Review

© Future Drugs Ltd. All rights reserved. ISSN 1473-7159 527

CONTENTS

Optical properties of noble metal nanoparticles

Using nanoparticles for biomolecule sensing

LSPR sensing of large proteins & antibodies on arrays of nanoparticles

Reversibility of the sensor

Selectivity of the sensor

Detection of disease markers

Miniaturization of the sensor

Comparison with an industry standard: SPR sensors

Outlook & conclusions

Expert opinion

Five-year view

Key issues

References

Affiliations

www.future-drugs.com

Preliminary studies and potential applications of localized surface plasmon resonance spectroscopy in medical diagnosticsAmanda J Haes and Richard P Van Duyne†

Miniature optical sensors that specifically identify low concentrations of environmental and biological substances are in high demand. Currently, there is no optical sensor that provides identification of the aforementioned species without amplification techniques at naturally occurring concentrations. Recently, it has been demonstrated that triangular silver nanoparticles have remarkable optical properties and that their enhanced sensitivity to their nanoenvironment has been used to develop a new class of optical sensors using localized surface plasmon resonance spectroscopy. The examination of both model and nonmodel biological assays using localized surface plasmon resonance spectroscopy will be presented in this review. It will be demonstrated that the use of a localized surface plasmon resonance nanosensor rivals the sensitivity and selectivity of, and provides a low-cost alternative to, commercially available sensors.

Expert Rev. Mol. Diagn. 4(4), 527–537 (2004)

†Author for correspondenceNorthwestern University, Department of Chemistry, 2145 Sheridan Road, Evanston, IL 60208 –3113, USATel.: +1 847 491 3516 Fax: +1 847 491 7713 vanduyne@chem.northwestern.edu

KEYWORDS:biosensing, disease diagnosis, localized surface plasmon resonance nanosensor, single nanoparticle spectroscopy

Advancements in technology due to nanoscalephenomena will become optimized when thechemical and physical properties of materialsare more thoroughly understood. Relevant tothis work, the potential to develop highly sen-sitive and specific sensors for biological targetsmotivates a portion of the research in thisfield. Previously, results have been presentedbased on the nanoscale limit of surface plas-mon resonance (SPR) sensors. This novel,nanoscale development, termed the localizedsurface plasmon resonance (LSPR) nanosen-sor, is a refractive index-based sensing devicethat relies on the extraordinary optical prop-erties of noble (e.g., Ag, Au and Cu) metalnanoparticles [1–5].

Optical properties of noble metal nanoparticlesNoble metal nanoparticles exhibit a strongultraviolet (UV)-visible (vis) absorption bandthat is not present in the spectrum of the bulkmetal [6–12]. This absorption band results whenthe incident photon frequency is resonant withthe collective oscillation of the conduction

electrons and is known as the LSPR. LSPRexcitation results in:

• Wavelength selective absorption withextremely large molar extinction coefficients(~3 × 1011 M-1 cm-1)

[13]

• Resonant Rayleigh scattering [14,15] with an efficiency equivalent to that of 106 fluorophors [16]

• The enhanced local electromagnetic fieldsnear the surface of the nanoparticle, whichare responsible for the intense signalsobserved in all surface-enhanced spectroscopies [5,14,17–19]

The simplest theoretical approach availablefor modeling the optical properties of nano-particles is the Mie theory estimation of theextinction of a metallic sphere in the longwavelength, electrostatic dipole limit. This isexplained in EQUATION 1 [20]:

E λ( )24πNAa3εm

3 2⁄

λ 10( )ln⋅--------------------------------------

εi

εr χεm+( )2 ε2i+

-----------------------------------------×=

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528 Expert Rev. Mol. Diagn. 4(4), (2004)

• E(λ) = extinction (viz., sum of absorption and scattering)

• NA = areal density of nanoparticles

• A = radius of the metallic nanosphere• εm = dielectric constant of the medium surrounding the

metallic nanosphere (assumed to be a positive, real numberand wavelength independent)

• λ = wavelength of the absorbing radiation

• εi = imaginary portion of the metallic nanoparticle’s dielectric function

• εr = real portion of the metallic nanoparticle’s dielectric function

• χ describes the polarization factor that corresponds to the aspect ratio of the nanoparticle (equal to 2 for a sphere)

It is abundantly clear that the LSPR spectrum of an iso-lated metallic nanosphere embedded in an external dielectricmedium will depend on the nanoparticle radius A, the nano-particle material (εi and εr) and the nanoenvironment’sdielectric constant (εm). Furthermore, when the nanoparti-cles are not spherical, as is always the case in real samples, theextinction spectrum will depend on the nanoparticle’s in-plane diameter, out-of-plane height and shape (χ). The valuesfor χ increase from 2 (for a sphere) up to and beyond valuesof 17 for a 5:1 aspect ratio nanoparticle. In addition, many ofthe samples considered in this work contain an ensemble ofnanoparticles that are supported on a substrate. Thus, theLSPR will also depend on interparticle spacing and substratedielectric constant.

