Detection of virus-like nanoparticles via scattering using a chip-scaleoptical biosensorRongjin Yan, N. Scott Lynn, Luke C. Kingry, Zhangjing Yi, Tim Erickson et al. Citation: Appl. Phys. Lett. 101, 161111 (2012); doi: 10.1063/1.4758294 View online: http://dx.doi.org/10.1063/1.4758294 View Table of Contents: http://apl.aip.org/resource/1/APPLAB/v101/i16 Published by the American Institute of Physics. Related ArticlesRole of electropores on membrane blebbing—A model energy-based analysis J. Appl. Phys. 112, 064703 (2012) Local streamline generation by mechanical oscillation in a microfluidic chip for noncontact cell manipulations Appl. Phys. Lett. 101, 074102 (2012) Wireless power transfer to a cardiac implant Appl. Phys. Lett. 101, 073701 (2012) Microfluidic impedance spectroscopy as a tool for quantitative biology and biotechnology Biomicrofluidics 6, 034103 (2012) Treatment of enterococcus faecalis bacteria by a helium atmospheric cold plasma brush with oxygen addition J. Appl. Phys. 112, 013304 (2012) Additional information on Appl. Phys. Lett.Journal Homepage: http://apl.aip.org/ Journal Information: http://apl.aip.org/about/about_the_journal Top downloads: http://apl.aip.org/features/most_downloaded Information for Authors: http://apl.aip.org/authors

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Detection of virus-like nanoparticles via scattering using a chip-scale opticalbiosensor

Rongjin Yan,1 N. Scott Lynn,2 Luke C. Kingry,3 Zhangjing Yi,1 Tim Erickson,1

Richard A. Slayden,3 David S. Dandy,2,4 and Kevin L. Lear1,4

1Department of Electrical and Computer Engineering, Colorado State University, Colorado 80523-1373, USA2Department of Chemical and Biological Engineering, Colorado State University, Colorado 80523-1370, USA3Department of Microbiology, Immunology and Pathology, Colorado State University, Colorado 80523-2025,USA4School of Biomedical Engineering, Colorado State University, Colorado 80523-1376, USA

(Received 20 July 2012; accepted 26 September 2012; published online 17 October 2012)

A local evanescent array coupled biosensor is used to detect spherical polystyrene nanoparticles

with diameters of 40 nm and 200 nm, whose sizes and refractive index are similar to virus particles.

The sensitivity is �1%/particle for 200 nm particles and 0.04%/particle for 40 nm particles. Mie

scattering in an evanescent field theory is used to model the scattered light intensity for both sizes

of nanoparticles. VC 2012 American Institute of Physics. [http://dx.doi.org/10.1063/1.4758294]

In recent years, a number of viral diseases, such as bird

flu and H1N1, have motivated increased interest in the devel-

opment of new point-of-care virus detection techniques.1,2

Timely diagnosis is necessary in order to quarantine the sick

and provide effective antiviral treatment. Traditional viral

diagnostic tools, such as enzyme-linked immunosorbent

assay (ELISA),3 restriction enzyme analysis (REA),4 and

real time polymerase chain reaction (RT-PCR),5 are effective

but have several drawbacks that can inhibit widespread,

rapid testing. These drawbacks include 2–10 days for cell

culture, complicated sample preparation procedures, and

requirements for immunoassay procedures within specialist

laboratories. Compared with these established techniques, a

desired biosensor for virus detection should be rapid, inex-

pensive, sensitive, and label-free.6

Label-free photonic biosensors provide a relatively easy

and inexpensive way to detect molecular interactions and quan-

tify their kinetics.7–9 One of the most popular detection princi-

ples is based on the evanescent field modulation. The binding

of molecules in optical evanescent field causes the refractive

index change in the binding area, which can be measured using

the induced guided optical mode’s modulation. Many biosen-

sors based on binding in evanescent fields provide very high

sensitivities including surface plasmon resonance (SPR) bio-

sensors10–12 and micro ring resonator biosensors.13,14 However,

most of these sensors cannot simultaneously sense multiple

analytes on a single waveguide since the modification of the

transmitted optical mode’s phase or amplitude is measured at

the end the waveguide or optical path.

