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Cellphone-Enabled Microwell-Based Microbead Aggregation Assay for Portable Biomarker Detection

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Quantitative biomarker detection methods featured with rapidity, high accuracy, and label-free are demonstrated for the development of point-of-care (POC) technologies or "beside" diagnostics. Microbead aggregation via protein-specific linkage provides an effective approach for selective capture of biomarkers from the samples, and can directly readout the presence and amount of the targets. However, sensors or microfluidic analyzers that can accurately quantify the microbead aggrega-tion are scared. In this work, we demonstrate a microwell-based microbeads analyzing system, by which online manipula-tions of microbeads including trapping, arraying, and rotations can be realized, providing a series of microfluidic approaches to layout the aggregated microbeads for further convenient characterizations. Prostate specific antigen is detected using the proposed system, demonstrating the limit of detection as low as 0.125 ng/mL (3.67 pM). A two-step reaction kinetics model is proposed for the first time to explain the dynamic process of microbeads aggregation. The developed microbeads aggrega-tion analysis system has the advantages of label-free detection, high throughput, and low cost, showing great potential for portable biomarker detection.
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Cellphone-Enabled Microwell-Based Microbead Aggregation Assay
for Portable Biomarker Detection
Weiwei Cui,
,,§
Meihang He,
,§
Luye Mu,
Zuzeng Lin,
Yanyan Wang,
Wei Pang,
Mark Reed,
and Xuexin Duan*
,
State Key Laboratory of Precision Measuring Technology & Instruments, College of Precision Instrument and Optoelectronics
Engineering, Tianjin University, Tianjin 300072, China
Department of Electrical Engineering and Yale University, New Haven, Connecticut 06520, United States
*
SSupporting Information
ABSTRACT: Quantitative biomarker detection methods featured
with rapidity, high accuracy, and label-free are demonstrated for the
development of point-of-care (POC) technologies or beside
diagnostics. Microbead aggregation via protein-specic linkage
provides an eective approach for selective capture of biomarkers
from the samples, and can directly readout the presence and amount
of the targets. However, sensors or microuidic analyzers that can
accurately quantify the microbead aggregation are scared. In this work,
we demonstrate a microwell-based microbeads analyzing system, by
which online manipulations of microbeads including trapping,
arraying, and rotations can be realized, providing a series of
microuidic approaches to layout the aggregated microbeads for
further convenient characterizations. Prostate specic antigen is
detected using the proposed system, demonstrating the limit of
detection as low as 0.125 ng/mL (3.67 pM). A two-step reaction kinetics model is proposed for the rst time to explain the
dynamic process of microbeads aggregation. The developed microbeads aggregation analysis system has the advantages of label-
free detection, high throughput, and low cost, showing great potential for portable biomarker detection.
KEYWORDS: microbead aggregation, microuidics, cellphone-enabled detection, PSA detection, two-step reaction, POC
Since biomarkers are objectively measured and evaluated as
indicators of normal biological processes, pathogenesis, or a
pharmacological response to therapeutic intervention,
13
accurate detection and quantication of biomarkers in a
convenient way is a central goal of modern biotechnology.
For clinical applications, sandwich immunoassays are the most
widely applied assay formats for target biomarker detections.
Generally, these methods require some type of labeling
(enzymatic or uorescent) for selective and sensitive report
of target analytes.
47
Specialized instruments (e.g., plate
reader) and numerous washing steps are required as well.
Recently, advances in quantitative biomarker detection
methods featured with rapidity, high accuracy, portability, and
label-free readout have enabled the concept of point-of-care
(POC) technologies or bedsidediagnostics.
813
Among
them, anity biosensors such as impedance spectroscopy,
1417
potentiometric sensor,
1821
surface-enhanced Raman scatter-
ing technologies,
22,23
and gravimetric sensor
2427
have been
developed to directly recognize and quantify the target
biomarkers by immobilizing specic acceptors on the trans-
ducers, which do not require any type of labeling. Another
promising label-free detection method is based on biofunction-
alized micro/nanoparticles, which have been demonstrated and
widely applied as molecule probes or carriers for direct capture
of target protein or DNA from complex samples.
13,2830
Colorimetric method or particle size analyzer has been applied
for biomarker detection through analyzing the aggregation
status of these particles.
3134
Aggregations of nanoparticles are
induced by the specic protein interactions (e.g., antibody
antigen) between the receptor immobilized on the particle
surface and the target analyte in solution. Generally, more and
larger aggregated clusters will be generated with higher
biomarker concentrations. However, the accurate relationship
between nanoparticle aggregations and the biomarker concen-
tration has not been thoroughly understood because of the lack
of compatible methods or tools to precisely quantify the
aggregation status, including information regarding the number
of the nanoparticles within each cluster. In addition to
nanoparticles, aggregation of antibody functionalized microbe-
ads for antigen detection has been recently demonstrated using
an impedance sensor to count the number of the aggregated
clusters.
35,36
Though an impedance sensor can directly read out
the number of clusters, it suers problems such as easy
Received: November 22, 2017
Accepted: January 19, 2018
Published: January 19, 2018
Article
pubs.acs.org/acssensors
Cite This: ACS Sens. 2018, 3, 432440
© 2018 American Chemical Society 432 DOI: 10.1021/acssensors.7b00866
ACS Sens. 2018, 3, 432440
congestion, low throughput, and limited resolution for dierent
sized beads. Besides, the developed impedance analyzing
system requires dedicated uid delivery setup and expensive
data acquisition system for fast and low noise electrical signal
processing which hindered the development of such systems
into portable assays. Microuidic analyzers with the capability
of processing the samples in a high throughput and well-
controllable manner are attractive for microbeads character-
izations. Especially, arraying microbeads into a small space
contributes to more ecient observation and characterization
of their aggregations.
