognition processes are generally highly efficient and specific, which are often based on synergic functions of ‘molecular machinaries’ that comprise multiple biomolecules in molecule-crowding intracellular settings. Construction of biomacromolecular nanostructures represents a promising means to mimic their intracellular counterpart and engineer biosensing interfaces in a programmable way [14]. In this review, we aim to summarize recent progress in using biomacromolecule nanostructures for interfacial engineering of biosensors primarily on gold surfaces and with particular focus on structured protein and DNA probes. Biosensors based on other surfaces such as graphene oxide, metal oxides and 2D transition-metal dichalogenides nanomaterials have also been explored and could be found in several excellent reviews [15–18]. We also provide perspectives on their applications in precision diagnosis. PROTEIN-BASED INTERFACIAL ENGINEERING Bioactivities of protein molecules (i.e. antibodies and enzymes) are responsible for high efficient antigen recognition or direct electron transfer at the sensor interface. Nevertheless, it is challenging to retain their natural bioactivities and conformations when proteins are immobilized on the sensing interface via physical adsorption or covalent conjugation [19]. Since proteins can interact with the interface via their various surface groups, they tend to adsorb on the interface with random orientations that often do not favor their bioactivities [20]. Moreover, surface attachment might also alter the native conformation of proteins [21], thereby degrading their functions. Hence, a range of static and dynamic interfacial engineering approaches have been proposed to assemble proteins on the interface with better orientation and freedom for conformational change and functional modulation. STATIC ENGINEERING BY CONTROLLED IMMOBILIZAITON OF PROTEINS Orientation control of proteins at the interface is highly required to retain their bioactivities. However, proteins often adopt random orientations at the biosensing interface due to their multi-site interactions with the substrate. For example, Immunoglobulin Gs (IgGs) (Fig. 1a) consist of one constant fragment (Fc) and two variable antigen-binding fragments (Fab), resulting in random orientations of interfacial proteins: head-on (both Fab on surface), tail-on (Fc on surface), side-on (Fc and one Fab on surface) or flat-on (all fragments on surface) (Fig. 1b). To achieve optimal antigen-binding activity, antibodies should be immobilized with favorable orientation of tail-on (Fig. 1c). Efforts to address this problem have been focused on precisely defined protein adorsption (non-covalent) and anchoring strategies (covalent) [19]. Figure 1. View largeDownload slide Structure and interfacial orientation of natural antibodies. (a) Schematic depiction of the structure of natural antibody with two Fab and one Fc domain, as well as multiple reactive groups, such as amine, sugar and disulfide group. Antibody can be immobilized in either a random (b) with four typical orientations or oriented (c) fashion. Figure 1. View largeDownload slide Structure and interfacial orientation of natural antibodies. (a) Schematic depiction of the structure of natural antibody with two Fab and one Fc domain, as well as multiple reactive groups, such as amine, sugar and disulfide group. Antibody can be immobilized in either a random (b) with four typical orientations or oriented (c) fashion. Intermediate proteins (e.g. protein A or protein G) (Fig. 2a) with multiple binding domains specific to the Fc part of the antibody were employed to immobilize the antibody with favorable tail-on orientation. As compared to their randomly immobilized counterparts, the oriented antibodies improved the biosensing sensitivity and specificity [22,23]. However, even partially oriented intermediate proteins may induce random orientations of antibodies, which still poses a limit. To further improve the protein orientation, a highly organized IgG-protein with 3D structures and exposed antigen-binding sites was designed [24]. However, this oriented immobilization strategy is heavily dependent on solution pH that negatively affects the protein G orientation on gold and subsequent antibody binding [25]. Cysteine-functionalized protein G (Cys-protein G) has proven successful for controlled orientation of protein A or protein G at the interface [26]. A recombinant Cys-protein G trimer synthesized by repeated linking of protein G via flexible linkers can enhance the bioactivities of antibodies immobilized on magnetic silica nanoparticles. Compared to a Cys-protein G monomer probe, the trimer improved the sensitivity by 10-fold. By genetically fusing protein A domain to a Cys-exposing variant of Escherichia coli protein ompA, favorable binding orientation and antibody presentation were realized [27]. Similarly, a ‘super-oriented IgG’ was constructed to attach to the oriented protein A at the interface via enzyme conjugation, leading to enhanced affinity up to ∼100 fold over the partially oriented IgG physical immobilized at interface [28]. Figure 2. View largeDownload slide Protein-based interface modulation. (a) Protein A/G mediated non-covalent orientation of antibodies. (b) Covalent orientation of antibodies by either reactive groups distributed on the surface of antibody or fusion protein technology. Adapted from [34]. (c) Protein-DNA conjugates mediated orientation of antibodies. (d) Proteins reversibly modulate the sensing interface. Adapted from [36]. (e) MBP dynamically modulates the electrochemical interface. Adapted from [40]. Figure 2. View largeDownload slide Protein-based interface modulation. (a) Protein A/G mediated non-covalent orientation of antibodies. (b) Covalent orientation of antibodies by either reactive