A planar dielectric waveguide based local evanescent

array coupled (LEAC) sensor has been proved to be a promis-

ing platform for point-of-care disease diagnostics,15 and in

previous work,16–18 a LEAC biosensor has been employed to

detect different antigen/antibodies complexes. The immunoas-

say mechanism of the LEAC sensor relies on specific binding

of analytes or targets to one of several localized regions of im-

mobilized biological molecular probes, such as antibodies or

ssDNA, to modify the waveguide cross-section. The increased

refractive index of the upper cladding shifts the local evanes-

cent field up and hence a decreases photocurrent in the detec-

tor array under the waveguide.15 When applied to the

detection of virus particles, the LEAC sensor operates with an

entirely different mechanism. Viruses can be captured on to

the top surface of the LEAC waveguide via a variety of immo-

bilized affinity receptors, including antibodies, aptamers, and

engineered affinity proteins. Instead of the thin uniform

adlayer (n¼ 1.55) formed by the immobilized antigen/anti-

body molecules, these captured virus particles form isolated

“islands” on the waveguide. Due to the refractive index differ-

ence between the air (n¼ 1) and the particles (n¼ 1.55) in the

evanescent field, light is scattered out of the guided mode.

Some of the scattered light is incident on the detectors under

the waveguide, leading to an increase in photocurrent in the

region where particles are present, as shown in Fig. 1.

Compared to SPR biosensors or ring/disk resonator bio-

sensors, LEAC sensor does not require a laser source but can

be operated with LEDs that offer both noise and system simpli-

fication advantages. Furthermore, the LEAC biosensor is a

non-resonant device with minimal temperature dependence.

LEAC sensor is a silicon photonic device, which makes it easy

to integrate photodetector and integrated circuit into a single

silicon chip. An entire multianalyte LEAC system including

optical source, detector array, integrated signal processing, and

telemetry electronics with a simple capillary force driven

microfluidics system should occupy a volume <1 cm3.

In order to understand the relationship between the scat-

tered light intensity and the particle size, the scattering

mechanism of small particles in an evanescent field is stud-

ied. The size parameter p is defined as p ¼ pd=k, where d is

the diameter of the particle and k is the wavelength of the

incident light. For the 200 nm polystyrene nanoparticles

(p¼ 0.96) used in our experiment, the Rayleigh approxima-

tion is not applicable, so Mie theory,19 also called Lorenz–

Mie theory, is used in the following calculations.

Since the virus particles are immobilized above the wave-

guide, the incident light in this case is the evanescent field that

exponentially decays into the upper cladding with a decay

constant, c, calculated as c ¼ 2pk ðn2

ef f � n2upperÞ

1=2, where neff

is the effective refractive index of the waveguide mode and

nupper is the refractive index of the waveguide upper cladding.

0003-6951/2012/101(16)/161111/4/$30.00 VC 2012 American Institute of Physics101, 161111-1

APPLIED PHYSICS LETTERS 101, 161111 (2012)

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For this problem, we will assume that the captured virus

particle in the shape of a sphere with radius r. Compared

with the definition of scattering cross section for plane-wave

excitation,20 variable intensity over the cross-sectional area

of the particle is considered in the evanescent wave case.

The guided mode in the LEAC sensor is a TE mode.16 The

Mie scattering cross section, Asca, is

Asca ¼k2

2pn2upper

X1n¼1

ð2nþ 1ÞðjanðrÞj2PnðhiÞ þ jbnðrÞj2CnðhiÞÞ;

(1)

where PnðhiÞ and CnðhiÞ are the size-independent weighting

factors normalized from Mie coefficients to account for the

incident wave being an evanescent field, an(r) and bn(r) are

the Mie coefficients that depend on r, and hi is the incident

angle as shown in Fig. 2.

The averaged incident light intensity over the cross-

sectional area of the particle perpendicular to the Poynting

vector of the evanescent wave is21

�I0 ¼ I0e�2cr np

nuppersinhi

I1ð2crÞcr

; (2)

where I1(2cr) is the modified Bessel function of order 1 with

argument 2cr. I0 is incident evanescent wave intensity at

z¼ 0.