In this work, a cellphone-enabled microwell-based POC
system is developed for biomarker detections by analyzing the
aggregation status of the biomarker functionalized microbeads.
The microuidic chip is made of a (PDMS) pattern using a
reusable silicon mold, with microsized features dened by
standard optical lithography. Using the developed system, we
demonstrated the successful detection of prostate specic
antigen (PSA), a critical biomarker in early stage detection of
prostate cancer using anti-PSA functionalized polystyrene (PS)
beads. The system provides a dedicated tool for microbead
trapping and arraying via uidic approaches. Aggregation status,
dened as the percentage of microbeads dimers, was acquired
by a cellphone enabled portable imaging system. The limit of
detection (LOD) of this system is demonstrated to be as low as
3.67 pM, which is beyond the requirement of clinical PSA
detections. Based on the experimental results, the kinetics of
microbeads aggregation is carefully studied, and a two-step
reaction model is proposed to explain the aggregation process.
These results demonstrate that this platform provides a simple,
label-free, low-cost, and point-of-care system for biomarker
detections.
EXPERIMENTAL SECTION
Materials. Carboxyl-terminated PS beads (108/mL) with diameter
of 5 μm were purchased from Sigma; polyclonal PSA (prostate specic
antigen) and anti-PSA (1 mg/mL) were purchased from Sigma-
Aldrich. Poly-L-lysine grafted with oligoethylene glycol (PLL-OEG,
MW = 1530 kDa) is synthesized in our lab according to procedures
described in previous publications.
37,38
HEPES (N-2-hydroxyethylpi-
perazine-N-ethane-sulfonicacid, 10 mM, pH = 7) containing 0.2% v/v
Tween (Sigma) as the surfactant was used as the solvent. Full serum
(F2442, MFCD00132239) was purchased from Sigma. The portable
microscope was purchased from Shenying Optics (SAGA002, Suzhou,
China), and the cellphone can be any commercial cellphone with a
camera of 12 Megapixels (more details about the cellphone enabled
portable microscope system are presented in the Supporting
Information, Figure S1).
Surface Modications of PS Microbeads. Anti-PSA with a
concentration of 8 μg/mL was applied to modify PS beads using
EDC/NHS (N-(3-(dimethylamino)propyl)-N-ethylcarbodiimide hy-
drochloride and N-hydroxysuccinimide) activation approach.
39
Microbeads were washed twice with HEPES buer (10 mM, pH =
7.4), and then activated by EDC (N-3-dimethylanimopropyl-N-
ethylcarbodimide) and sulfo-NHS (N-hydroxysulfosuccinimide) at a
typical concentration of 2 mg/mL (EDC) and 0.5 mg/mL (NHS) in a
shaker (800 rpm) at room temperature. After activation for 30 min,
the microbeads were separated by centrifugation and washed twice
again with HEPES. Antibodies (anti-PSA, 8 μg/mL) diluted in HEPES
were then quickly added to the activated particles. After incubation at
room temperature for 40 min, the supernatant was removed after
centrifugation. Ethanolamine aqueous solutions (20 mM, pH = 7.4)
were then added and reacted with microbeads for another 40 min to
deactivate the unbound carboxyl groups on the surface, followed by
washing twice with HEPES buer. Finally, the antibody conjugated
microbeads were resuspended in HEPES and stored at 4 °C.
Chip Design and Fabrication. The MicroWell-based Microuidic
Chip (MW-MFC) consists of a microchamber for uid delivery and
microwell arrays on the bottom, both of which are made of PDMS.
The microchamber is designed to be 30 mm in length, 5 mm in width,
and 20 μm in height supported by a 4 ×25 pillars array to prevent
channel collapse. Two input channels and one exit channel are
connected with to the chamber, and the width of each is 600 μm.
Microwells on the bottom of the chamber are 10 μm in depth and 20
μm in diameter. Silicon mold was fabricated by reactive ion etch
process, and a releasing agent peruorodecyltriethoxysilane (PFDTS)
was vacuum deposited overnight and then cured in 85 °C for 30 min.
Two-part Sylgard 184 was mixed with a mass ratio of 10:1 and then
degassed for 30 min. The mixture was poured over the silicon mold
and cured at 85 °C for 90 min, after which the cured PDMS was
carefully peeled oand holes were punched to serve as the inlet and
outlet ports. Each PDMS chip was cleaned with ethanol and deionized
water (1:1) solution and gently dried with nitrogen, followed by
treatment with O2plasma cleaner under 120 W for 30 s. The PDMS
was aligned and irreversibly bound together after 30 min curing under
85 °C. The inner surface of PDMS chip was incubated with the PLL-
OEG solution (100 mM, dissolved in HEPES buer, pH = 7.4) for 30
min to achieve a hydrophilic coating.