groups distributed on the surface of antibody or fusion protein technology. Adapted from [34]. (c) Protein-DNA conjugates mediated orientation of antibodies. (d) Proteins reversibly modulate the sensing interface. Adapted from [36]. (e) MBP dynamically modulates the electrochemical interface. Adapted from [40]. Although the interfacial orientation control with intermediate proteins can enhance biosensing performance, this non-covalent interaction between intermediate protein and interface is generally weak and sensitive to environmental changes, such as pH, ionic strength and temperature [25]. In response, direct covalent immobilization of proteins provides a feasible way to engineer the biosensing interface. Classic reactions of inherent reactive groups and tailor-made modifications on protein molecules have been used to improve the orientation and homogeneity of proteins at the interface (Fig. 2b). Amine (NH2) groups in the lysine side-chain on the antibody surface are usually used for relatively random covalent immobilization, although it generally lacks orientation control [29,30]. To improve the orientation control, in contrast, the unique carbohydrate moiety at the Fc portion of an antibody is more amenable to orient proteins via covalent immobilization. In the presence of periodate sodium [31] or boronic acid [32], the carbohydrate vicinal hydroxyl groups would be oxidized to aldehydes and cyclic boronate esters, respectively, which provide oriented covalent antibody immobilization (Fig. 2b). Analogously, the disulfide group in the hinge region of antibody was also exploited for oriented immobilization of antibody via reduction into thiol groups (Fig. 2b) [33]. These methods of covalent coupling of antibody at the interface has proven widely useful for antibodies with multiple activatable sites on antibodies. In addition, DNA can also be used to immobilize antibody, serving as both spacer and binding domain. Heath et al. [34] established a DNA-encoded antibody library for spatially multiplexed detection of nucleic acids, proteins and cells (Fig. 2c). Importantly, the hybridized duplex probe is rigid enough to maintain a favorable orientation (upright) of the antibody at the interface [35]. More importantly, this DNA-antibody conjugate renders the biosensing interface regenerable via facile water rinsing and robust to resist the denaturation induced by surface effect. Scaffolded protein nanostructures provide a powerful means to oriented assemble proteins at the interface. For example, Zhang et al. fused a coding sequence of streptavidin recognition peptide (streptag) to a specific site (3΄ end) of the phoA gene that codes for E. coli alkaline phosphatase (EAP) [1]. This fused protein was used for oriented immobilization of EAP enzyme on microtiter plates via streptag–streptavidin binding. With the insertion of a flexible linker peptide coding sequence between the streptag and EAP sequences, the enzyme activity of immobilized EAP increased by ∼8.4-fold over the fused protein without linker sequence. The linker peptide provides a spacer that minimizes the steric hindrance between EAP and streptavidin, enhancing the orientation effect. In a further step, by incorporating a linker peptide with a cysteine residue at the C-terminal of glucose oxidase (GOx), they engineered a fusion structure (GOx-linker-cysteine) that can be immobilized on gold surfaces with Au–S bond or on a silanized glass surface via disulfide bond (Fig. 2b) [3]. Importantly, this fused cysteine enables the GOx to be immobilized at the interface with well-controlled orientation, thus forming a homogeneous biosensing interface. With the synergistic effect of the linker spacer and interfacial homogeneity, a higher and more stable electrochemical current response was obtained as compared to the GOx without fused cysteine and linker spacer. Although the covalent orientation of proteins is effective, the high-affinity interaction (e.g. biotin–streptavidin binding) is usually irreversible, which makes the biotin–streptavidin based assembly non-regenerable. To realize a regenerable biosensing interface (e.g. surface plasmon resonance, SPR), two streptavidin affinity tags, nano-tag and streptavidin-binding peptide (SBP-tag), were employed (Fig. 2d) [36]. Both of them can specifically interact with streptavidin while the binding affinity (KD ∼4-17 nM) is weaker compared to the biotin–streptavidin binding (KD ∼1 fM), thus enabling the easily controlled association and dissociation. With this tunable binding mechanism, Zhang and coworkers developed a SPR biosensor chip coated with streptavidin for reversible, site-directed protein immobilization mediated by the nano-tag [36]. The streptavidin surface could be regenerated repeatedly without loss of activity even by injection of 50 mM of NaOH solution. This reversible biosensing interface could be readily generalized to build other SPR biosensors, permitting anchoring of various proteins on the streptavidin surface in a stable, site-directed and reversible fashion. DYNAMIC ENGINEERING BY CONFORMATIONAL CHANGE OF PROTEINS In nature, biomolecular recognition is generally accompanied by conformational changes of proteins, such as ligand-induced protein folding or unfolding. Such a special conformation-switchable mechanism can be used to engineer the biosensing interface in a dynamic manner [37]. It can also be extended to design a biosensor exclusively based on the conformational switch of proteins [38]. Nevertheless, most proteins, unlike single-stranded (ss)DNA, do not undergo significant conformational changes upon ligand binding. Hence, a highly sensitive method that can precisely probe the weak conformational alterations is desirable. Extensive studies have shown that electrochemical sensors are capable of tracing the relatively small