From Eqs. (1) and (2), the relative amount of scattering

can be quantified with the ratio of scattered power for par-

ticles of radius r compared to that for radius r0

PðrÞ

¼e�2ar I1ð2crÞ

cr

X1n¼1

ð2nþ1ÞðjanðrÞj2PnðhiÞþjbnðrÞj2TnðhiÞÞ

e�2ar0I1ð2cr0Þ

cr0

X1n¼1

ð2nþ1Þðjanðr0Þj2PnðhiÞþjbnðr0Þj2TnðhiÞÞ:

(3)

For a wavelength of k¼ 654 nm and the single mode wave-

guide used in LEAC sensor,16 the normalized scattered

power for particles with different diameters from 20 to

200 nm, which covers most of the virus sizes seen in nature,

is plotted in Fig. 3. The scattered power is strongly depend-

ent on the particle size. The scattered power of a single

200 nm particle is 39 times larger than a 40 nm particle.

Nanoparticle scattering was measured experimentally

using LEAC biosensors fabricated using a commercial

0.35 lm CMOS process.15 A buried metal-semiconductor-

metal photodetector array was integrated on the biosensor

chip beneath a SiN/SiO2 waveguide.

For measurement purposes, light was coupled into the

waveguide using end-fire coupling with 654 nm laser light

through a single mode fiber (4/125 lm core/cladding diame-

ter). A 16.8 Hz3V peak-to-peak AC signal with 2.6 V offset

was generated by a HP8116A function generator and was

used to both modulate the 654 nm laser diode and serve as

the reference signal for the lock-in amplifier. A PAR-5210

dual-phase lock-in amplifier was used to measure the photo-

currents. Signal from a third needle that probed a fixed pad

was used to normalize the signal collected from other pads,

in order to account for change in the light power coupled to

the waveguide resulting from fiber movement.

FIG. 1. (a) Local scattered light detection mechanism; (b) SEM image of the

LEAC sensor cross section.

FIG. 2. Geometry of the scattering problem. FIG. 3. Normalized scattered power vs. particle diameter.

161111-2 Yan et al. Appl. Phys. Lett. 101, 161111 (2012)

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To verify the results shown in Fig. 3, we utilized the af-

finity interaction between biotin and NeutrAvidid to selec-

tively capture both 40 and 200 nm fluorescently labeled

NeutrAvidin coated polystyrene nanospheres onto the LEAC

sensor. The polystyrene nanoparticles have similar optical

properties to virus particles (n¼ 1.57),22 where the 40 and

200 nm particles represent the smaller and larger characteris-

tic sizes for diagnostically relevant viruses.23 Biotinylated

bovine serum albumin (BSA) was printed in discrete 288 lm

diameter spots onto the LEAC waveguide24 using a microar-

ray printer, after which the remaining surface was passivated

in 45 mgmL-1 BSA in PBS. These LEAC biosensors were

then incubated in a solution containing NeutrAvidin coated

particles (0.01, 0.005, or 0.001% solids) 45 mgmL-1 BSA,

and 1% tween-20 in PBS.

After measurement of the baseline photocurrent distri-

bution of each buried photodetector, the LEAC sensor was

incubated for 30 min in a solution containing 10-2–10-3% w/

w NeutrAvidin coated, fluorescent microspheres (Invitro-

gen), 45 mgmL-1 BSA, and 1% w/w tween-20 in PBS. The

affinity existing between immobilized biotin on the wave-

guide and NeutrAvidin coated on the particles was utilized

to mimic the affinity interaction between an immobilized

biorecognition element and its corresponding virus. After

incubation, the sensor was sequentially rinsed and dried as

described above. Figure 4(a) displays the selective capture of

200 nm diameter particles, where it can be seen there is a

clear boundary between the region printed with biotinylated

BSA and the BSA blocked region that did not capture par-

ticles. Figure 4(b) displays the fluorescent intensity of the

LEAC sensor surface after incubation with 200 nm particles,

where it can be seen that the particles are isolated to the

waveguide above the photodetectors 8–11. For each sensor,

the number of particles immobilized to the waveguide above

each photodetector was quantified using either SEM (for 40

and 200 nm diameter particles) or fluorescent microscopy

(200 nm particles), where particles were counted if in physi-

cal contact with the waveguide. There were no significant

differences between the SEM and microscopy methods per-

taining to the 200 nm particles.