Principles of Microbeads Manipulations and Quantica-
tions. The sensing mechanism of our system is based on the counting
of the number of the aggregated beads. Anti-PSA modied PS beads
are mixed with dierent concentrations of PSA samples. After
incubation at room temperature, the PS beads will aggregate together
through the specic interactions between PSA and anti-PSA. In
principle, when the number of the beads is designed to be in excess of
the target biomarkers, dimers (two PS beads linked together as Figure
1a shows) are the most likely aggregation status in the nal
product.
35,36
The reacted microbeads were then processed by the
microwell analyzing system to obtain the percentage of the dimers.
Figure 1b presents the arrayed microbeads within the MW-MFC,
which is imaged with a cellphone equipped with a portable optical
microscope. The acquired images were processed with a home
developed MATLAB program and the percentage of the dimer, i.e.,
aggregation ratio, was calculated (as Figure 1c shows) to represent the
aggregation status by eq 1
a
ggregation ratio number of dimers
number of arrayed microbeads 100% (1)
Figure 1. (a) Diagram of the microbeads aggregation process: mixing
PSA into anti-PSA modied polystyrene microbeads suspension forms
bead aggregations induced by the antibodyantigen interactions. (b)
Image of the arrayed microbeads within microwells. (c) Statistics of
the aggregated microbeads acquired by processing the photos. Scale
bar in (b) is 40 μm.
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As Figure 2c shows, the microbeads suspended within the MW-MFC
would be dragged by the uidic drag force
40
μγν=−FKVVR()(/)
dP
21/
2
(2)
where K= 81.2, Ris the radius of the microbead, νis the kinematic
viscosity, and γis the uidic shear rate induced by velocity gradient
near the microbead. Vand Vprepresent the ow velocity and the
microbead moving velocity, respectively. When the ow rate is very
low (below the ow rate Q1labeled in Figure 2c, which represents the
case of a balance between gravity and drag force inside the microwell),
most of the microbeads is deposited onto the substrate by gravity, G.
While the microbeads can be lifted from the substrate and the
microwells when the ow rate is increased to be higher than Q2
(corresponding to the ow rate that induces a lift force equaling to the
gravity). By tuning the ow rate (between Q1and Q2), the
microvortices induced by the microwells
4143
would contribute to
trapping microbeads into the microwells. Finally, HEPES buer was
injected into the chip with a ow rate of 400 nL/min (referring to Q2),
which is high enough to generate a shift force to release the
microbeads outside the microwells and wash them away from the
chamber, while the trapped beads inside the microwells are well
protected from lifting out. When increasing the ow rate up to 500
nL/min (higher than Q2), the trapped microbeads will also be released
by the lift force induced by the ow rotation,
44
following the eq 3
πρω=−FRVV
6
()
L3
fp(3)
Here, ωis the rotation angular velocity of microbeads.
Image Acquisition and Microbeads Detection. PSA of
dierent concentrations was respectively mixed with anti-PSA
modied PS beads suspension in a 500 μL centrifugal tube and then
incubated for 30 min at room temperature. The nal concentration of
microbeads in the incubated mixture was 2 ×106/mL. After
incubation, the suspension was diluted four times with HEPES buer.
Typically, 20 μL of aggregated microbeads suspension was introduced
into the MW-MFC. By controlling the uidic ow as outlined above,
the suspended microbeads are well trapped and arrayed for imaging
analysis. The manipulations of microbeads trapping and arraying
requires 10 min, and the picture acquisition of the arrayed microbeads
in the region of interest (ROI) takes 5 min. For all the experiments, a
total population of approximately 3000 microbeads was trapped and
arrayed corresponding to trap ratio of 3%. Typically, 70 images were
taken to include a sucient number of beads. To eectively acquire
the beads aggregation information from these images, a MATLAB
image-processing program was developed to directly count the
number of the dimers. Microbeads with diameter of 5 μm can be
easily identied with the Hough-transform algorithm. Furthermore, a
cellphone APP has been developed, enabling the whole assay job,
including the image capture, analysis, and data processing all operated
within a single cellphone (the source code of the APP can be
downloaded at the GitHub page https://github.com/linzuzeng/
Microsphere/releases, and detailed information can be found in the
Supporting Information, Figures S2S4).
RESULTS
Chip Fabrication and System Design. Figure 2a,b shows
the schematic of the MW-MFC and the fabrication process.
The chamber is designed as 20 μm in height supported by
micropillars. Two connecting channels are used to introduce
microbeads suspensions and HEPES buer via syringe pumps.
The depth of the microwell is set as 10 μm for high eciency
trapping of 5 μm microbeads in combination with the
microchamber. Hydrophilic modication of the inner surface
of the chip is required to reduce the nonspecic adsorption of
suspended microbeads. This is achieved through coating the
chips with PLL-OEG.
37,38
Figure 2c presents the principles of
the microbeads trap and release with MW-MFC, which forms
the basis of the microbeads array. Since the size of the particle
and microwell is rather large, a cellphone camera connected to
a standard portable microscope is used to acquire the images of
the arrayed microbeads by directly attaching the camera to the
objective lens of the microscope (Figure 2d). Figure 2e
presents a comparison of the MW-MFC images obtained via a
professional microscope (top image) and the cellphone
(bottom image). The aggregation status of the microbeads
can be easily identied from both images, proving the capability
of the developed portable system for image-based microbeads
Figure 2. (a) Schematic and (b) fabrication process of the MicroWell-based Microuidic Chip. (c) Principles of the microbeads trap and release,
wherein the dotted arrow presents the integration force on the microbead. (d) Setup of the cellphone-enabled image acquiring system. (e)
Comparison of the images taken by a professional microscope (top) and by the cellphone (bottom). The labeled rectangle in (b) represents the ROI
which is located at the center of the chip. Images of the particles were taken within the ROI to avoid the errors induced by the higher trapping
eciency near the entrance, side wall, and exit of the chip.