Using nanoparticles for biomolecule sensingIt is apparent from EQUATION 1 that the location of the extinc-tion maximum of noble metal nanoparticles is highly depen-dent on the dielectric properties of the surrounding environ-ment. Wavelength shifts in the extinction maximum ofnanoparticles can be used to detect molecule-induced changessurrounding the nanoparticle. As a result, there are at least fourdifferent nanoparticle-based sensing mechanisms that enablethe transduction of macromolecular or chemical-bindingevents into optical signals based on changes in the LSPRextinction or scattering intensity shifts in LSPR λmax, or both.These mechanisms are:

• Resonant Rayleigh scattering from nanoparticle labels in amanner analogous to fluorescent dye labels [15,16,21–28]

• Nanoparticle aggregation [29–33]

• Charge-transfer interactions at nanoparticle surfaces [1,20,34–37]

• Local refractive index changes [1–5,38–46]

In this review, the refractive index change in the environ-ment near noble metal nanoparticles is demonstrated to be aneffective platform for highly sensitive detection techniques. Inorder to systematically study the LSPR response of noble metalnanoparticles to changes in their dielectric environment, atechnique that produces nanoparticles with size and shapemonodispersity is required. The chemical synthesis of noblemetal nanostructures is often employed for this purpose [47–52].

These approaches are used to prepare high concentrations of avariety of shapes and sizes of nanoparticles with varyingdegrees of monodispersity and tunable optical properties(FIGURE 1). For solution-phase LSPR-based sensing, signaltransduction depends on the sensitivity of the surface plasmonto interparticle coupling. When there are multiple particles insolution that support a localized surface plasmon and are inclose proximity (i.e., interparticle spacings less than the nano-particle diameter), they are able to interact electromagneticallythrough a dipole coupling mechanism. This broadens and redshifts the LSPR, a change easily detected using UV-vis spec-troscopy. Two methods of detection readily lend themselves tomonitoring these changes in the position of the LSPR: UV-visextinction (absorption plus scattering) and resonant Rayleighscattering spectroscopy.

200 400 600 800 1000Wavelength (nm)

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40 nm 40 nm

1 2

6

3

50 nm

33 nm

A

B

4 5

(1)

(4)

(2)(5)

(3)(6)

Figure 1. Tunable Ag and Au nanoparticle solutions fabricated via the citrate reduction of metal salts. (A) Corresponding transmission electron micrographs. (1) The solution consists of homogenous Au nanospheres (13 nm diameters). (2) The solution consists of inhomogenous Ag nanoparticles (nanospheres, trigonal prisms and polygon platelets). (3) The solution consists of Ag nanoparticles (nanospheres, trigonal prisms, and polygon platelets). (4) The solution consists of Ag nanoparticles (trigonal prisms and polygon platelets). (5) The solution consists of Ag nanoparticles (trigonal prisms with rounded tips and polygon platelets). (6) The solution is made up of inhomogenous oblong Ag nanoparticles. (B) Ultraviolet-visible extinction spectra of the corresponding solutions.

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Recently, several papers have been published on a gold nano-particle-based UV-vis technique for the detection of DNA.This colorimetric detection method is based on the change inabsorbance spectra (i.e., color) as particles are brought togetherby the hybridization of complementary DNA strands[30–33,53–55]. The limits of detection (LOD) reported are in therange of tens of femtomoles of target oligonucleotide. Thesenanoparticle aggregation assays represent a 100-fold increase insensitivity over conventional fluorescence-based assays [54].

An alternative approach to solution-phase nanoparticle sens-ing is to synthesize nanoparticles bound to substrates. In thisformat, signal transduction depends on changes in the nanopar-ticles’ dielectric environment induced by solvent or target mole-cules (not via nanoparticle coupling). One approach to fabricatethese nanoparticles is to use nanosphere lithography (NSL) [56],a powerful fabrication technique that inexpensively producesarrays of nanoparticles with controlled shape, size and interpar-ticle spacing. NSL begins with the self-assembly of size-mono-disperse nanospheres of diameter, D, to form a 2D colloidalcrystal deposition mask. The substrate is prepared so that thenanospheres freely diffuse until they reach their lowest energyconfiguration. This is achieved by chemically modifying thenanosphere surface with a negative charge that is electrostaticallyrepelled by the negatively charged substrate, such as mica orchemically treated glass. As the solvent (water) evaporates, capil-lary forces draw the nanospheres together and they crystallize ina hexagonally close-packed pattern on the substrate. As in allnaturally occurring crystals, nanosphere masks include a varietyof defects that arise as a result of nanosphere polydispersity, siterandomness, point defects (vacancies), line defects (slip disloca-tions) and polycrystalline domains. Typical defect-free domainsizes are in the 10–100 µm range. Following self-assembly of thenanosphere mask, a metal or other material is then deposited bythermal evaporation, electron beam deposition, or pulsed laserdeposition from a collimated source normal to the substratethrough the nanosphere mask to a controlled mass thickness, dm.Following metal deposition, the nanosphere mask is removed,typically by sonicating the entire sample in a solvent, leavingbehind surface-confined nanoparticles that have a triangularfootprint. The optical properties of these nanoparticles can beeasily tuned throughout the visible region of the spectrum bychanging the size or shape of the nanoparticles (FIGURE 2).