conformational changes of a surface-confined, redox-labeled macromolecule. Since heterogeneous electron transfer between an electrode and a surface-confined redox molecule exhibits an exponential dependence on both distance and the Marcus coupling factor [39], even small conformational changes in proteins can produce large variations in electron-transfer rates, which in turn translate into measurable changes in electrochemical signals. For example, Benson et al. [40] exploited the ligand-induced hinge-bending motions in the electrode surface-confined maltose binding protein (MBP) that they used to engineer the biosensing interface in a dynamic fashion (Fig. 2e). In this report, a gold electrode was first assembled with a functional monolayer, serving as a binding interface for site-specific immobilization of proteins [40]. Subsequently, the MBP protein was immobilized on the electrode surface with a specific orientation such that a redox reporter group (ruthenium complex) was fixed on the electrode for electrochemical readout. As the target ligand of maltose binds, the target-induced hinge-bending motion in the protein proceeds, which moves the Ru(II) reporter away from the electrode. This target-responsive dynamic modulator in turn triggers a concentration-dependent decrease in the observed electrochemical response, which provides the electronic detection of maltose at mM concentrations. It should be noted that the relatively low sensitivity (at the mM level) of this dynamic sensor is limited by the low affinity of MBP against maltose (KD = 4 mM) but does not pose a fundamental restriction of this strategy. Apparently, this method can be generalized to detection of other related natural or designed MBP-like proteins that would undergo a similar ligand-induced conformational change [41]. Allosteric protein switches are ubiquitous in the biological signal transduction system, which enable cells to sense and respond to specific molecular events. Inspired by nature, tailor-engineering of protein switches with custom input and output functions is significant in molecular diagnostics and physiological decoding [42]. To demonstrate this concept, RG13, a fusion protein of MBP and TEM1β-lactamase, was engineered as a model protein switch with electrochemically activated switching behavior by reducing the disulfide bonds in the switches [43]. In this approach, an electrochemical signal can be used as an exogenous input to control the on/off state of protein switches by modulating the oxidation state of an introduced disulfide bond on the electrode surface. The presence of maltose is the key to activating the enzyme activity due to the induced large hinge-bending conformational change in the MBP domain of RG13. This strategy allows the allosteric protein switch to dynamically regulate the interface signaling. Although the protein switch or folding is effective to engineer the biosensing interface, the ligand-induced conformational change in protein switches is relatively small. In contrast, a more flexible peptide, like aptamer, can trigger a large structure change by binding to its corresponding antibody. Lai and coworkers [44] reported an electrochemical peptide-based dynamic biosensor for the sensitive and specific detection of HIV anti-p24 antibodies where a highly antigenic epitope from the HIV-1 capsid protein, p24, was used as the recognition peptide by modifying with MB (methylene blue) and thiol at two ends, respectively. Of note, this epitope is a short linear peptide lacking defined secondary structure, and adopts predominantly an α-helical conformation in native state. Thus, the binding of the target antibody facilitates the formation of a more rigid complex that induces a large conformation change. This dynamic peptide-based biosensor achieved a detection limit of 10 nM, and had a dynamic range that is broader than the typical concentration range commonly observed with HIV-infected patients. DNA-BASED INTERFACIAL ENGINEERING Single-stranded (ss-)DNA can hybridize to its complementary DNA strictly obeying the Watson-Crick base-pairing rules, which allows DNA recognition and structural assembly in a highly predictable and finely programmable manner. These advantages open opportunities for biosensing interfacial engineering. It was reported that the large inter-probe distances and upright orientation of surface-tethered DNA probes is imperative to realize efficient hybridization. However, engineering an upright and accessible DNA recognition layer is challenging because of the unexpected surface adsorptions, disordered conformations and inhomogeneity of grafting density of DNA bioprobes at the biosensing interface. Static and dynamic interface engineering using DNA and DNA structures have been developed to overcome this challenge. STATIC ENGINEERING BY CONTROLLED IMMOBILIZATION OF DNA PROBES In a typical ssDNA probe-based DNA biosensor, an efficient probe–target hybridization process is fundamental for improving the biosensing performance, particularly in sensitivity and specificity. Nevertheless, high hybridization efficiency depends on a favorable orientation (upright) of DNA probes and rational inter-probe distance (Fig. 3a) [45]. It is expected that the ssDNA probes can adopt an upright orientation at the biosensing interface via a single-point attachment. By taking thiolated DNA (SH-DNA) as an example, however, significant interactions exist between the DNA bases and the gold surface via multiple nitrogen atoms [46], which allows ssDNA molecules to lie down on the Au surface, resulting in a largely limited accessibility of the target sequences with reduced hybridization efficiency (Fig. 3b). Such non-specific adsorption onto the Au via Au–N interactions has been confirmed by Tarlov and coworkers [47]. Further characterization, such as X-ray