ACS Sensors Article
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434
analysis. It is noted that we used a portable microscope to
demonstrate the portability of the system. It is also possible to
use other commercially available cellphone attachments to
achieve even better portability.
45,46
The arrayed microbeads within the MW-MFC are presented
in Figure 2e. Generally, there are four cases of the trapped
microbeads within one microwell as shown in Figure 3a: (1) no
microbeads; (2) single; (3) dimer; (4) two separated
microbeads. There exists one case that two microbeads may
physically sink together into one microwell and contact each
other without the real linkage through the antigens which
would form a false dimer and hinder the detection accuracy. To
solve this issue, rotation or swing of the trapped microbeads is
introduced by changing the ow rate which could stretch the
physically linked beads within the microwells. The advantage of
using a uidic method is not only due to the capability to move
microbeads into and out of the microwells by tuning the ow
rate in real time, but the uidic manipulation helps to reduce
the false dimers. As Figure 3b shows, the trapped microbeads
rotate or swing under the ow forces. The rotating motions of
the microbead will divide the false dimers into two single beads
as Figure 3c shows.
In addition, the ratio of trapped microbeads within the MW-
MFC, μ(dened as the ratio of number of trapped microbeads
to the number of microwells), is an important factor. The
probability (P) of the trapped microbeads number (x) within
one microwell obeys to the Poisson distribution
47
expressed as
eq 4:
μ
==
!
μ
P
Xx x
() e
x
(4)
When μ= 0.03, the probability of a false dimer in a single well
is 0.044%. To reduce the false aggregated microbeads, the value
of μis kept below 0.03 in the experiments.
PSA Detection. Figure 4a presents the size distribution of
the applied microbeads with average diameter of 5.1 μm. The
commercial counter instrument (Beckman Coulter Multisizer
3) calculates the equivalent diameter of the counted particle by
assuming it with spherical shape by tting the electrical signal.
To test the variation of the dimer number, PSA with
concentration of 12.5 ng/mL, 0.125 ng/mL, and control
group solution were, respectively, introduced into the anti-PSA
modied microbeads suspensions and incubated for 30 min.
Figure 4b shows the microbeads size distribution changes
because of the microbeads aggregation. It is easy to count the
bead number, but the size is dicult to precisely distinguish
between aggregations and single beads via the diameter
distribution curve. This is due to low-ecient of microbead
aggregations via protein specic interactions. Moreover, the
equivalent size of a dimer is comparable with single ones.
Considering the wide distribution range of the single bead size
shown in Figure 4a, the diameter distribution changes will be
dicult to evaluate. Besides, the instrument is expensive and
sample-consuming. While, the developed MW-MFC system
can directly recognize the single and aggregated beads via
microuidic chip-assisted image methods, enabling an accurate
description of the microbeads size distribution.
To demonstrate the capability of using microbeads
aggregation for protein detections, PSA of dierent concen-
trations (250, 125, 12.5, 1.25, and 0.125 ng/mL, corresponding
to 7.35 nM, 3.67 nM, 0.367 nM, 0.0367 nM, and 3.67 pM)
were, respectively, detected using the developed system. Figure
4c presents the experimental results of PSA detection, showing
a negative correlation between dimer percentage and PSA
concentrations. The decrease of dimer percentage with
increased biomarker concentration can be explained by the
fact that the high concentrations of PSA will block the anti-PSA
sites on the PS beads, thus preventing particle aggregations.
Additional evidence is that the dimer percentage even dropped
to a value close to the negative control for the highest
concentration used in the experiment. It means that high
concentrations of protein would block the particle aggregation
process and increase the ratio of individual microbeads. As
lower PSA concentration leads to more aggregations in Figure
4c, it is possible to reduce the LOD of PSA detection. To
further study the LOD of this method, PSA with concentrations
of 367 and 36.7 fM was respectively tested following the same
process. The results are plotted in Figure 4d, which shows a
non-monotonous response curve of the relationship between
dimer percentage and PSA concentration. The result indicates
that the microbeads aggregation system can achieve a rather
low LOD. In practical applications, the clinical concentration of
PSA is about 10 ng/mL (corresponding to 0.3 nM),
48,49
which
is shown as the shaded range in Figure 4c. For real diagnosis
applications, the target sample and a diluted solution of the
target sample can both be measured, and the dierence of
aggregation ratio between the two experimental groups can be
used to determine the relevant concentration region of the
response curve.
In order to test the practicality of this method for clinical
applications, experiments in serum have also been performed to
detect PSA with concentrations of 0, 2, 5, 25, and 100 ng/mL
(corresponding to 0, 0.0588, 0.147, 0.735, and 2.94 nM). The
experiments were conducted following the detection procedure
as buer samples. For each case, PSA serum solution was added
into and uniformly mixed with the anti-PSA functionalized
microbeads suspension. Figure 5 presents the results of the
Figure 3. (a) Four cases of the trapped microbeads within one
microwell. (b) Local rotation of the microbeads. (c) Principle of
rotation to identify the false dimer and the imaging algorithm to
distinguish and count the single bead and dimers arrayed within the
MW-MFC system. The diameter of the microwells in (a)(c) is 20
μm.