LSPR sensing of large proteins & antibodies on arrays of nanoparticlesUsing NSL, the authors have demonstrated that nanoscale chemo-sensing and biosensing could be realized through shifts in the LSPRextinction maximum (λmax) of these triangular silver nanoparticles[1,2,4,43]. Instead of being caused by the electromagnetic coupling ofthe nanoparticles, these wavelength shifts are caused by adsorbate-induced local refractive index changes in competition with charge-transfer interactions at the surface of nanoparticles. It should benoted that the signal transduction mechanism in this nanosensor isa reliably measured wavelength shift rather than an intensitychange as in many previously reported nanoparticle-based sensors.

As required for all new sensor development, a model systemmust be chosen to test the capabilities of the sensor with a well-characterized system. For that reason, the well-studiedbiotin/streptavidin system [2], with its extremely high bindingaffinity (Ka ~1013 M-1), and the antigen/antibody couple,biotin/antibiotin [4], were chosen to illustrate the attributes ofthese LSPR-based nanoscale affinity biosensors. For theseexperiments, macroscale UV-vis extinction spectra were col-lected in standard transmission geometry with unpolarizedlight using a miniature spectrometer.

In the streptavidin studies, NSL was used to create surface-confined triangular Ag nanoparticles supported on a glass sub-strate [2]. The Ag nanotriangles have in-plane widths ofapproximately 100 nm and out-of-plane heights of approxi-mately 51 nm as determined by atomic force microscopy. Toprepare the LSPR nanosensor for biosensing events, the Agnanotriangles are first functionalized with a self-assembledmonolayer (SAM) composed of 3:1 1-octanethiol:11-mercap-toundecanoic acid resulting in a surface coverage corresponding

200 400 600 800 1000

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Wavelength (nm)

(1) 432 nm (3) 623 nm(2) 552 nm

250 nm

2

250 nm

1

250 nm

3

A

B

C

1 2 3

Figure 2. Tunable Ag nanoparticle substrates fabricated using nanosphere lithography. (A) Photographs of nanoparticle substrates. (B) Atomic force microscopy images of nanoparticle substrates. (C) Ultraviolet-visible extinction spectra of Ag nanoparticle substrates. In all sections, Ag nanoparticle substrate D = 400 nm, dm = 50.0 nm. (1) After thermal in vacuum for 1 h at 600°C, (2) after thermal annealing in vacuum for 1 h at 300°C, and (3) as fabricated (no annealing).

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530 Expert Rev. Mol. Diagn. 4(4), (2004)

to 0.1 monolayer of carboxylate binding sites [57]. Since themaximum number of alkanethiol molecules per nanoparticle is60,000 [1], this results in an equivalent of approximately 6000carboxylate binding sites per nanoparticle. Next, an amine-ter-minated biotin was covalently attached to the carboxylategroups using 1-ethyl-3,3´(dimethylaminopropyl)carbodiimide(EDC). The number of resulting biotin sites is determined bythe yield of the EDC coupling reaction. Since this is likely to beapproximately 1–5% efficient [58], one expects there to be only60–300 biotin sites per nanoparticle at maximum coverages. Inthese studies, the λmax of the Ag nanoparticles were monitoredduring each surface functionalization step. An example of anassay for the detection of streptavidin follows. First, the LSPRλmax of the bare Ag nanoparticles were measured to be561.4 nm. To ensure a well-ordered SAM on the Ag nano-particles, the sample was incubated in the thiol solution for24 h. After careful rinsing and thorough drying with N2 gas,the LSPR λmax after modification with the mixed SAM wasmeasured to be 598.6 nm. The LSPR λmax shift correspondingto this surface functionalization step was a +38 nm wavelengthshift. Next, biotin was covalently attached via amide bond for-mation with a two-unit polyethylene glycol linker to carboxy-lated surface sites. The LSPR λmax after biotin attachment wasmeasured to be 609.6 nm, corresponding to an additional+11 nm shift. Given the magnitude of this shift and theexpected magnitude of the response (see EQUATION 2, vidainfra), this shift is confirmed to arise from an approximately5% submonolayer coverage of biotin. The LSPR nanosensor

