Perspective Cite This: J. Am. Chem. Soc. 2019, 141, 1162−1170
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Challenges and Solutions in Developing Ultrasensitive Biosensors Yanfang Wu, Richard D. Tilley, and J. Justin Gooding*
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School of Chemistry, Australian Centre for NanoMedicine, ARC Centre of Excellence in Convergent Bio-Nano Science and Technology, The University of New South Wales, Sydney, New South Wales 2052, Australia what is regarded as “ultralow”. Furthermore, according to IUPAC, sensitivity is the “slope of the calibration curve”, and hence the ability of an analytical method to discriminate between small changes in concentration, as distinguished from the detection limit, which is the “minimum single results which, with a stated probability, can be distinguished from a suitable blank value”.5 Detection limit and sensitivity are naturally linked but should not be used interchangeably. The term “ultrasensitive”, however, is here to stay. As such, we seek to define an ultrasensitive bioanalytical sensor as a biosensor with sufficient sensitivity and low background to allow sub-picomolar detection limits. The distinction of sub-picomolar detection limits is made because it represents a concentration range relevant to the detection of many disease biomarkers where there is an unmet commercial imperativefor example, cancer markers in the blood, blood-borne bacteria, and many viral infections in bodily fluids.6 The emphasis on sample-type in which the biomarker is found is deliberate. The sample being readily accessible and acquired using minimally invasive methods is intertwined with the utility of a device. Further, the required detection limits are often determined by not only the number of biomarkers produced but also the remoteness of the sample collection from the site of the pathology. For example, the blood−brain barrier prevents a sufficient concentration of proteins to cross from the brain into the bloodstream for conventional sensors to detect, but with ultrasensitive bioanalytical sensors, the detection of brain analytes directly in the blood might be possible.7 Hence, bioanalytical sensors with sub-picomolar detection limits will provide new opportunities for monitoring a host of new biomarkers as well as established biomarkers at lower concentrations. The detection of species at previously undetectable concentrations8 could also provide new biological insights. If the affinity of the biorecognition molecule for the target analyte is sufficiently high, then sub-picomolar detection limits are simple to envisage. For example, the dissociation constant for the binding of biotin and streptavidin is on the order of 10−15 M.9 With such a dissociation constant, femtomolar detection limits are expected to be achieved. However, for the vast majority of biomolecular binding pairs, such as most antibody− antigen pairs where the dissociation constants are of the order of 10−8−10−12 M, detection limits in the nanomolar to picomolar range are expected. The development of ultrasensitive sensors is not without its challenges, and some of these challenges are unique to low concentrations. Naturally, one of the challenges is to develop
ABSTRACT: This Perspective focuses on the latest strategies and challenges for the development of bioanalytical sensors with sub-picomolar detection limits. Achieving sub-picomolar detection limits has three major challenges: (1) assay sensitivity, (2) response time, and (3) selectivity (including limiting background signals). Each of these challenges is discussed, along with how nanomaterials provide the solutions. One strategy to gain greater sensitivity involves confining the sensing volume to the nanoscale, as used in nanopore- or nanoparticlebased sensors, because nanoparticles are ubiquitous in amplification. Methods to improve response time typically focus on obtaining an intimate mixture between the sensor and the sample either by extending the length scale of nanoscale sensors using nanostructuring or by dispersing magnetic nanoparticles through the sample to capture the analyte. Loading nanoparticles with many biorecognition species is one solution to help address the challenge of selectivity. Many examples in this Perspective explore the detection of prostate-specific antigen which enables a comparison between strategies. Finally, exciting future opportunities in developing single-molecule sensors and the requirements to go even lower in concentration are explored.
1. INTRODUCTION The field of biosensing had incredible commercial successes in the 1980s with the development of glucose sensors and the lateral flow device such as pregnancy test kits. These successes led to a research emphasis on portable devices for use by nonspecialists.1 Ironically, despite incredible developments in biomolecule immobilization, signal transduction and device integration, glucose sensors and lateral flow pregnancy test kits still dominate commercialized biosensors. This dominance can partly be attributed to the market size, the medical necessity in the case of glucose sensing, and the versatility of the lateral flow device which can be adapted to many different analytes.2 These two technologies are of course not the only successes. If a bioanalytical sensor is regarded as a biological recognition species integrated with a signal transducer, then there are many other examples,3 DNA chips being the most notable of these. There are many future opportunities in the sensing field with regard to detecting low concentrations of analyte.4 It is developments in sensors with very low detection limits, and the challenges this class of biosensor face, that are the focus of this Perspective. The term “ultrasensitive” is frequently used in the literature to describe sensors with ultralow detection limits without defining © 2018 American Chemical Society
Received: August 30, 2018 Published: November 21, 2018 1162
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Journal of the American Chemical Society more sensitive assays. Most frequently this is achieved by using more sensitive detectors or amplification schemes. Amplification can be done either by amplifying the amount of biomarker, as in nucleic acid amplification methods, or by amplifying the signal from each captured molecule, as with enzyme labels.4 However, clearly a more sensitive assay alone is not enough. At low concentrations of analyte, response times are also important. This is evident from the fact that nanopore sensors can detect as low as a single molecule, but quantitative analysis is often limited to micromolar or nanomolar concentration ranges due to the slow response times.10 Hence, the mass transport of the analyte to the sensing interface is a key obstacle in getting lower detection limits.11,12 The third major challenge for ultralow detection limit sensors is selectivity, most notably counterproductive nonspecific effects generating spurious signals in the sensor. Nonspecific effects are an enduring challenge for all sensing applications which are exacerbated with ultralow detection limits. These three challenges(1) assay sensitivity, (2) response time, and (3) selectivityare, however, not insurmountable challenges. There are now a multitude of bioanalytical sensors published, and some even commercialized, that have risen to these three challenges. There are some common features in many of these technologies and also a number of especially unique solutions. The purpose of this Perspective is to highlight the recent achievements in developing ultralow detection limit bioanalytical sensors and discuss the lessons learned with regard to the three key challenges listed above. This Perspective will conclude with a discussion of future prospects, such as the need for single-molecule quantitative sensors and rare cell sensors.