photoelectron spectroscopy and Fourier transform infrared spectroscopy, revealed that the nonspecifically adsorbed ssDNA could not be removed, even with extensive rinsing or heating to 75°C, and the ‘specifically’ anchored ssDNA monolayer is not oriented perpendicularly to the surface, especially at low densities. Short DNA probes tended to orient parallel to the surface, whereas the relative long strands preferred to form disordered film due probably to adjacent entanglement [48]. Although the densely packed short DNA probes may assume an upright conformation on the surface, the restricted target accessibility would offset this advantage of favorable probe orientation (Fig. 3c). Figure 3. View largeDownload slide ssDNA-based interface modulation. (a) Schematic depiction of an ideal DNA recognition layer with a large intermolecular distance and linear upright orientation. (b) Sparse probes are prone to lying flat-on the surface due to non-specific adsorption. (c) Densely packed short DNA probes adopt an upright conformation, but yield poor hybridization efficiency because of reduced accessibility. (d) Diluents, such as MCH and OEG, co-assembled with ssDNA to favor the upright orientation on the surface and reduce non-specific adsorptions. Figure 3. View largeDownload slide ssDNA-based interface modulation. (a) Schematic depiction of an ideal DNA recognition layer with a large intermolecular distance and linear upright orientation. (b) Sparse probes are prone to lying flat-on the surface due to non-specific adsorption. (c) Densely packed short DNA probes adopt an upright conformation, but yield poor hybridization efficiency because of reduced accessibility. (d) Diluents, such as MCH and OEG, co-assembled with ssDNA to favor the upright orientation on the surface and reduce non-specific adsorptions. To help address this dilemma, Tarlov et al. [49] introduced a coassembling small molecule (mercaptohexanol, MCH), synergizing with SH-DNA to engineer the recognition ability of bioprobes at the interface (Fig. 3d). This ‘helper’ molecule is able to largely remove the nonspecifically adsorbed DNA and meanwhile protrudes the surface-attached DNA probes into solution phase via the repulsion between the net negative dipole of alcoholic terminus and the negatively charged DNA backbones. These mixed ssDNA/alkylthiol monolayers have been extensively investigated by varied surface-measuring techniques and confirmed with a favorable upright orientation of ssDNA bioprobes [50]. This synergic static modulation strategy was widely employed in engineering DNA biosensors and biochips [51–53]. However, the co-assembled small molecule diluents (MCH) cannot resist the non-specific protein adsorption effectively. In response, a ssDNA/oligo-ethylene glycol (OEG) mixed monolayer was used to improve the protein resistance of the biosensing interface, particularly in complex matrices, such as blood [54,55]. We also note that relatively rigid double-stranded (ds) DNA molecules represent another route to improving interfacial probe arrangement with enhanced sensitivity and specificity, especially for toehold design [56]. Although helix structures or the co-assembled diluents can aid DNA probes to adopt a favorable orientation at the interface to some extent, the probe density is still a critical factor controlling the kinetics of target/probe hybridization [57]. Interestingly, a simple optimization of the assembly concentration of DNA probes can accurately modulate the interfacial probe average density. One concern in this approach is that local lateral interactions inevitably exist in DNA films, particularly for long sequences [58,59]. This enables the prediction of the most favorable hybridization in the ‘Langmuir’ (L) regime hard to reach, because it only exists in the limits of sparse films where probes are so far apart that they do not interact with each other [60]. From a single-molecule view [61], the aggregation patches on the Au surface can significantly reduce target accessibility, which, nevertheless, could not be easily eliminated using MCH as the dilution molecule. To better address this challenge, these empirically ‘static’ modulation methods dependent on the scheduled probe/diluent ratio may need to couple with nanostructured surfaces or conceptually new probe-design strategies. DNA nanostructures can be used to construct a scaffolded biosensing interface to tune the sensitivity of biosensors in a programmable fashion [14,62]. So far, researchers have engineered a variety of DNA nanostructures with well-defined dimension, topography and precisely controlled functions by using DNA nanotechnology. These functional DNA nanostructures have been actively exploited to develop in-vivo or in-vitro biocomputing and biosensing devices [63–65]. Because the sensitivity of biomolecular detection depends not only on the affinity between biomolecules, but also on the interfacial properties of the biosensors [66], the size reduction of biosensors, particularly to the nanoscale, usually accelerates the mass transport rate and improves the sensitivity [67,68]. However, the limited space available in nanosensors restricts the effective probe numbers and biorecognition events. To address this challenge of size reduction, a trans-scale biosensor that incorporates nano-architectures into macroscopic surfaces is necessary [5]. Nevertheless, reproducible engineering of nanostructured surfaces with well-defined topography remains technically difficult, though high-cost photolithography potentially offers a route to the fabrication of nanostructures at the wafer-scale [69]. Recently, Fan et al. [14] developed a conceptually new ‘soft lithographic’ strategy to reproducibly engineer and programmably modulate a biosensing interface using well-defined 3D DNA nanostructures (Fig. 4a). By