has now been prepared for exposure to the target analyte. Toensure that molecular binding had saturated, approximately25 µl of the molecule sample was exposed to the nanoparticlesubstrate for 3 h [2]. Exposure to 100 nM streptavidin resultedin LSPR λmax = 636.6 nm, corresponding to a +27 nm shift fora full monolayer coverage of streptavidin (FIGURE 3). Repeatingthis experiment with exposure of the nanosensor surface toonly 1 pM streptavidin results in a marked decrease of theresponse to a small but reproducibly detected +4 nm shift.This concentration limit is as sensitive as other commerciallyavailable optical sensors.

An identical approach was taken to detect antibiotin anti-body (sigma). However, instead of exposing the LSPR sensorchip to streptavidin, it was exposed to antibiotin antibody. Thebinding affinity of antibiotin antibody to biotin is significantlylower (1.9 × 106 to 4.98 × 108 M-1) than that of biotin/strepta-vidin [59,60]. In this case, upon exposure to 700 nM antibiotinantibody, a wavelength shift of +42.6 nm was detected. Whenthe solution concentration was decreased to 700 pM antibiotinantibody, a small but reproducible +2.5 nm LSPR responseshift was observed.

Clearly, as the concentration of the solution decreases, theLSPR wavelength shift response decreases. This encouraged theauthors to determine the full response of the sensor over a wideconcentration range. For this reason, variable analyte concen-trations were exposed to a biotinylated LSPR chip to test thesensitivity of the system to different molecules. Specifically, theLSPR λmax shift, ∆R, versus [analyte] response curve was

0

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40

Rmax = 26.5 nm

LOD < 1 pM

Ka, surf = 1 × 1011 M-1

Streptavidin

10-16 10-15 10-14 10-13 10-12 10-11 10-10 10-9 10-8 10-7 10-6 10-5 10-4 10-3

LSP

R r

espo

nse

(nm

)

500 540 580 620 660 700

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Biotin Streptavidin

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Wavelength (nm)540 580 620 660

R max = 38.0 nm

LOD < 700 pM

Ka, surf = 4.5 × 107M-1

Antibiotin

Concentration (M)

A

B

Figure 3. The specific binding of streptavidin (A) and antibiotin antibody (B) to a biotinylated Ag nanobiosensor is shown in the response curves. All measurements were collected in a N2 environment. The solid line is the calculated value of the nanosensor's response. The left inset represents an LSPR nanosensor experiment in which biotinylated nanoparticles are exposed to 100 nM streptavidin. As shown, the LSPR of biotinylated Ag nanoparticles shifts to the red by 27 nm upon the specific linkage of a monolayer of streptavidin to its surface. The right inset represents an LSPR nanosensor experiment in which biotinylated nanoparticles are exposed to 7 mM antibiotin antibody. As shown, the LSPR of biotinylated Ag nanoparticles shifts to the red by 38 nm upon the specific linkage of a monolayer of antibiotin antibody to its surface.Ka,surf: Surface-confined thermodynamic binding constant; LSPR: Localized surface plasmon resonance; LOD: Limit of detection; Rmax: Saturation response.

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measured over the concentration range 1 × 10-15 M < [strepta-vidin] < 1 × 10-6 M and 7 × 10-10 M < [antibiotin] < 7 × 10-6

M (FIGURE 3) [2,4]. The lines seen in the figure are not a fit to thedata. Instead, the line was computed from a response model (acomplete analysis of this model is described in [4]). It was foundthat this response could be interpreted quantitatively in termsof a model involving:

• 1:1 binding of a ligand to a multivalent receptor with different sites but invariant affinities

• The assumption that only adsorbate-induced local refractiveindex changes were responsible for the operation of the LSPR nanosensor