Figure 1. (a) Cross section of a heptameric α-hemolysin (αHL) pore sitting in a lipid bilayer as used in a nanopore sensor. (b) Concentration dependence of a cyclic peptide, cyclo[(L-Arg-D-Leu)4-], (RL)4, in the monitored current traces from a single αHL pore under a constant voltage (+80 mV). Note that as the concentration of analyte decreases, the frequency of the resistance spikes decreases such that the challenge of low detection limits becomes one of mass transport rather than the sensitivity of the transducer. Adapted with permission from ref 15. Copyright 2000 American Chemical Society.
levels where the quantification is achieved by counting the number of resistance spikes. The frequency of the resistance spikes decreases as the concentration decreases, as can be seen in Figure 1. The frequency of resistance spikes becomes quite low even at 4 μM (RL)4 present in the sample solution. It has been estimated for a nanopore sensor that if femtomolar concentrations were required then resistance spikes would be observed at the rate of one every 10 min.16 This low number of detected resistance spikes is not an issue of the sensitivity of the detector but one of mass transport of the target analyte to the sensor. Thus, the challenge of achieving ultralow detection limits is transferred from an issue related to the sensitivity of the transducer to one of mass transport. We will discuss this further, as well as how to overcome this challenge, in the next section. The strategy used in nanopore sensors to allow the detection of single molecules, that is also replicated in many ultrasensitive detectors, is to reduce the measurement volume down to very small amounts, femtoliter or less. Similar femtoliter measurement volumes are interrogated for single-molecule fluorescence, scanning probe microscope measurements,17,18 nanowire field effect transistor-based sensors,19,20 and surface-enhanced Raman spectroscopy (SERS).21−24 A consequence of reducing the measurement volume to the nanoscale is that just a few molecules in that volume equate to a significant concentration. For example, a single molecule in 1 fL gives a concentration of 2 nM. In nanopore sensors, the nanopore itself defines the measurement volume in which the molecule is confined. An alternative is to confine the molecules to be measured to the surface of a nanoscale object. Turning single plasmonic nanoparticles into individual detectors for single-nanoparticlebased assays has received significant attention recently.25 Sensors based on the surface plasmon resonance of individual nanoparticles are an excellent example of this. Around each nanoparticle, the surface plasmons decay within the surrounding medium, typically within less than 10−20 nm from the nanoparticle surface.26 As such, the sensing volume is a few zeptoliters (10−21 L) for a single plasmonic nanoparticle. The volume is so small that the presence of a single molecule or single-molecule adsorbates within this zone can change the local refractive index enough for a measurable shift in localized surface plasmon resonance (LSPR). The sensitivity of the LSPR makes
2. ASSAY SENSITIVITY There have been astonishing advances in the past few years with regard to more sensitive bioanalytical sensors. These advances are derived from either more sensitive transducers or new amplification schemes that generate sufficient signal from just a few analyte molecules. The nanoscale is pervasive in methods that provide more sensitive detectors and in amplification. 2.1. More Sensitive Transducers. Strategies for the development of more sensitive transduction methods for ultrasensitive sensors are exemplified by nanopore sensors.13,14 Nanopore sensors employ nanoscale pores in an insulating membrane to monitor the passage of molecules, nanoparticles, and even particle-like microbes through the pore. The sensing is achieved by measuring the conductivity between electrodes placed on each side of the insulating membrane. As the conductivity is determined by ion fluxes through the pores, the analyte is detected by a reduction in conductivity as the analyte temporarily blocks the nanopore as it passes through from one side of the membrane to the other. The selection of the pore diameter is determined by the size of target analyte and is typically nanoscale, so that when an analyte molecule translocates through the nanopore, there is a significant change in resistance. As such, nanopore sensors detect analytes one molecule at a time. This single-entity sensitivity, in principle, gives a nanopore sensor the ultimate resolution of a single molecule. The strength and weakness of nanopore sensors for quantitative sensing are demonstrated in the very early publications on this topic. For example, Sanchez-Quesada et al.15 monitored the current from a single α-hemolysin (α-HL) pore for the detection of a cyclic peptide, (RL)4 (see Figure 1). This single-nanopore sensor detected (RL)4 in the micromolar 1163