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serum test, which reveals a relationship consistent with that in
buer solution as shown in Figure 4. The microbeads
aggregation due to the beads interactions and nonspecic
serum protein adsorption is calculated to be 4.41%, which is a
little higher than that in buer solutions. The linear relationship
between dimer percentage and PSA concentration demon-
strates the great potential of the microbeads aggregation
strategy for biomarker quantications in clinical diagnostics.
DISCUSSION
To understand the special behaviors of the formation of the
dimers, the contact probability and the aggregation kinetics of
the microbeads were further studied. For microbeads,
aggregation is less ecient compared with nanoparticles
because of their dierent diusion rate and size. As Brownian
diusion for the number of micro/nanoparticles scales with 1/
R2, the diusion of microbeads is much slower compared with
nanoparticles,
44
which prevents the aggregation of the microbe-
ads.
As Figure 6a shows, PSA in the solution is primarily binding
with anti-PSA and adsorbed onto the surface of the microbeads
at the original period
50
(this reaction step is named as process
I) which follows eq 5:
+⇌
A
RAR()
K
K
11
off
on
(5)
where A1and Rare molecule capture sites and target
biomarker, respectively. In process I, anti-PSA on the
microbeads surface works as the capture molecule (A1)to
interact with PSA (R) in the bulk solution. Kon and Koff are the
association and dissociation rate of PSA and anti-PSA
interaction, and they dene the equilibrium dissociation
constant (KD).
==KK
K
Nr
D
Don
off
s
(6)
Figure 4. (a) Microbeads size distribution and (b) microbeads aggregations measured via a commercial counter. Experimental results of PSA
detection using MW-MFC: (c) negative corrective relationship between dimer percentage and PSA concentration from 3.67 pM to 7.35 nM,
meanwhile the concerned concentration in diagnostic application is in the range from 30 to 300 pM. (d) Response curve of PSA detection using
microbeads aggregation method.
Figure 5. Experimental results of PSA serum detections. PSA serum
with concentrations in the region that is relevant in the clinical
diagnostics is detected, showing a negative corrective relationship.
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436
where Nsis the minimum number of molecules to be captured
for detection, Dis the diusion coecient of the molecules, and
ris the radius of micro/nanoscale matters. For a simplied
discussion, the molecule and microbeads are all considered as
spheres in the following analysis. The number of captured PSA
(N) by anti-PSA modied microbeads
50,51
can be calculated
with eq 7:
ρ=−
N
tKNNKN
d
d()
on 00off (7)
where N0is the real density of the receptor molecules (anti-
PSA on the microbead surface) and ρ0is the bulk concentration
of analyte molecule (PSA). The diusion of analyte molecules
obeys the following eq 8:
ρ
ρ
=∇
N
tD
d
d
2
(8)
As Figure 6b shows, PSA captured on a microbead can react
with residual anti-PSAs immobilized on an adjacent microbead
to generate a dimer (process II). For microbeads aggregation,
the interaction of PSA coated microbeads (A1R) and anti-PSA
modied microbeads (A1) can be expressed as eq 9:
+
HIoooAR A A R
(
)()
K
K
111
2
off
on
(9)
where A1Ris the PSA coated microbeads formed in the process
Iand A12Ris the aggregated microbeads formed in the process
II. (A1R) and (A12R) represent the number of anti-PSA
modied microbeads and PSA coated microbeads, respectively.
The aggregation ratio (β) of the microbeads in the above
process can be described with HeidelbergerKendall (HK)
model.
52
β
=bSARAf5.2 4 ( )( )
hh1 1
3/2 (10)
where bhis the steric hindrance coecient for available binding
sites on one PS bead bound to another bead (0 < bh< 1), and
Shis the corresponding steric hindrance coecient for available
binding sites on one PSA molecule bound to the anti-PSA
modied PS beads (0 < Sh< 1). fis the surface fraction
occupied by one binding site on the particle given by
=
b
R4
2
2(11)
where bis the radius of a circular binding site of the receptor
(PSA-binding site) and Ris the radius of the particle.
For three beads aggregation, the forming ratio is β2. And for
three more microbeads aggregation, the ratio should be higher
order (β3,β4, ...) of the dimers because they are formed on the
basis of dimers. As revealed by the experimental results in
Figure 5, the aggregation ratio is low, generally lower than 11%.
Therefore, the generation of three or three more aggregations is
more dicult. In fact, there is no three or three more
aggregations observed or detected in the MW-MFC system.
Combining eqs 10 and 11, the relationship between
processes I and II can be further understood to elucidate the
dierence between microbeads aggregation and nanoparticle or
molecular interactions. In the experiment of PSA detection,
when PSA is mixed with anti-PSA modied PS beads, processes
I and II would take place. The value of bhand Shis assumed to
be constant. The aggregation ratio of the two processes will be
dierent as it is determined by the radius of the bounded
particles or molecules from eqs 10 and 11.Thus,the
relationship between processes I and II can be described as
in Figure 6c, in which particle size is a critical factor, revealing
the special features of microbeads aggregation strategy. In
addition, the sum value of (A1R) and (A1) equals the number of
coated anti-PSA sites on the microbeads surface; thus, the
maximum aggregation ratio obtained by eq 10 would occur
when the number of A1equals the A1R, corresponding to the
concentration locating at the peak curve in Figure 4b.