In this work, we utilize the photocurrent modulation ra-

tio for the quantification of the scattering effects due to the

presence of captured nanoparticles on the waveguide surface.

The photocurrent modulation ratio is defined as (Ip� I0)/I0,

where I0 and Ip are the photocurrents measured before and

after nanoparticle incubation, respectively. Figure 4(c) plots

the modulation ratio as a function of the photodetector num-

ber for photodetectors 4–15. Figure 4(c) also includes the

number of particles captured onto the waveguide above each

photodetector (seen in Fig. 4(b)), measured via fluorescent

microscopy methods. It can be seen that on detectors 4–7,

where there were less than 2 particles captured per photode-

tection region, the modulation ratio is approximately zero.

The modulation ratio on detectors 8–11 sharply increases

from evanescent scattering due to the presence of the poly-

styrene nanoparticles in higher densities. The raw modula-

tion ratio drops to a negative value after detector 12, as

increased levels of scattering in the particle immobilized

region reduces the overall power of guided light in the

waveguide.

To compensate for the decreased power measured at the

ith detector due to increased levels of scattering in regions of

high particle capture densities, we introduce a modulation

correction factor, Ki,

Ki ¼ e�

�aw�ði�1Þþ

Pi�1

n¼1

lnðe�aw � DSnÞ�; (4)

where aw is the propagation loss along the waveguide before

the particle patterning. DSn is the particle induced scattering

ratio change on detector n, which can be defined as DSn

¼ Nn

N1M1ð1� e�awÞ. Nn is the particle number on detector n

and M1 is the modulation ratio on detector 1. Figure 4(c)

also plots the corrected modulation ratio as a function of

photodetector number.

Figure 5(a) shows the corrected modulation ratio vs. a

particle count, where the data were collected from 11 photo-

detectors over 3 different LEAC biosensor chips. As

expected, it can be seen that corrected modulation ratio is a

linear function of N. The sensitivity for these 200 nm par-

ticles is approximately 1%/particle.

The relationship between the corrected modulation ratio

and particle number for 40 nm particles is plotted in Fig. 5(b)

for a single biosensor chip. The sensitivity for 40 nm par-

ticles is �0.04%/particle. The sensitivity for 200 nm particle

FIG. 4. (a) SEM picture of immobilized polystyrene nanoparticles with di-

ameter of 200 nm. Dotted line shows a clear boundary between the region

printed with biotinylated BSA and the BSA blocked region that did not cap-

ture particles. (b) Fluorescent microscope photo showing the patterned

region on the chip. (c) Particle number distribution along the waveguide and

the modulation ratio measured from different detectors. The particle diame-

ter is 200 nm.

161111-3 Yan et al. Appl. Phys. Lett. 101, 161111 (2012)

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is approximate 25 times larger than 40 nm particles. This ex-

perimental ratio is slightly smaller with respect to the results

calculated using Eq. (3), where it is expected that 200 nm

particles should provide a sensitivity 39 greater than that

expected by the 40 nm particles. This difference may due to

the local field shift effect caused by the immobilized par-

ticles, which is not considered in the derivation of Eq. (3).

Polystyrene nanoparticles with two different diameters,

40 nm and 200 nm, were detected by a LEAC biosensor. The

sensitivity for the 200 nm particle is �1%/particle and it is

0.04%/particle for the 40 nm particle. The particle scattering

light intensity was calculated using the Mie scattering

method in evanescent field described in reference, and the

calculated ratio of scattered light intensity for 40 nm and

200 nm particles was 39, which is larger than the experimen-

tal ratio, 25. The experiment results show the potential of

this biosensor in the virus detection application. Considering

the top surface of the biosensor is SiNx, it would be an easy

process to integrate microfluidic delivery of the virus par-

ticles to the LEAC biosensor. This will give the biosensor

real-time monitoring ability and largely reduce the test time.

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FIG. 5. (a) Dependence of corrected modulation ratio on particle number for

200 nm diameter polystyrene nanoparticles. The sensitivity is �1%/particle.

(b) Dependence of corrected modulation ratio on particle number for 40 nm

diameter polystyrene nanoparticles. The sensitivity is �0.04%/particle.

161111-4 Yan et al. Appl. Phys. Lett. 101, 161111 (2012)

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