patterning the macroscopic gold electrode with tetrahedral DNA nanostructures (TDNs) varying in sizes, the detection limit of DNA sensors can be programmably tuned over four orders of magnitude. Figure 4. View largeDownload slide DNA nanostructure-based programmable modulation on the sensing interface. (a) Assembling DNA nanostructures with varying sizes for programmable modulation on the sensing interface [14]. (b) A universal biosensing platform based on tetrahedron-structured DNA probes (TSPs) [70]. Reprinted with permission from [14,70]. Copyright 2015 and 2010 by WILEY-VCH, Weinheim, respectively. (c) TSPs-based E-DNA sensor for microRNA detection [85]. (d) TSP-conjugated antibody for sensitive PSA detection amplified with HRP-AuNP [79]. (e) TSP-conjugated aptamer for sensitive exosome detection [82]. (f) TSP-conjugated aptamer-HCR for sensitive detection of cancer cells [83]. (g) TSP-mediated capillary microarray for multiplexed bioassays, achieving detection limits of 1 μM and 0.1 nM for small molecules (ATP and cocaine), respectively [93]. Reprinted with permission from [79,82,83,93]. Copyright 2014 and 2017 by the American Chemical Society. Figure 4. View largeDownload slide DNA nanostructure-based programmable modulation on the sensing interface. (a) Assembling DNA nanostructures with varying sizes for programmable modulation on the sensing interface [14]. (b) A universal biosensing platform based on tetrahedron-structured DNA probes (TSPs) [70]. Reprinted with permission from [14,70]. Copyright 2015 and 2010 by WILEY-VCH, Weinheim, respectively. (c) TSPs-based E-DNA sensor for microRNA detection [85]. (d) TSP-conjugated antibody for sensitive PSA detection amplified with HRP-AuNP [79]. (e) TSP-conjugated aptamer for sensitive exosome detection [82]. (f) TSP-conjugated aptamer-HCR for sensitive detection of cancer cells [83]. (g) TSP-mediated capillary microarray for multiplexed bioassays, achieving detection limits of 1 μM and 0.1 nM for small molecules (ATP and cocaine), respectively [93]. Reprinted with permission from [79,82,83,93]. Copyright 2014 and 2017 by the American Chemical Society. A typical DNA tetrahedron-structured probe (TSP) was self-assembled with a relatively long, probe-bearing oligonucleotide and three thiolated oligonucleotides, which carries a pendant DNA probe (hybridizable domain) at one vertex and three thiol groups (-SH) at the other three vertices (Fig. 4b), allowing it to firmly anchor on the gold surface via Au–S bonds [70]. Of note, this 3D-structured probe demonstrates several advantages over its single-stranded or duplex counterparts in the engineering of biosensing interfaces. First, the high mechanical rigidity of TSP allows it to adopt a highly ordered, upright orientation at the Au surface, which was stabilized by the three thiol ‘legs’ that greatly increase the stability of the surface-confined probes by ∼5000 times as compared to the mono-thiolated DNA strands [71]. Second, the TSP structures avoid inter-probe entanglement (in ssDNA probes) via spatially segregated pendant probes, and also reduce the surface effects by positioning the probes in solution-phase-like settings [72]. Importantly, it is able to precisely anchor a single DNA probe on a single TDN, which is hard to achieve with inorganic nanostructures. Third, the TSP-decorated surfaces are protein-resistant, and the ‘helper’ molecule, MCH, is not necessary in this case, both of which allow the TSP-based sensors to be deployed directly in complex matrices, such as serum, for practical applications. In addition, diffusion and convection at the nanostructured interfaces are also expected to be higher than those at macroscopic ones [68]. In TSP-based sensors, the lateral distance between probes is dominated by the TDN scaffold, and thus this distance can be precisely tuned at the nanometer scale by varying the size of TDNs [73]. To this end, five types of TDNs (TDN-7, TDN-13, TDN-17, TDN-26 and TDN-37) with different sizes have been rationally designed [14]. The numbers (7, 13, 17, 26 and 37) designate the base pairs of each edge in the TDNs. Because each base pair was separated by 0.34 nm in a double helix, the edge lengths of the TDNs were able to be calculated with precision, equal to 2.4, 4.4, 5.8, 8.8 and 12.6 nm, respectively (Fig. 4a). It was observed that the electrode surface density of TDN probes was inversely proportional to the size of the TDN, which was further confirmed by statistical analysis that revealed the nearly linear relationship between the lateral distance and TDN size. Such an ability to precisely regulate the lateral distance can also tune the surface hybridization capabilities, including hybridization kinetics and efficiency, which depends heavily on the probe distance. By regulating the probe distance via the deployment of TDNs with varied size, the detection limit of the biosensor was programmed [14]. This result is consistent with that reported by the Howorka group [74], in which the molecular recognition of receptors on TDN-scaffolded surface was improved by an order of magnitude. With the use of TDNs, a variety of ultrasensitive biodetections, such as nucleic acids [70,75–78], proteins [79–81], exosomes [82] and cells [83], have been conducted. By using TSP-patterned gold electrodes, Pei et al. [70] engineered a biosensor with a typical sandwich hybridization strategy, in which the upright probe can selectively capture its complementary sequence with high discrimination ability towards non-cognate sequences (Fig. 4b). The free domain was subsequently flanked by the biotinylated reporter probe that induced the production of electrochemical signals via the substrate catalysis of horseradish peroxidase (HRP). In combination with hybridization chain reaction amplification, this tetrahedral platform improved the