As shown in Figure 6c, the area of the sharing region, C,is
used to represent the relationship between processes I and II.
For nanoparticle aggregations, the area of Cis comparable to
the size of the biomolecules. With increase in particle size, C
decreases as the aggregation ratio scales with R3.For
microbeads with a diameter of 5 μm, the area of Cwould be
very small and even close to zero, indicating that the reaction
process will be divided into two steps as described above. In
process II, aggregation of the microbeads takes place through
the interactions between the adsorbed PSA on microbeads and
the residual anti-PSA coated on other microbeads. At low PSA
concentration, there exist many residual anti-PSA sites on the
microbeads, while less PSA is bound to the particle for the
aggregation reaction. On the contrary, most of the anti- PSA
sites would be blocked in cases of high PSA concentrations, and
few anti-PSA binding sites remain that can allow for further
formation of the aggregates which mainly occur in process II.
This case is similar to the low PSA concentration, and
aggregation process is highly blocked as the number of PSA
sites and anti-PSA sites deviate far from maximum concen-
tration predicted by eq 11.
Equation 10 reveals that improving the concentration of
microbeads (corresponding to (A1R) and (A12R)ineq 10
would contribute to forming more aggregated dimers in process
II. Because of the limited diusion of microbeads, higher beads
concentration in conned space leads to higher contact
probability, which further contributes to generating more
aggregated microbeads. Microbeads of dierent concentrations
diluted from the original microbeads suspension by 20, 100,
200, and 500 times have been tested for PSA detections
Figure 6. Two-step reaction model of microbeads aggregations: (a)
process I indicates the adsorption of PSA onto anti-PSA modied
microbeads, and (b) process II presents the process of microbeads
aggregations. (c) Relationship between aggregation dynamics and
particle size, which reveals that the aggregation process is dependent
on the micro/nanoparticle size.
ACS Sensors Article
DOI: 10.1021/acssensors.7b00866
ACS Sens. 2018, 3, 432440
437
following the same procedure. In the experiment, the nal
concentration of PSA is set as 0.25 μg/mL. The results are
shown in Figure 7, which indicates that higher microbeads
concentration leads to greater aggregation rate which agrees
with our theoretical analysis. This conclusion is dierent from
the results of nanoparticles,
35
in which nanoparticle concen-
tration seems to have no inuence on the aggregation result,
showing the dierent feature of the microbeads. Besides, the
portable MW-MFC system developed here is suitable for
microbeads aggregation experiments using higher beads
concentration due to its inherent advantages of high
throughput and hydrodynamic control.
CONCLUSIONS
In this work, we have demonstrated a cellphone-enabled
portable particle analysis system for biomarker detection based
on the microbeads aggregation strategy. With the MW-MFC,
aggregated microbeads are trapped and arrayed in a high
throughput and hydrodynamically controllable way. The bead
aggregation status can be simply analyzed by processing the
images with a custom-built MATLAB program or using the
developed cellphone APP directly. The detection limit of PSA
is demonstrated to be as low as 0.125 ng/mL (3.67 pM) in
buer, satisfying the clinic diagnostics demands. Moreover, the
clinically relevant range for PSA was measured in serum with a
linear relationship. The developed new platform does not
require any bulky and expensive optical or complex electrical
setup, thereby making immunoassay even cheaper and suitable
for point-of-care applications. Signicantly, the behaviors of
protein induced microbeads aggregation have been carefully
studied, and a two-step kinetics model has been developed to
explain the response curve of PSA detection, with which the
kinetics of microbeads aggregation process can be gured out.
The proposed microbeads aggregation strategy and measure-
ment system shows great potential for a wide range of
biomarker detections.
ASSOCIATED CONTENT
*
SSupporting Information
The Supporting Information is available free of charge on the
ACS Publications website at DOI: 10.1021/acssen-
sors.7b00866.
Structure of the portable microscope system; picture
processing with the cellphone-enabled application; the
algorithm used in the cellphone application (PDF)
AUTHOR INFORMATION
Corresponding Author
*E-mail: xduan@tju.edu.cn.
ORCID
Luye Mu: 0000-0002-6810-7598
Xuexin Duan: 0000-0002-7550-3951
Author Contributions
§
W.C. and M.H. contributed equally.
Notes
The authors declare no competing nancial interest.
ACKNOWLEDGMENTS
The authors gratefully acknowledge nancial support from the
Natural Science Foundation of China (NSFC No. 61176106,
91743110), and the 111 Project (B07014).
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Supplementary resource (1)

Data
January 2018
Weiwei Cui · Meihang He · Luye Mu · Zuzeng Lin · Xuexin Duan
... At date, several smartphone-based microscope sensors have been used for biomarker diagnosis and they can be divided into bright field microscope and fluorescence microscope (see Table 11.4). For example, Cui et al. developed a portable smartphone-based brightfield microscope for the diagnosis of clinical biomarkers using a PoC platform based on an aggregation assay (Fig. 11.6C) (Cui et al., 2018). Polystirene (PS) microbeads were functionalized with antibodies and then integrated in a microfluidic device. ...