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Journal of the American Chemical Society detection platforms using single plasmonic nanoparticles a promising technology for the next generation of highthroughput single-molecule sensing that can detect analytes in ultralow concentrations. Single-nanoparticle LSPR has been successfully demonstrated for the detection of single-molecule binding events on bare and receptor-functionalized gold nanorods.26−28 It has been revealed recently that even single metal ions (e.g., zinc or mercury ions) can be monitored by their interaction with single gold nanorods.29 One key advantage of nanoscale detectors such as plasmonic nanoparticles for ultralow detection limit biosensors over nanopore sensors where single molecules are detected sequentially is the potential for massively parallel detection with an optical readout from many single plasmonic nanoparticles (Figure 2).30 In essence this converts the measurement
Figure 3. Current vs time recording for a monoclonal anti-prostate specific antigen (PSA) antibody-modified p-type silicon nanowire, tracked by alternating the delivery of protein and pure buffered solutions: (1) 9 pg/mL PSA, (2) 0.9 pg/mL PSA, 0.9 pg/mL PSA + 10 μg/mL bovine serum albumin, (4) 10 μg/mL bovine serum albumin, and (5) 9 pg/mL PSA. The important feature to note is the reversibility of antigen binding, as evident from the dissociation of the PSA from the nanowires with each switch from protein solution to buffer. Adapted with permission from ref 20. Copyright 2005 Nature Publishing Group.
Such devices can detect proteins down to 0.49 fM and DNA down to 0.1 fM under optimized conditions.36 The ability to observe the reversibility of binding reactions does mean that the lifetimes of molecular binding pairs must be considered. In conventional sensors, with the high-affinity constants of biorecognition molecules, a few molecules dissociating from the surface will make no appreciable change in the response of the sensor. This only becomes a challenge with ultrasensitive sensors, where dissociation of a few molecules or a single molecule can be determined, as in the example in Figure 3. If one considers typical affinity constants of biorecognition species being in the micromolar to nanomolar range, the lifetime of molecular binding pairs is between 1 s and 3 min.11,37 Hence, if the measurement time is longer than this, the analytical result may be compromised. Again, nanomaterials provide a solution to this challenge by loading many biorecognition molecules onto a single nanoparticle. A nanoparticle can have many biorecognition species on its surface. This multivalency of the biorecognition species on a nanoparticle gives a surface effect where there is an apparent decrease in the off rates for biomolecular binding pairs. This apparent decrease in the rate of dissociation of the biomolecular binding pair can be attributed to an analyte dissociating but then binding to another biorecognition species on the nanoparticle before it diffuses away.7,38,39 The apparent decrease in off rate also serves the purpose of increasing the apparent binding constant. This effectively pushes the concentration range at which the analyte can be detected to lower concentrations.40 Thus, multivalent binding by loading nanoparticles with many biorecognition species is also used in many amplification strategies. 2.2. Signal Amplification Systems for Conventional Assays. The alternative to developing increasingly more sensitive transducers to achieve ultralow detection limits is either to amplify the signal from a small number of analyte molecules or to amplify the number of analyte molecules.4 Both strategies have been exploited for decades in sensing. Amplification of the signal is exemplified by enzymatic labeling and the polymerase chain reaction (PCR). Microfluidics and microfabrication have seen impressive advances in signal
Figure 2. Monitoring of many individual nanorod-based plasmonic sensors for single-molecule plasmon sensing in a wide field format. A glass coverslip is modified with 150−250 gold nanorods per field of view such that each diffract limited spot is one gold nanorod (circled), anti-biotin antibody is flowed across the surface, and the light scattered by each gold nanorod is monitored using a microscope. Each step in the intensity−time trace is attributed to a single antibody binding. Adapted with permission from ref 30. Copyright 2015 American Chemical Society.