detection limit to 100 aM [76]. For microRNA detection, a synergic TSP-based electrochemical sensor was constructed by integrating with rolling circle amplification (RCA) and silver nanoparticles that were attached to the RCA products in tandem [84]. This TSP-RCA sensor achieved a limit of detection as low as 50 aM. To further improve the sensitivity, Wen et al. [85] increased the inter-probe distance using interfacial engineering and multi-enzyme amplification, which yields an extremely low detection limit of 10 aM (∼600 microRNA molecules in 100 μL) (Fig. 4c). In addition to direct use as a capture probe, the TDN can also serve as a probe scaffold to build immuno-sensing layers by conjugating with antibodies. For example, antibodies of tumor-necrosis-factor alpha (TNF-a) can be easily and reversibly anchored onto DNA nanostructured surfaces via a DNA-bridged antibody linking with EDC-NHS chemistry, which converts the DNA recognition layer into a protein recognition one [80]. Next, Zuo et al. [79] reported another antibody immobilization strategy that uses highly efficient click chemistry instead of a carbodiimide reaction to conjugate the PSA antibody with a DNA-bridged TDN scaffold (Fig. 4d). By precisely controlling the nanospacing of anchored antibodies with a nanostructured scaffold, they achieved an extremely high sensitivity (1 pg/mL) in PSA detection. Instead, a more direct, DNA-bridge-free approach was introduced to anchor anti-IgG on the TDN scaffold for ultrasensitive IgG and bacteria quantification [81,86,87]. By replacing the pendant probe with an expanded nucleotide-containing aptamer, Tan et al. [82] developed a TSP-assisted aptasensor for direct capture and detection of hepatocellular exosomes (Fig. 4e). This TSP-assisted aptasensor demonstrated improved accessibility and detected the exosome with 100-fold higher sensitivity as compared to the ssDNA-functionalized aptasensor. In combination with a multibranched hybridization chain reaction that offers multiple biotins and branched arms, this TSP-functionalized interface has been engineered to realize multivalent capture and highly sensitive detection of cancer cells (Fig. 4f) [83]. The TSP-based biosensing interface has also proven to be an effective platform for the detection of small molecules, ions and even multiplex targets, because the suspended probe can be easily replaced with various functional nucleic acids. As a proof of concept, a split aptamer for a small molecule, cocaine, was incorporated into TDN to develop a performance-enhanced electrochemical sensor for cocaine [88]. To realize higher sensitivity, another electrochemical cocaine sensor was presented based on structure conversion from a frustum pyramid to an equilateral triangle [89]. In this approach, the presence of cocaine triggered the aptamer-bearing DNA nanostructure change from ‘Close’ to ‘Open’, achieving a detection limit of 0.21 nM. Likewise, an anti-ATP aptamer was immobilized at the top of the tetrahedron, thereby forming an ATP sensor that could detect a concentration as low as 0.2 nM [90]. In conjunction with mercury-specific oligonucleotides, Yin et al. [91] constructed a turn-on Hg2+ sensor with a detection limit of 100 pM, which was 100-fold lower than that obtained with a ssDNA-based sensor. As a versatile probe scaffold, TSP can easily evolve into a microarray platform by replacing the pendant probe with the corresponding target-responsive probes. Recently, Li and coworkers fabricated a high-throughput addressable microarray by covalently coupling TSP onto glass substrates for multiplexed analysis with improved sensitivity and specificity [92]. The detection limits of this sensor array for miRNA (let-7a), PSA and small molecule (cocaine) were 10 fM, 40 pg/mL and 100 nM, respectively. To accelerate the target binding, Qu et al. [93] reported a rapid (<5 min) arrayed DNA-nanostructure-supported aptamer pull-down (DNaPull) assay under convective flux in a glass capillary (Fig. 4g). This DNaPull assay allowed multiplexed analysis of the contents in droplets with nano- or picoliter volumes, achieving detection limits of 1 μM, and 0.1 nM for small molecules (ATP and cocaine) and a biomarker (thrombin), respectively. DYNAMIC ENGINEERING BY CONFORMATIONAL CHANGE OF DNA Dynamic engineering is probably more attractive due mainly to the target-induced conformational change. These conformation-switchable DNA molecules, such as hairpin (quadruplex and pseudoknot with intramolecular secondary structures), aptamer and DNAzyme, possess relatively rigid and ordered structures that may, in principle, prevent inter-strand entanglement at the interface and/or increase the orderliness of DNA bioprobes. In 2003, Heeger and coworkers [94,95] developed a conceptually new electronic equivalent of the molecular beacon, the electrochemical DNA (E-DNA) sensor, which exploits target-induced conformational changes of hairpin (stem-loop) probes on Au electrode surfaces (Fig. 5a). This surface-confined ‘dynamic’ probe was labeled with an electroactive ferrocene (Fc) at the 5΄ end and a thiol at the 3΄ terminal. Initially, the hairpin localizes Fc proximal to the electrode surface, thus enabling a rapid electron transfer. After hybridization, the stem opened and extended to a rigid duplex, forcing Fc away from the electrode surface. This dynamic response allows separation between Fc and the electrode surface reaching several nanometers that yields a large and measurable change in electrochemical signal because of the exponential decay of electron transfer with distance at the interface. This sensor achieved a selective detection of 10 pM target DNA featuring in single-step and electronic (electrochemical) detection. More importantly, this design utilizes only the intrinsic conformational change of surface-confined DNA probes to