... The device is able to detect PSA with a LOD of 0.125 ng mL 21 (3.67 pM) thus allowing its clinical quantification. The final platform is portable, cost-effective, high throughput and label-free displaying a great potential for diagnosis of other biomarkers (Cui et al., 2018). Wang's group proposed a brightfield smartphone imaging device for the detection of protein biomarkers using a simple microbubbling digital assay and magnetic enrichment technology (Fig. 11.6D). ...
... (C) A portable smartphone PoC biosensor based on microwell-microfluidic chip and microbead aggregation assay. Scheme of the microbeads aggregation process (top); Statistics of the aggregated microbeads acquired by processing the photographs acquired by cellphone (Cui et al., 2018). (D) A smartphone-based bright-field microscopy biosensor for quantitation of femtomolar-level protein biomarkers detection integrating a simple microbubbling digital assay. ...
Chapter
Biomarkers are nucleic acids, peptides, proteins, lipids metabolites, or other small molecules in human tissues or biological fluids whose accurate detection contributes to the prediction and determination of diseases and their status. In recent years, platforms that allow the detection of important biomarkers at the point-of-care (PoC) are moving the field of diagnostics towards personalized medicine. In order to comply with the REASSURED criteria stated by the WHO, the most efficient solution is to integrate PoC devices with smartphones. Without a doubt the smartphone camera is a “smart detector,” and almost all the optical-based methods have been integrated, including absorbance, fluorescence, microscopic bio-imaging, surface plasmon resonance, chemiluminescence, bioluminescence, and photoluminescence. In this chapter, we explain how smartphones can be used as smart detectors in diagnostic devices and we will provide an overview of recent developments of smartphone-based optical PoC devices.
... Microbeads have also long been studied as a labeling material that can detect target molecules easily and quickly. (11)(12)(13) Receptor molecules that react with the target molecules are fixed on the surface of the microbeads or the substrate, and the reaction with the target molecules results in the formation of a specific bond between the beads or between the beads and the substrate. The detection and quantitative evaluation of the target molecules can be realized by observing the microbeads bonded to the substrate or the bead aggregates with optical technology or the naked eye. ...
... (12,18,19) However, even when such a method is introduced, trial and error is necessary to find the optimal conditions such as the modification density and the length of the molecular chain to obtain a sufficient effect. (20) To evaluate the effect of such modification, an evaluation system that is similar to an actual assay was constructed, and a magnetic force (12) or a shear force imposed by a fluid (13) was applied to estimate the non-specific interaction. In this method, since it is necessary to construct an experimental system using an electromagnet or a liquid feed pump in every test, the operation becomes complicated and systematic examination is difficult. ...
... Herein, we will highlight several of the most commonly used classes, including microfluidic arrays, droplet microfluidics, microfluidic paper-based devices, and microfluidic slip-driven devices. Other microfluidic devices such as continuous flow chip [14,15], microwell-based platform [16,17], and self-powered microfluidic chip [18][19][20] have also been reported to be applied for multiplex assays and readers could refer to more complete descriptions elsewhere [21]. ...
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... The simplicity of the assays has ensured their wide distribution. The simplicity of the detection of aggregation processes allows implementing simple methods of on-site testing on their basis, for example, using cameras of mobile phones [4]. However, the sensitivity of aggregation assays is, in most cases, inferior to other bioanalytical approaches. ...
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This paper reports a surface functionalization strategy for protein detections based on biotin-derivatized poly(L-lysine)-grafted oligo-ethylene glycol (PLL-g-OEGx-Biotin) copolymers. Such strategy can be used to attach the biomolecule receptors in a reproducible way simply by incubation of the transducer element in a solution containing such copolymers which largely facilitated the sensor functionalization at an industrial scale. As the synthesized copolymers are cationic in physiology pH, surface biotinylation can be easily achieved via electrostatic adsorption on negatively charged sensor surface. Biotinylated receptors can be subsequently attached through well-defined biotin-streptavidin interaction. In this work, the bioactive sensor surfaces were applied for mouse IgG and prostate specific antigen (PSA) detections using quartz crystal microbalance (QCM), optical sensor (BioLayer Interferometry) and conventional ELISA test (colorimetry). A limit of detection (LOD) of 0.5 nM was achieved for PSA detections both in HEPES buffer and serum dilutions in ELISA tests. The synthesized PLL-g-OEGx-Biotin copolymers with different OEG chain length were also compared for their biosensing performance. Moreover, the surface regeneration was achieved by pH stimulation to remove the copolymers and the bonded analytes, while maintaining the sensor reusability as well. Thus, the developed PLL-g-OEGx-Biotin surface assembling strategy is believed to be a versatile surface coating method for protein detections with multi-sensor compatibility.
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Point-of-care (POC) molecular diagnostics plays a pivotal role for the prevention and treatment of infectious diseases. In spite of recent advancement in microfluidic based POC devices, there are still rooms for development to realize rapid, automatic and cost-effective sample-to-result genetic analysis. In this study, we propose an integrated rotary microfluidic system that is capable of performing glass microbead based DNA extraction, loop mediated isothermal amplification (LAMP), and colorimetric lateral flow strip based detection in a sequential manner with an optimized microfluidic design and a rotational speed control. Rotation direction-dependent coriolis force and siphon valving structures enable us to perform the fluidic control and metering, and the use of the lateral flow strip as a detection method renders all the analytical processes for nucleic acid test simplified and integrated without the need of expensive instruments or human intervention. As a proof of concept for point-of-care DNA diagnostics, we identified the food-borne bacterial pathogen which was contaminated in water or milk. Not only monoplex Salmonella Typhimurium but also multiplex Salmonella Typhimurium and Vibrio parahaemolyticus were analysed on the integrated rotary genetic analysis microsystem with a limit of detection of 50 CFU in 80min. In addition, three multiple samples were simultaneously analysed on a single device. The sample-to-result capability of the proposed microdevice provides great usefulness in the fields of clinical diagnostics, food safety and environment monitoring.