from a near-field measurement to a wide-field measurement.11 One of the impediments to such a strategy has been the speed of the readout of the optical signatures of each nanoparticle. Improvements in CMOS (complementary metal oxide semiconductor) cameras now make it possible to rapidly read out the optical properties of thousands of nanoparticles in less than 1 s using a color analysis approach.31,32 Further advances in multiparallel nanoscale detectors can be achieved using microfluidic transport of the sample to the surface of the nanoparticles.33 One of the interesting aspects of decreasing the measurement volumes to the nanoscale, where there may be only a few molecules in the measurement volume, is that stochastic behavior may be observed. In a conventional sensor, with a high-affinity biorecognition species, a few analyte molecules dissociating from their surface-bound affinity ligands may not be detected. However, with these ultrasensitive transducers, with a low number of molecules in the sensing volume, the unbinding of a few molecules, or even a single molecule, can be observed. This is ably demonstrated by the silicon nanowire field-effect transistors (NW-FETs) described by Lieber and co-workers,19,20,34 where, as the analyte flowed past the nanowires, proteins were observed to bind and then unbind from the surface with a concomitant change in the resistance of the NW-FET (Figure 3). The reversibility of this binding was attributed to the nanowire possessing fewer than 100 antibodies on each wire.35 1164
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Journal of the American Chemical Society amplification, for example, with PCR-on-a chip devices.41 Nanomaterials, however, have provided a whole new raft of amplification options which will be discussed here. One logical way that nanomaterials can enhance the sensitivity and detection limit of bioassays is through their exceptional optical properties. For example, the use of very bright luminescent lanthanide-doped nanoparticles is shown by Xu et al.,42 where simply changing the label to an ultrabright, sub-5-nm lanthanide-doped lutetium oxyfluoride nanoparticle in a commercial kit gave a 200-fold improvement in detection limit of prostate-specific antigen (PSA) to 15.2 fM. Similarly, Chang et al.43 developed ultrabright SERS labels composed of many silver nanoparticles coated with a SERS-active dye that are assembled on a silica core and then encapsulated in a silica shell. In effect, the SERS label uses the idea of the label having a high mass loading of the species that is being detected. Using this strategy, a conventional immunoassay with detection limits for PSA of 3.4 fM has been achieved. Rather than use a very bright label, changing the optical properties of the labels has also given ultralow detection limits. An elegant example is described by Stevens and co-workers,44 who used a conventional sandwich immunoassay format with polyclonal antibodies conjugated to gold nanostars for capturing the PSA. Monoclonal anti-PSA antibodies followed by secondary antibodies with conjugated glucose oxidase were used to generate the signal. The addition of glucose resulted in the glucose oxidase producing hydrogen peroxide, which then reduced silver ions in solution into silver that deposited on the nanostars. The result was a change in the optical properties in the form of a shift in the plasmonic signal. Detection limits for PSA of 40 zM were reported. The amazing aspect about this core idea is that De La Rica et al.45 later showed close to singlemolecule immunoassays, maybe down to five molecules, in microtiter plate wells.46
interface more rapidly and methods that extend the length scale of nanosensors further out into the sample solution.12 In all interfacial sensors, eventually diffusion is the mode of mass transport by which the analyte finally interacts with the sensing surface. The time constant for diffusional processes has a squared dependency on the distance that the diffusing species must travel. The obvious way, therefore, to increase the flux of analyte to a sensing interface is to reduce the thickness of the diffusion layer via convection. This has been the approach used in sensing for decades.48 A simple theoretical treatment to estimate the impact of convection, the size of the sensor, and the kinetics of the biorecognition binding reaction on the time constants of nanoscale sensor has been developed by Squires et al.49 It is also acknowledged in their paper that there are several examples in the literature where the time constants of the response of some nanosensors can be orders of magnitude quicker than that predicted by simple physicochemical models due to as yet undetermined factors.49 Strategies to improve the response time to reduce the detection limits of nanopore sensors have employed migration to increase the flux to the nanoscale sensor rather than convection. For instance, to increase the flux of DNA to the nanopore, Meller et al.50 employed asymmetric electrolyte solutions on each side of the membrane containing the nanopore to enhance the electric field that projected out from the entrance of the nanopore to attract the DNA to the pore (Figure 4). Using such a strategy, a 30-fold increase in translocation events was observed for 3.8 pM for the detection of DNA. Similarly, Freedman et al.51 used gold-coated glass nanopipets and dielectrophoretic trapping to increase the flux and give a translocation rate of ∼5.25 events per second,
3. RESPONSE TIME The discussion in this Perspective thus far has focused entirely on maximizing the signal from a low number of analyte molecules. The observant reader will notice that some of the strategies employed to reach ultralow detection limits using labels are in fact quite old. The same can be said about transducers that are capable of detecting as low as single molecules. For example, the nanopore sensor depicted in Figure 1 dates back to the year 2000.15 So, clearly, sensitivity is not the only challenge to overcome to develop useful ultrasensitive sensors. From a purely practical perspective, sensors are helpful only if accurate quantitation of an analyte is accomplished within a reasonable time frame.47 Sensor response time is a huge challenge for ultralow detection limit sensors. The challenge with regard to response time is one of mass transport, that is, getting the analyte to the sensor.12 As discussed above, with a nanopore, if the analyte concentration was in the ultralow concentration range of 1 fM, then the time between resistive spikes would be of the order of 10 min.16 In effect, there is a mismatch in length scales. The sensors are designed at the nanoscale with the biomolecular interactions occurring at the 1−10 nm scale. The static boundary layer adjacent to a sensing surface where mass transport occurs by diffusion is of the order of 10−100 μm, with the analytical sample size being microliters to milliliters and sometimes even liters in volume.12 There have been a number of elegant approaches to solving this mismatch of length scales, which can be subdivided into methods that bring the analyte to the sensing