modulate interface signaling, avoiding the introduction of exogenous reagents. By using such a dynamic, one-step response strategy, Lubin et al. [96] successfully detected DNA authentication tags that are associated with paper or drugs. Later, Lai and coworkers [97] demonstrated a sequence-specific detection of unpurified amplification products of the gyrB gene of Salmonella typhimurium, signifying a promising way towards rapid, sample-to-answer pathogen detection. Figure 5. View largeDownload slide DNA-based dynamic interface modulation. (a) The stem-loop-structured probe design for a reagentless electrochemical DNA (E-DNA) sensor [94]. (b) A signal-on stem-loop-structured E-DNA sensor with HRP-catalyzed signal amplification [98]. (c) Stand displacement reaction-based signal-on electronic detection [101]. Reprinted with permission from [94,101]. Copyright 2003 and 2006 by National Academy of Sciences, USA. (d) Two-stem-loop-structured E-DNA sensor [102]. (e) Three-stem-loop-structured E-DNA sensor [103]. (f) A conformation-switchable triplex E-DNA sensor [104]. (g) DNAzyme-based dynamic probe as an ion (Pb2+) sensor [105]. (h) Dynamic DNA aptamer as an ATP sensor [107]. Reprinted with permission from [98,102–105,107]. Copyright 2007, 2008, 2009 and 2014 by the American Chemical Society. Figure 5. View largeDownload slide DNA-based dynamic interface modulation. (a) The stem-loop-structured probe design for a reagentless electrochemical DNA (E-DNA) sensor [94]. (b) A signal-on stem-loop-structured E-DNA sensor with HRP-catalyzed signal amplification [98]. (c) Stand displacement reaction-based signal-on electronic detection [101]. Reprinted with permission from [94,101]. Copyright 2003 and 2006 by National Academy of Sciences, USA. (d) Two-stem-loop-structured E-DNA sensor [102]. (e) Three-stem-loop-structured E-DNA sensor [103]. (f) A conformation-switchable triplex E-DNA sensor [104]. (g) DNAzyme-based dynamic probe as an ion (Pb2+) sensor [105]. (h) Dynamic DNA aptamer as an ATP sensor [107]. Reprinted with permission from [98,102–105,107]. Copyright 2007, 2008, 2009 and 2014 by the American Chemical Society. The hairpin structure of DNA bioprobes at the interface can be employed to sterically hinder the accessibility of enzymes to bioactive sites at one end of the hairpin DNA, and thus a target-responsive ‘signal-on’ signaling would be generated. Liu et al. [98] developed a highly sensitive enzyme-based E-DNA sensor that pushed the detection limit down to low femtomolar concentrations (Fig. 5b), nearly three orders of magnitude lower than that of the initial E-DNA setup [94], where the sensitivity is limited to picomolar concentrations, partially because one Fc label can only transfer one electron from/to the Au electrode. In this work, however, one HRP can convert ∼104 reactions of hydrogen peroxide to water; the target binding-induced signal change could be greatly amplified. Wei et al. [99] developed a similar sensor for electrochemical detection of salivary mRNA targets with 0.4 fM sensitivity. This hairpin capture probe can also be designed with an external toehold and immobilized on the gold electrode surface of quartz crystal microbalance (QCM) [100]. Through the toehold-mediated strand displacement reaction, this QCM biosensor achieved a highly selective and sensitive detection of single-nucleotide polymorphism (SNP) in the p53 tumor suppressor gene. In addition to typical hairpin DNA probes, it is also possible to design other dynamic probes, such as partial duplex [101], pseudoknot [102], triple-stem [103] and triplex structures [104], to engineer biosensing interfaces. For example, a partially complementary duplex probe was designed for a signal-on, label-free electronic DNA sensor that achieved a sub-picomolar concentration detection limit with a strand displacement mechanism (Fig. 5c) [101]. This improved detection limit is believed to be attributed to the rapid shift from rigidity (duplex) to flexibility (ssDNA) of the sensing probe. Next, a special pseudoknot probe that consists of two stem-loop structures sharing one strand as the stem or loop was deployed directly in complex matrices, such as blood serum for highly sensitive DNA detection (Fig. 5d) [102]. Furthermore, Xiao et al. [103] used a triple-stem structured probe to perform SNP assay even in the complex medium of blood serum (Fig. 5e). A thermodynamic analysis suggested that this triple-stem probe takes advantage of the large thermodynamic changes in enthalpy and entropy resulting from major conformational rearrangements in target-responsive binding, leading to exquisite sensitivity to single-base mismatches [103]. Another highly specific E-DNA sensor was also developed by using a triplex probe (Fig. 5f) [104]. In this approach, a clamp-like DNA probe can bind with its complementary target sequence through two distinct hybridization modes, namely Watson-Crick recognition and Hoogsteen interaction. These bindings lead to the formation of a triplex that can improve both the affinity and specificity of molecular recognition. Functional nucleic acids, such as DNAzymes and aptamers (RNA or DNA), can also be used to dynamically engineer the biosensing interface. A typical Pb2+-responsive DNAzyme probe complex was employed for electrochemical detection of Pb2+ [105]. In the presence of Pb2+, the rigid structure of the complexes would be cleaved, which produced specific electrochemical signals. Similarly to ion-dependent DNAzyme, aptamers change their conformations in response to ions (Fig. 5g). The presence of K+ can induce a contractile DNA duplex probe (K+ aptamer), which incorporated two short separated motifs of G/G mismatches, which transited from the duplex DNA to the G-quadruplex, and thus changed the charge-conducting properties of the DNA structure [106]. Beyond