The binding curve provides three important characteristicsregarding the system being studied. First, the mass and dimen-sions of the molecules affect the magnitude of the LSPR shiftresponse. Comparison of the data with theoretical expectationsyielded a saturation response of ∆Rmax = 26.5 nm for streptavi-din, a 60-kDa molecule, and 38.0 nm for antibiotin antibody, a150-kDa molecule. Clearly, a larger mass density at the surface ofthe nanoparticle results in a larger LSPR response. Next, the sur-face-confined thermodynamic binding constant Ka,surf can becalculated from the binding curve and is estimated to be1 × 1011 M-1 for streptavidin and 4.5 × 107 M-1 for antibiotinantibody. These numbers are directly correlated to the thirdimportant characteristic of the system, the LOD. The LOD isless than 1 pM for streptavidin and 100 pM for antibiotin anti-body. As predicted, the LOD of the nanobiosensor studied islower for systems with higher binding affinities, such as for thewell-studied biotin–streptavidin couple, and higher for systemswith lower binding affinities as seen in the antibiotin antibodysystem. Given this information and analysis, similar treatmentcan be made for virtually any ligand receptor system. It should benoted that the LOD of the system corresponds to the smallestreliable wavelength shift response induced by a given solutionconcentration. These real LODs are often translated to the com-monly reported surface coverage in terms of molecules. For theLSPR nanosensor, the surface coverage detection corresponds toapproximately 25 streptavidin molecules/nanoparticle. In previ-ously described experiments, approximately 3 × 108 nanoparticleswere probed (nanoparticle density approximately 1 × 1010 nano-particles/cm2 with a spot size with a 1-mm radius) and this corre-sponds to the detection of 75 × 108 streptavidin molecules. Aclear method to decrease the number of molecules detectedwould be to decrease the number of nanoparticles probed. Thishas been recently demonstrated and is discussed later.

Reversibility of the sensor In order for LSPR nanobiosensors to fulfill their mandate,they must be biocompatible and work under physiologicalconditions. Some binding interactions, such as poly-L-lysineto a negatively charged surface, can interact reversibly, whileother couples with higher surface binding affinities interactirreversibly. A commercially viable nanobiosensor should beentirely reusable. In the case of this study, this means that the

analyte detection must be entirely removable, rendering thesensor reusable. Reusability has a large impact on the costeffectiveness and the simplicity of biosensor use. The reversi-bility of the LSPR nanobiosensor has been demonstrated byexposing an antibiotin antibody functionalized chip with anexcess concentration of biotin [4]. In less than 30 s, the sen-sor’s capability to be reused to detect antibiotin antibody hadbeen fully regenerated. While this experiment is more diffi-cult if the affinity between the ligand and receptor is stronger(as is the case with biotin/streptavidin), these results clearlyindicate that an LSPR sensor can be reused multiple times forantigen/antibody interactions.

Selectivity of the sensorAlthough LSPR spectroscopy is a totally nonselective sensorplatform, a high degree of analyte selectivity can be conferredusing the specificity of surface-attached ligands and passivationof the sensor surface to nonspecific binding. For this reason, aset of control experiments were performed to show that thestreptavidin and antibiotin antibody binding to the sensor sur-face containing no capture ligand (biotin), prebiotinylatedstreptavidin binding to a sensor surface with biotin and bovineserum albumin in large excess, simulating a clinical sample,binding to a sensor surface with biotin, all produce wavelengthshift responses less than that corresponding to the LOD [2]. Thepositive results found with these model systems were gainedafter the development of proper surface chemistries and correctrinsing buffers. In order to realize the full potential of this tech-nique, the development of highly specific biomarkers that cap-ture specific analytes will aid in the continued success of theLSPR sensor chip as well as other devices that rely on molecularinteractions for a signal transduction.

Detection of disease markers Alzheimer’s disease is the leading cause of dementia in peopleover 65 years of age and affects an estimated 4 million Ameri-cans [61]. Although first characterized almost 100 years ago byAlois Alzheimer, who discovered brain lesions, now known asplaques, and tangles in the brain of a middle-aged woman whodied with dementia in her early fifties [62], the molecular causeof the disease is not understood. Also, an accurate diagnostictest has yet to be developed. However, two inter-related theoriesfor Alzheimer’s disease have emerged that focus on the putativeinvolvement of neurotoxic assemblies of a small 42-amino acidpeptide known as amyloid-β [63,64]. Although a normal proteincatabolite, amyloid−β is abundant in Alzheimer’s disease braintissue, where it polymerizes into extremely large amyloid fibrils.Insoluble deposits of amyloid fibrils constitute the proteina-ceous cores of Alzheimer’s disease plaques. The widely investi-gated amyloid cascade hypothesis suggests that the amyloidplaques cause neuronal degeneration and, consequently, mem-ory loss and further progressive dementia. In this theory, theamyloid β protein monomers, present in all normal individuals,do not exhibit toxicity until they assemble into amyloid fibrils[65]. However, there is a poor correlation between plaque sites

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532 Expert Rev. Mol. Diagn. 4(4), (2004)

and disease progression, and more recent studies have estab-lished that amyloid plaques are not the only toxic form of amy-loid-β [66]. The other toxins are known as amyloid-β-deriveddiffusible ligands (ADDLs). In contrast to the fibrils, ADDLsare small, globular and readily soluble, comprising of 3–24-meroligomers of the amyloid-β monomer [67]. ADDLs are potentand selective CNS neurotoxins. Perhaps even more signifi-cantly, they inhibit mechanisms of synaptic information storagewith great rapidity prior to any neurodegenerative conse-quences [67]. ADDLs have now been confirmed to be greatlyelevated in autopsied brains of Alzheimer’s disease subjects [68].Therefore, it would be of great value to develop new tests capa-ble of quantifying ADDLs in patient populations. Such anultrasensitive method for ADDL/anti-ADDL antibody detec-tion could potentially emerge from LSPR nanosensor technol-ogy, providing an opportunity to develop the first clinical labo-ratory diagnostic test for Alzheimer’s disease.