Figure 4. (a) Schematic illustration of creating an asymmetric salt gradient across the nanopore, where there is a significantly lower salt concentration on the cis side of the nanopore where DNA is present. (b) Current−time traces at a 3.5 nm nanopore for monitoring translocation of 3.8 nM 400 bp DNA molecules under asymmetric salt concentrations. As can be seen, when there are lower concentrations of KCl on the cis side where the DNA is present compared with the trans side, there are higher rates of capture events and hence resistive pulses. Adapted with permission from ref 50. Copyright 2010 Nature Publishing Group. 1165
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The most common method for extending the sensing interface out into the sample solution uses magnetic nanoparticles. By dispersing the magnetic nanoparticles throughout the sample solution to capture the analyte, followed by using a magnet to collect the nanoparticles for measurement, dramatic decreases in detection limit are achieved. The detection limit is lowered because the majority of the analyte in the sample is collected by the nanoparticles. This is distinct from conventional sensors, where only a small proportion of the analyte is captured. As dispersing the nanoparticles throughout the sample reduces the distance over which species must diffuse, the response time is also decreased.56 In effect, this approach switches the sensing paradigm from making the analyte find the sensor to making the sensor find the analyte. There are a myriad of examples of this strategy.57−59 In most cases, magnetic nanoparticles are used as a method of preconcentrating the analyte before bringing the analyte to the sensing interface. An example of this method, combined with the use of ultrasensitive detectors, was reported by Wanunu et al.,60 using nanopores for the detection of microRNA. Magnetic microparticles are used to scavenge the microRNA using the p19 protein conjugated to magnetic microparticles. P19 binds 21−23-bp dsRNA. Probe DNA is bound to the miRNA of interest, which is then captured by the p19 beads to preconcentrate it. The miRNA−probe RNA duplex is then released from the magnetic microparticle and measured through the nanopore. Even greater performance can be achieved by incorporating the magnetic nanoparticles into the signal generation as well as the collection of analyte. This was demonstrated by the Gooding group in their “dispersible electrode” concept.56 Dispersible electrodes are gold-coated magnetic nanoparticles (Au@ MNPs) that are modified with biorecognition molecules to facilitate selective detection of a species of interest. The Au@ MNPs are dispersed in the sample matrix, where they capture the analyte. Applications of a magnetic field rapidly bring Au@ MNPs back to a macroscale electrode, where quantitation of captured analytes can be realized electrochemically. The magnetic nanoparticles were coated in gold so that they were conducting, such that they formed an electrode when brought back to the macroelectrode. The “dispersible electrode” moniker is used as the idea is that an electrochemical sensor is broken up into nanoparticles, dispersed through the sample solution to reduce response times by reducing diffusional pathlengths and to increase sensitivity by collecting more of the analyte, and then reassembled at the macroelectrode by the application of a magnetic field. The concept was first demonstrated for the detection of metal ions56 and has since been reported for detecting proteins,61 small organic molecules,62 and microRNA.63 The benefits of this strategy are shown for the detection of PSA, where the detection limits is 3.0 fM with a response time of 4 min, compared with a 2500 times higher detection limit on planar surfaces and a response time of over 30 min. Further, for the detection of microRNA, an amplification system gave detection limits for miR-21 of 10 aM in blood.63 The combination of magnetic nanoparticles to collect the analyte and labels that amplify the signal has been utilized by Walt and Rissin64 to give ultralow detection limits based on single-molecule measurements (Figure 6). The single-molecule assays are referred to as SiMoA and are the basis of the company Quanterix. This digital enzyme-linked immunosorbent assay (ELISA) technique has been successfully demonstrated for quantitation of a variety of small amounts of biological analytes from samples such as single-cell extracts, serum, and urine.
enabling a decrease in the limit for DNA detection down to 5 fM. Further, Schibel and Ervin52 reported that pressure differences can also be used to enhance the flux of analyte to an antibodymodified glass capillary nanopore. This strategy decreased the detection limit by 6 orders of magnitude compared to the limit when the pressure differential was absent. The effectiveness of extending the reach of the sensing interface further into the sample solution to reduce response times and detection limits was powerfully demonstrated by Kelley et al.53,54 using nanostructuring of electrodes via electrodeposition of palladium onto an electrode of 500 nm. The electrodes were nanostructured by varying the electrodeposition conditions such that the fractal dimensions of an electrodeposited metal could be controlled while the geometric area the plated electrodes covered remained similar. The three electrodes shown in Figure 5 were modified with a probe strand
Figure 5. (a) Scanning electron microscope images for three fabricated Pd electrodes with different levels of nanostructuring. (b) Comparison of detection limit and dynamic range for biosensors based on electrodes of these three types of structures: electrocatalytic nucleic acid detection at nanostructured microelectrodes. Adapted with permission from ref 53. Copyright 2009 Nature Publishing Group.