ions, aptamer can also specifically bind with small molecules and biomacromolecules. Zuo et al. [107] designed an adenosine triphosphate (ATP) sensor using an anti-ATP aptamer (Fig. 5h). Their study verified that ATP could stabilize the tertiary aptamer structure at the surface and could responsively denature the initial aptamer duplex. By incorporating a DNA aptamer into the three-way junction elements, a rationally designed probe architecture for the detection and quantification of thrombin (a plasma protein) was reported [108]. Compared to static engineering, dynamic engineering has shown superior properties in several aspects. First, the enhanced rigidity over the linear probes offers better-ordered orientation of DNA probes at the biosensing interface. Second, because of their internal hybridized nature, these dynamic probes are inherently resistant to inter-probe interactions at the interface. Third, thermodynamic studies with foldable probes (single- or triple-stem structure) have demonstrated their capability to sensitively discriminate single-base mismatched DNA. However, it should be noted that foldable probes are not sufficiently rigid to survive on highly crowded surfaces; hence, appropriate modulation of the surface density is still critically important for high-performance sensors. In addition, MCH is still required to help stem-loop structures to stand upright at the surface, which makes the surface inevitably a heterogeneous one. PERSPECTIVES In this review, we summarize the biomolecular nanostructures-mediated interfacial engineering, in which proteins and nucleic acids are capable of engineering the biosensing interface in static or dynamic manners to improve sensitivity and specificity. In particular, the precise gene manipulation or DNA nanotechnology allows the design and engineering of proteins with site-specific modification and DNA nanostructures, respectively, which considerably improves the capacity to control the probe density, orientation, homogeneity and accessibility at the interface. Based on these interface engineering strategies, a variety of biosensors have been tailored to probe biochemical substances, antigen-antibody binding or nucleic acid hybridization for highly sensitive detection of health-related enzyme substrates and molecular biomarkers, such as microRNA, DNA and protein, as well as tumor-derived exosomes and cancer cells. Despite the progress, several formidable challenges exist. First, in protein-based interface engineering, the indirect use of protein A or G reduces the efficiency. Thus, a more direct approach for anchoring oriented protein at the interface should be explored in the future, such as direct coupling of the Fc portion of the antibody with a rigid spacer (e.g. dsDNA) for covalent immobilization, which permits proteins with favorable spatial orientation and homogeneity in a rational density. Designing one-step conformation-switch protein with large distance change is another way to enhance the target response of the protein interface in future dynamic sensors [38]. Second, in DNA-based interface engineering, the rigidity- and flexibility-tunable DNA probe is highly desirable in designing surface-confined single-step DNA sensors by controlling the upright orientation and conformation switch. DNA tetrahedron is a versatile rigid scaffold that allows the incorporation of different soft functional nucleic acids into one or two edges of the 3D structure, thereby forming dynamic structures in response to different target molecules [109,110]. DNA-conjugated protein probes should be explored, which can combine the advantages of each other (e.g. high stability of DNA and rapid binding rate of antigen antibody), and meanwhile offset shortcomings in part, like the instability of proteins in practical use and regeneration [111]. Of note, the synergistic complex structure composed of DNA nanostructure and antibody (TDN-antibody) may provide additional advantages: (i) the framework of TDN serves as a rigid spacer that endows the antibody with an orderly upright orientation and solution-phase setting; (ii) the comparative size between TDN and antibody (∼5 nm) enables the protein probes to be immobilized with a controlled lateral distance and favorable accessibility; (iii) through simple thermal denaturation of bridged DNA, the antibody layer could be regenerated, guaranteeing a renewable sensing interface. The purity of the assembled DNA nanostructures or DNA-protein conjugates is another concern, which is limited by the separation techniques, such as electrophoresis and high-performance liquid chromatography. Hence, the approaches with higher separation efficiency, for example, single-step magnetic isolation, are required to be developed in future. Moreover, advanced synthetic biology [42], smart bioresponsive materials [112] and genome-editing technology (e.g. CRISPR-Cas9) [113] might be exploited to design and modify the functional protein structures with more favorable interfacial orientation and accessibility compared to the routine means. These artificial protein or DNA probes may also be synergized with the high-curvature nanostructured substrates [67] to enhance the interface-modulated capacity. When interfaced with a single-step, multiplexed microfluidic system, the functional structured bioprobes would be promising to enable precise quantification of clinically related multi-targets at different levels, particularly in resource-limited settings [114–117]. We anticipate such biomolecule-driven interface-modulated strategies will ultimately pave the way towards a new generation of molecular diagnosis and therapy. 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Journal

National Science ReviewOxford University Press

Published: Sep 1, 2018

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