Preliminary results indicate that the LSPR nanosensor canbe used to aid the diagnosis of Alzheimer’s disease [69,70]. Inthese studies, antibodies that specifically interact withADDLs were decorated onto the silver nanoparticle surface.The sensor’s viability was tested by exposing this surface tomodel target ADDL molecules. An additional antibody thenamplified the LSPR nanosensor’s response. After the demon-stration of the model assay, the target molecules that weresandwiched in the previous experiment were substituted bycerebral spinal fluid from diseased and nondiseased patients[69]. The fluid was centrifuged to remove large pieces of cellu-lar material; otherwise, the samples were not further modi-fied. In preliminary assays that took less than 2 h to perform,huge differences were observed upon analysis of the two types(diseased and nondiseased) of samples.

Miniaturization of the sensorClearly, an obvious method to improve the limit of detection ofthe nanoparticle array system would be to reduce the number ofnanoparticles probed. A key to exploiting single nanoparticles assensing platforms is to develop a technique to monitor the LSPRof individual nanoparticles with a reasonable signal-to-noiseratio. UV-vis absorption spectroscopy does not provide a practi-cal means of accomplishing this task. Even under the most favo-rable experimental conditions, the absorbance of a single nano-particle is very close to the shot noise-governed LOD. Instead,resonant Rayleigh scattering spectroscopy is the most straight-forward means of characterizing the optical properties of individ-ual metallic nanoparticles. Similar to fluorescence spectroscopy,the advantage of scattering spectroscopy lies in the fact that thescattering signal is being detected in the presence of a very lowbackground. The instrumental approach for performing theseexperiments generally involves using high magnification micros-copy coupled with oblique or evanescent illumination of thenanoparticles. Recently, the authors demonstrated that a singlesilver nanoparticle can be used to detect a submonolayer coverageof streptavidin [25]. In these experiments, chemically synthesizedAg nanoparticles were dispersed on a glass coverslip in a flow cell

and a dark field image was collected. After incubation in 10 nMstreptavidin, a +12.7 nm shift was monitored, a responseattribute to less than 700 streptavidin molecules.

Comparison with an industry standard: SPR sensorsDuring the course of these findings, it was realized that the sen-sor transduction mechanism of this LSPR-based nanosensor isanalogous to that of flat surface, propagating SPR sensors(TABLE 1). For approximately 20 years, SPR sensors – copper,gold or silver planar films – have been used as refractive index-based sensors to detect analyte binding at or near to a metalsurface. They have also been widely used to monitor a broadrange of analyte–surface-binding interactions, including theadsorption of small molecules [71–73], ligand–receptor binding[74–77], protein adsorption on self-assembled monolayers [78,79],antibody–antigen binding [80], DNA and RNA hybridization[81–83], and protein–DNA interactions [84].

Just as in LSPR spectroscopy, the sensing mechanism ofSPR spectroscopy is based on the measurement of smallchanges in refractive index that occur in response to analytebinding at or near to the surface of a noble metal (e.g., Au, Agand Cu) [85]. Chemosensors and biosensors based on SPRspectroscopy possess many desirable characteristics including:

• Refractive index sensitivity on the order of one part in 105–106 corresponding to an areal mass sensitivity of approximately 1–10 pg/mm2

[71,72,86]

• Long-range sensing length scale determined by the exponential decay of the evanescent electromagnetic field, Lz approximately 200 nm [71]

• Multiple instrumental modes of detection (viz., angle shift, wavelength shift and imaging) [85]

• Real-time detection on the 10-1–103 s time scale for measurement of binding kinetics [72,73,87,88]

• Lateral spatial resolution on the order of 10 µm, enabling multiplexing and miniaturization, especially using the SPR imaging mode of detection [85]