of DNA for the detection of a complementary target. The detection limit of the otherwise identical sensors decreased from 100 fM to 1 fM to 10 aM as the structure was changed from smooth to moderate nanostructuring to fine nanostructuring. The incredible improvement in performance is not just due to the approach of extending the length scale of the sensing interface. In later work,55 the same authors also showed that the nanostructuring influences the accessibility and density of the probe DNA on the sensing surface, which promotes hybridization of the target DNA and hence also contributes to the lowering of the detection limit. This point will be returned to in the next section. 1166
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sufficiently high affinity for the analyte of interest to give a stable signal and the presence of nonspecific adsorption giving a background signal. We will discuss solutions to each of these in turn. With ultralow detection limit sensors that have nanoscale dimensions, the number of analyte molecules that bind to the biorecognition surface can be of the order of a few hundred down to single molecules.11 As discussed for Lieber’s work in Figure 3, a consequence of such small numbers of biomolecular interactions is that the reversibility of binding, e.g., between an antigen and an antibody, becomes observable. Such reversible interactions are not typically observable in conventional sensors, where millions of molecular interaction are generating the signal.35 The discussion above on the lifetime of biomolecular complexes could then make one wonder how single-molecule sensors, such as that shown by Rissin et al.,64 are possible. A common theme in many ultralow detection limit sensors is that nano- or microparticles modified with biorecognition molecules are used. In the SiMoA technology, it is estimated that there are 274 000 antibodies per bead.7 A consequence of having so many surface-bound antibodies on a single bead is that the effective off rate for the biomolecular complex is reduced compared with that of antibodies in solution. This is attributed to both rebinding effects on surfaces with so many adjacent antibodies and possible multivalent effects.7 The influence of surface receptor density on the rebinding of species in solution has been treated theoretically by Thompson and co-workers for planar surfaces.65 In cases where the amount of analyte is very low, the theory shows that the likelihood of prompt rebinding increases with increased association rates and the density of surface receptors but decreases with a square root dependence on increases in dissociation rates and the diffusion coefficient. The notion of length scales is also important in rebinding. If the distance a molecule can diffuse in solution during the average time the molecule is surface bound is less than length described by the receptor surface density divided by the concentration of analyte in solution, then rebinding is favored. From the perspective of assay design for ultralow detection limit sensors, this simply means it is favorable to have more surface-bound receptors. The assay format also helps, of course. As discussed by Rissin et al.,7 the common receptors in bioassays, such as protein antibodies, have a disassociation constant in the nanomolar range. Therefore, they exhibit a 103-fold cross-reactivity over nonspecific molecules. However, a sandwich structure constructed from a capturing probe, a labeling probe binding to a different site, and a signaling reporter is an excellent assay format. This is because overall this construction provides a capability to discriminate specific and nonspecific molecules of larger than 1011-fold difference in concentration. Therefore, it becomes realistic to detect an analyte of femtomolar or less when outnumbered by nonspecific molecules of millimolar concentration. The other key issue is nonspecific adsorption, which might give a false positive signal. The standard approach to reducing nonspecific adsorption has been to modify the sensing interface with not just the biorecognition species but also surface chemistry to resist nonspecific adsorption of proteins and cells such as poly(ethylene glycol) or zwitterionic species.66 Such strategies have been extensively explored and optimized but seldom achieve >99% reduction in nonspecific adsorption. As such, this is not sufficient for ultralow detection limit sensors. Many ultralow detection limit sensors use such low fouling
Figure 6. Digital ELISA based on arrays of femtoliter-sized wells for single-molecule detection of serum proteins at sub-femtomolar concentrations. (a) An antibody-modified magnetic bead collects the analyte of interest. A secondary biotinylated antibody is then added, followed by streptavidin conjugated to β-galactosidase. (b) The magnetic beads are captured in wells small enough that only a single bead can go into each well. Enzyme substrate is added which, upon reaction with β-galactosidase, gives a fluorescent product. Only wells that exhibit fluorescence have a bead that has captured an analyte. (c) Scanning electron micrograph of the wells with loaded beads. (d) Fluorescence image of the well array, with the wells where an analyte is captured exhibiting fluorescence. Adapted with permission from ref 64. Copyright 2010 Nature Publishing Group.
SiMoA uses antibody-modified paramagnetic microparticles to capture the analyte. The microparticles are collected, washed to remove nonspecific molecules bound on their surfaces, incubated with detection antibodies and streptavidin-labeled β-galactosidase enzyme conjugates, and finally loaded into arrays of femtoliter-sized wells (∼40 fL). The volume of the wells ensures that each well can have only a single microparticle. Fluorescent products are generated by the enzyme label when an analyte exists on a microwell. Quantification of the number of analyte molecules is realized by counting the number of fluorescent microwells. As the number of microparticles is in considerable excess relative to the number of analyte molecules, each well that fluoresces contains a single molecule. The SiMoA system has been shown to be able to detect as low as 0.4 fM PSA from patient sera.