• The system is available commercially

Important differences between the SPR and LSPR sensors arethe comparative refractive index sensitivities and the character-istic electromagnetic field decay lengths. SPR sensors exhibitlarge refractive index sensitivities (~2 × 106 nm/RIU) [71]. Forthis reason, the SPR response is often reported as a change inrefractive index units. The LSPR nanosensor, on the otherhand, has a modest refractive index sensitivity(~2 × 102 nm/RIU) [1]. Given that this is four orders of magni-tude smaller in comparison with the SPR sensor, initialassumptions were made that the LSPR nanosensor would be10,000-times less sensitive than the SPR sensor. However, thisis not the case. In fact, the two sensors are very competitive intheir sensitivities. The short (and tunable) characteristicelectromagnetic field ld, provides the LSPR nanosensor with itsenhanced sensitivity [5,18]. These LSPR nanosensor results indi-cate that the decay length, ld, is approximately 5–15 nm, or1–3% of the light’s wavelength, and depends on the size, shape

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and composition of the nanoparticles. This differs greatly fromthe 200 × 300-nm decay length, or approximately 15–25% of thelight’s wavelength, for the SPR sensor [71]. The smallest footprintof the SPR and LSPR sensors also differ. In practice, SPR sensorsrequire at least a 10 µm × 10 µm area for sensing experiments.For LSPR sensing, this spot size can be minimized to a largenumber of individual sensing elements (1 × 1010 nanoparticlesfor a 2-mm spot size, nanosphere diameter = 400 nm) down to asingle nanoparticle (with an in-plane width of ~20 nm) using sin-gle nanoparticle techniques [24]. The nanoparticle approach candeliver the same information as the SPR sensor, thereby minimiz-ing its pixel size to the sub-100 nm regimen. Due to the lowerrefractive index sensitivity, the LSPR nanosensor requires no tem-perature control, whereas the SPR sensor (with a large refractiveindex sensitivity) does. The final and most dramatic differencebetween the LSPR and SPR sensors is cost. Commercialized SPRinstruments can vary between US$150,000–300,000 (plus sup-plies), whereas the prototype and portable LSPR system costs lessthan US$5000 (plus supplies).

However, a unifying relationship between these two seem-ingly different sensors is that both sensors’ overall response canbe described using EQUATION 2 [71]:

where ∆λmax is the wavelength shift response, m is the refrac-tive index sensitivity, ∆n is the change in refractive indexinduced by an adsorbate, d is the effective adsorbate layerthickness, and ld is the characteristic electromagnetic fielddecay length. It is important to note that for planar SPR sen-sors, this equation quantitatively predicts an adsorbate’s affecton the sensor. When applied to the LSPR nanosensor, thisexponential equation approximates the response for adsorbatelayers but does not provide a fully quantitative explanation ofits response [5,18]. Similar to the SPR sensor, the LSPR nano-sensor’s sensitivity was realized to arise from the distancedependence of the average induced square of the electric fieldsthat extend from the nanoparticle’s surface. This work pro-vides important first steps towards the unified view of LSPRand SPR spectroscopies.

Outlook & conclusionsBriefly looking to the future, a reasonable extrapolation of thecurrent data leads us to expect that by optimizing these size-and shape-tunable nanosensor materials and by using singlenanoparticle spectroscopic techniques, it will be possible to:• Reach sensitivities of a few molecules, perhaps even a single

molecule, per nanoparticle sensor element• Reduce the time scale for real-time detection and the study of

protein binding kinetics by two to three orders of magnitude,

Table 1. Comparison between SPR and LSPR sensors.

Feature/characteristic SPR LSPRLabel-free detection Yes [73,75,82,89] Yes [1,2,4,24]

Distance dependence ~1000 nm [71] ~30 nm (size tunable) [5,18]

Refractive index sensitivity 2 × 106 nm/RIU [71,72,74,86] 2 x 102 nm/RIU [1,18]

Modes Angle shift [85]

Wavelength shiftImaging

Extinction [2]

Scattering [24,25]

Imaging [24,25]

Temperature control Yes No

Chemical identification SPR-Raman LSPR-Raman scattering

Field portability No Yes

Commercially available Yes No

Cost US$150,000–300,000 US$5000 (Multiple particles)US$50,000 (Single nanoparticle)

Spatial resolution ~10 µm × 10 µm [85,90] One nanoparticle [24,25,91]

Nonspecific binding Minimal (determined by surface chemistry and rinsing) [85–87,89,92]

Minimal (determined by surface chemistry and rinsing) [2]

Real-time detection Time scale = 10-1 to 103 s Planar diffusion [72,73,87,88,93]

Time scale = 10-1 to 103 s,Radial diffusion [24]

Multiplexed capabilities Yes [94,95] Yes - possible

Small molecule sensitivity Good [72] Better [18]

Microfluidics compatibility Yes Possible

LSPR: Localized surface plasmon resonance; RIU: Refractive index unit; SPR: Surface plasmon resonance.