4. SELECTIVITY The discussion thus far has really focused on generating the analytical signal through either enhancing signal from an analyte or actually getting the analyte to the sensor to generate that signal. Ultralow detection limits, however, exacerbate wellknown challenges in sensing related to the selectivity of the sensor. This leads to the question: Is an ultralow detection limit sensor for a given analyte even practical? What this question pertains to is the selectivity of the biorecognition species for the analyte of interest in the actual clinical sample. The detection of ultralow amounts of a species in a complex sample really challenges the selectivity of a biomolecule, simply because, if a molecule is found at a level of 1 molecule in 1000 of a cross reactant and the biorecognition species has a selectivity of only 99%, then the sensor will not be selective enough.47 If the analyte is ultrarare relative to any cross reactants, then this scenario will be a major issue. Other challenges exacerbated by ultralow detection limits include whether the biomolecules have 1167
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times. Similarly, strategies to increase the signals with ultrabright labels report similar detection limits of 15 fM43 or even lower (to 0.3 fM using labels with high mass loading).57 Combining strategies to solve the issues of mass transport, labels with high mass loading of the species detected, and limiting nonspecific effects sees another 2 orders of magnitude drop in detection limit to 30 aM68 and even the ability to gain appreciable analytical signal from single PSA molecules.64 Thus, the success of addressing these three challenges is apparent. It is worth asking: What is behind the rapid advances in this field? Underpinning the majority, if not all, of these solutions is either the use of nanomaterials or the performance of measurements at the nanoscale, or in many cases both. Consider the triple challenge in ultralow detection limit sensors of more sensitive detectors, getting the analyte to the sensor, and achieving selectivity. The most common approaches to developing detectors with sufficient sensitivity involve either restricting the measured volume to the nanoscale, such that the concentration within the measurement volume is appreciable, or using nanoparticles as labels that generate far greater signal than would otherwise be possible. The solution to the mass transport challenge is to address the length scale mismatch by using nanomaterials that extend further into solution to limit the diffusional pathlengths required and hence reduce the response time constant of the sensor. The solutions to ensuring the sensor is sufficiently selective revolve around using nanoparticles with large numbers of recognition species to bias the biorecognition equilibria toward molecular binding pairs and using magnetic nanoparticles to spatially and temporally separate the capture of the analyte from complex biological samples. If there is a hero of ultrasensitive biosensors and bioassays, it is the magnetic nanoparticle. If one considers the examples discussed above, many of them relate to magnetic nanoparticles. The obvious assistance they provide is to preconcentrate the analyte and bring it to the detector. The high loading of recognition species on the nanoparticle facilitates binding.7 Magnetic nanoparticles also allow the measurement to be performed in solutions free of interfering substances after the collection of the analyte from complex samples, as mentioned in the previous paragraph. They can also be the way in which the signal is generated.57,63,68 We feel these developments in ultrasensitive sensors are exceedingly important. As mentioned in the Introduction, there are many unmet commercial applications for sensors that require the ability to detect ultralow concentrations. Similarly, being able to detect lower amounts of analyte means that sensors can be developed for fluids that are easier to access but have lower concentrations of analyte than existing sensors can measure.4 Furthermore, as was discussed by Mirkin and coworkers,8 more sensitive sensors can provide new fundamental insights into diseases. This last point is true for many existing sensors as well. In our view, with the emphasis of many in the field on developing commercial sensors, using sensors for biological discovery has long been a missed opportunity for sensing science. So what does the future hold for this field? Two areas where there is major focus involve the development of single-molecule sensors and the requirement to go even lower in concentration. The development of sensors that can measure single-molecule interactions, also referred to as digital assays, was discussed here with regard to the nanopore devices and the SiMoA technology.69 Single-molecule sensors can also be deceptively simple, as demonstrated by the slipchip devices described by
layers combined with other strategies to further limit nonspecific effects. The most common strategy used in this field, and the one used in most of the examples above in which measurements were performed in complex biological samples, is to perform the measurement in a solution different to that used for the analyte collection. This is the second major reason, after being able to preconcentrate analytes, that magnetic nanoparticles are so commonly used. The ability provided by magnetic nanoparticles to easily spatially and temporally separate the sample collection from the measurement means that complex samples can be analyzed. The low apparent off rates discussed above when many biorecognition species are immobilized on the magnetic nanoparticles further facilitate the effectiveness of this approach. There are, however, other strategies that can be employed. These mostly revolve around differentiating specific from nonspecific signals. For example, with refractive index sensors, nonspecific adsorption is a major issue and typically causes the same red shift in the optical signature as the analyte of interest. However, Kilian et al.67 have configured a porous silicon sensor for the detection of matrix metalloprotease levels directly from live cells where the enzyme signal causes a blue shift in the reflectance peak while the nonspecific effects cause a red shift. The blue shift is achieved when the enzyme reaction removes an enzyme-degradable polymer, hence reducing the average refractive index of the photonic crystal, while nonspecific adsorption adds material. This approach does not remove the problem, as nonspecific protein adsorption still causes a change in signal, but it does point to an effective strategy. Zheng et al.24 have also shown that, with a single-molecule sensor based on surface-enhanced Raman spectroscopy, at the single-molecule level the magnitude of the diagnostic Raman peak can be used to differentiate analyte specifically adsorbed into the Raman hotspots as distinct from nonspecifically adsorbed Raman dye. This last example points to the idea of using the type of interaction with the sensing surface of specific analyte, versus nonspecific adsorbed species, as a way to enhance sensor selectivity when the sensor operates at the single-molecule level.
5. CONCLUSIONS AND FUTURE PERSPECTIVES Herein we have outlined advances in bioassays and biosensors that can be classed as ultrasensitive, as they have sufficient sensitivity and low background to allow sub-picomolar detection limits. There have been many incredible advances in the field over the past decade or so. These advances have been underpinned by a variety of solutions to the triple challenge of finding more sensitive detectors, getting the analyte to the sensor, and achieving selectivity. The fact that there are many stunning developments in this field where prostate-specific antigen was used as the model analyte is fortuitous. This allowed us to choose examples of technologies presented herein used for the detection of PSA over many other worthy examples that detected different analytes, so as to provide a means by which the reader could evaluate the effectiveness of some of these strategies toward detecting the same molecule. In our own paper on dispersible electrodes for detecting PSA, a comparison was made between the performance of the ultralow detection limit dispersible electrodes and the same sensor prepared in exactly the same way on a conventional planar electrode.61 The dispersible electrodes address the issue of mass transport with ultralow detection limit sensors. This advance alone lowered the detection limit from 7.5 pM on the planar sensor to 3.0 fM with the dispersible electrodes, a decrease in detection limit of 2500 1168
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Journal of the American Chemical Society Ismagilov and colleagues.70−72 The development of quantitative sensors based on measuring many single-molecule interactions is not just beneficial for ultrasensitive sensors;11molecular counting obviates the need for calibration,7,64 the bane of existence of most analytical methods. Furthermore, the ability to monitor the heterogeneity in how single molecules interact with a sensor surface would be a powerful solution to the challenge of nonspecific signals24 or would allow one sensing interface to measure many different biomarkers in a unique way. Turning our attention to the challenge of sensors with even lower detection limits, the emphasis of the technologies presented thus far has been on the detection of protein and nucleic biomarkers. Even lower detection limits are needed for the detection of circulating tumor cells (CTCs) as prognostic and diagnostic biomarkers for cancer and for the detection of pathogens in clinical and environmental applications. With CTCs, one cell in 10 mL can be important, while for pathogens it could be one microbial species per liter in water quality analysis. These volumes add an even greater challenge for sensor technology, as such low concentrations mean that one must be sure that the sample for analysis, to be representative of the bulk samples, must have a volume several times greater. Such large sample volumes can be a challenge even for microfluidics, although the centrifugal CTC separation described by Lim and co-workers addresses the volume challenge effectively.73,74 The CTC challenge is particularly interesting, as this field really started with a focus on enumeration of the cells.75 However, in more recent times, the understanding that the definition of a CTC based on a surface biomarker, or cell size, alone is not adequate76 has led to technologies that can not only capture and count rare cells but also provide information on their heterogeneity.77,78 The quantification of heterogeneity in species that are as rare as one cell in 1 mL is an incredible challenge that requires many of the solutions discussed above. For example, the stunning solution reported by Kelley et al.78 uses magnetic nanoparticles to preconcentrate cells with phenotypic markers, where an advanced microfluidics device separates and isolates rare cells of a given phenotype depending on how many magnetic nanoparticles are bound to each cell phenotype. To finish, this field is moving very quickly. Biosensors are pushing the limits of how few entities can be quantified reliably and creating a need for such devices at the same time. Being able to explore the heterogeneity in rare biological species is just one such frontier these sensors can address. Rapidly detecting even rarer species, such as very rare pathogens, is another. It is the capability to measure very small amounts of analyte which is ensuring a bright future for sensing and sensing research.
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Science and Technology (CE140100036) and an Australian Laureate Fellowship (FL150100060).
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AUTHOR INFORMATION
Corresponding Author
*
[email protected] ORCID
Yanfang Wu: 0000-0003-4201-2061 Richard D. Tilley: 0000-0003-2097-063X J. Justin Gooding: 0000-0002-5398-0597 Notes
The authors declare no competing financial interest.
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ACKNOWLEDGMENTS We acknowledge funding from the Australian Research Council for the ARC Centre of Excellence in Convergent Bio-Nano 1169
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