∆λmax m∆n 1 e2d–( ) ld⁄

–[ ]=( )

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534 Expert Rev. Mol. Diagn. 4(4), (2004)

since nanoparticle sensor elements will operate in a radialrather than planar diffusion mass transport regime

• Implement massively parallel bioassays for high-throughputscreening applications while maintaining extremely lowsample volume requirementsFinally, it is pointed out that LSPR nanosensors can be

implemented using extremely simple, small, light, robust,low-cost equipment for unpolarized, UV-vis extinction spec-troscopy in transmission or reflection geometry. The instru-mental simplicity of the LSPR nanosensor approach isexpected to greatly facilitate field-portable environmental orpoint-of-care medical diagnostic applications.

Expert opinionThe use of nanoparticles for the highly sensitive and selectivedetection of biomolecules has been demonstrated with bothmodel systems and complex human samples. Additionally, ithas been shown that this assay can be clearly minimized toone nanoparticle via the implementation of dark field micros-copy. The next challenge for sensor development lies in thecapability of formatting the sensor into an array format andintegrating this into a microfluidic chip. By improving thealready established (and commercially available) SPR sensingdevice, this is foreseeable in the near future. The choice ofnanoparticle size and shape will be critical for the optimiza-tion of the sensor’s response. Future studies will continue toanalyze complex biological samples and if fully realized, willprovide vital information regarding disease understanding.

Five-year viewTo date, most research in the area of nanoscience has beenapplied towards an understanding of the fundamental scienceinvolved rather than the implementation of successful applications

in nanotechnology. During the next 5 years, this will change.The development of nanodevices, including nanosensors thatare highly sensitive and selective (give low false positives andnegatives), will provide a major improvement over currenttechnologies. Instruments that provide high-throughputscreening, for example, array imaging and identification ofbiomarkers (i.e., Raman tags), for drug discovery and diseasediagnosis will uncover information vital to the understand-ing and treatment of disease that may lead to the design ofbetter drug candidates for its treatment or prevention. Theseaccomplishments will be met via the achievement of singlemolecule detection, incorporation of molecular identifica-tion using Raman or surface-enhanced Raman spectroscopy,implementation of in vivo sensing using labeled nanoparti-cles and introduction of microfluidics into the chip design.By overcoming the current obstacles in device development,this nanotechnology will become available to hospitals andpharmaceutical companies, as well as to any laboratory, at anaffordable price

AcknowledgementsThe authors acknowledge the support of the Nanoscale Sci-ence and Engineering Initiative of the National ScienceFoundation (NSF) under NSF Award Number EEC-0118025. Any opinions, findings and conclusions or recom-mendations expressed in this material are those of theauthors and do not necessarily reflect those of the NSF. AJHaes also wishes to acknowledge the American ChemicalSociety Division of Analytical Chemistry and DuPont for agraduate fellowship. The authors are grateful for useful dis-cussion, technical support and the expert assistance providedby Lei Chang, William Klein, Adam McFarland, GeorgeSchatz, Karen Shafer-Peltier and Shengli Zou.

Key issues

• Silver nanoparticles are extremely sensitive to small changes in their surrounding dielectric environment. For this reason, their surfaces can be chemically functionalized and used to provide highly sensitive and selective detection of biological targets.

• These localized surface plasmon resonance (LSPR) nanosensors have been demonstrated as successful binding affinity sensors for the detection of proteins and antibodies and in the future, for DNA, RNA and peptide nucleic acids.

• The continued development of surface chemistries that provide high specificity of molecular interactions and a minimum of nonspecific interactions is required for the continued development of a chip-based LSPR sensor. Along these lines, the development of more precise (or more general) biomarkers for molecular interactions will contribute to this success.

• The implementation of microfluidic channels into chip design will greatly increase the capabilities of the chip’s detection.

• Improvement of nanoparticle adhesion to the sensor substrate is vital for the long-term viability of the sensor. Current approaches provide sufficient nanoparticle adhesion but this should be improved prior to the development of acommercialized chip.

• To provide an affordable product for any laboratory, the integration of well plates with LSPR chips is being developed.

• Continued development of large arrays for multiplexed analysis and associated imaging modalities is vital for the widespread applicability of the LSPR nanosensor.

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Affiliations • Amanda J Haes, PhD

Research Assistant, Northwestern University, Department of Chemistry, 2145 Sheridan Road, Evanston, IL 60208–3113, USATel.: +1 847 491 2952 Fax: +1 847 491 7713a-haes@chem.northwestern.edu

• Richard P Van Duyne, PhD

Morrison Professor of Chemistry, Northwestern University, Department of Chemistry, 2145 Sheridan Road, Evanston, IL 60208–3113, USATel.: +1 847 491 3516Fax: +1 847 491 7713vanduyne@chem.northwestern.edu