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The advancement in micro/nanotechnologies has been continuously providing possibilities for inventing novel biochemical sensors. However, variations in the transducer type can cause different sensing results due to the differences in their mechanisms of analyzing biomolecular interactions. In this work, we focused on the comparative analysis of static and non-static assays for molecular interactions using on-chip integrated extended-gate field effect transistor (EGFET) as a static sensing interface and solidly mounted resonator (SMR) as a non-static sensing interface. Analysis of polyelectrolytes (PETs) surface assembly and antigen-antibody interaction using the two types of biochemical sensors presented consistent and complementary sets of information. Meanwhile, due to the difference in their operating mechanisms, variations on the detection efficiency, kinetics and thermodynamics were observed. Our results highlighted the critical dependence of signal detection on biochemical sensors’ operating mechanisms and provided a valuable guidance for static and non-static assays for biomolecular detections.
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In this review we discuss recent developments on the use of mobile phones and similar devices for biosensing applications in which diagnostics and communications are coupled. Owing to the capabilities of mobile phones (their cameras, connectivity, portability, etc.) and to advances in biosensing, the coupling of these two technologies is enabling portable and user-friendly analytical devices. Any user can now perform quick, robust and easy (bio)assays anywhere and at any time. Among the most widely reported of such devices are paper-based platforms. Herein we provide an overview of a broad range of biosensing possibilities, from optical to electrochemical measurements; explore the various reported designs for adapters; and consider future opportunities for this technology in fields such as health diagnostics, safety & security, and environment monitoring.
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Pathogen detection has traditionally been accomplished by utilizing methods such as cell culture, immunoassays, and nucleic acid amplification tests; however, these methods are not easily implemented in resource-limited settings because special equipment for detection and thermal cycling is often required. In this study, we present a magnetic bead aggregation assay coupled to an inexpensive microfluidic fabrication technique that allows for cell phone detection and analysis of a notable pathogen in less than one hour. Detection is achieved through the use of a custom-built system that allows for fluid flow control via centrifugal force, as well as manipulation of magnetic beads with an adjustable rotating magnetic field. Cell phone image capture and analysis is housed in a 3D-printed case with LED backlighting and a lid-mounted Android phone. A custom-written application (app.) is employed to interrogate images for the extent of aggregation present following loop-mediated isothermal amplification (LAMP) coupled to product-inhibited bead aggregation (PiBA) for detection of target sequences. Clostridium difficile is a pathogen of increasing interest due to its causative role in intestinal infections following antibiotic treatment, and was therefore chosen as the pathogen of interest in the present study to demonstrate the rapid, cost-effective, and sequence-specific detection capabilities of the microfluidic platform described herein.
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This perspective presents recent developments in the application of surface-enhanced Raman spectroscopy (SERS) to biosensing, with a focus on in vivo diagnostics. We describe the concepts and methodologies developed to date and the target analytes that can be detected. We also discuss how SERS has evolved from a ‘point-and-shoot’ stand-alone technique in an analytical chemistry laboratory to an integrated quantitative analytical tool for multimodal imaging diagnostics. Finally, we offer a guide to the future of SERS in the context of clinical diagnostics.
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Point-of-care diagnostics (PoC) and personalised medicine are highly valuable for the improvement of world health. Smartphone PoC platforms which precisely diagnose diseases and track their development through the detection of several bioanalytes represent one of the newest and most exciting advancements towards mass-screening applications. Here we focus on recent advances in both multiplexed and smartphone integrated PoC sensors.
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This study describes a simple and sensitive approach for visual and point-of-care detection of P. aeruginosa and its toxin genes based on multiple loop-mediated isothermal amplification (mLAMP) and lateral flow nucleic acid biosensor (LFNAB). Differentiation of the internal standard gene ecfX and toxin genes (ExoS and ExoU) in P. aeruginosa was determined using FITC-, hex-and digoxin-modified primers in the mLAMP process. In the presence of biotin-and FITC- (hex-, digoxin-) modified primers and Bst DNA polymerase large fragments, the mLAMP produced numerous biotin- and FITC- (hex-, digoxin-) attached duplex DNA products. The products were detected by LFNAB through dual immunoreactions (anti-biotin antibodies on the gold nanoparticle (Au-NP) and biotin on the duplex, anti-FITC (hex, digoxin) antibodies on the LFNAB test line and FITC (hex, digoxin) on the duplex). The accumulation of Au-NPs produced a characteristic red band, enabling visual detection of P. aeruginosa and its toxin genes without instrumentation. After systematic optimization of LFNAB preparation and detecting conditions, the current approach was capable of detecting concentrations as low as 20 CFU/mL P. aeruginosa or its toxin genes within 50 min without complicated instrument, which is more sensitive than PCR. Therefore, this approach provides a simple, pollution free, sensitive, and low-cost point-of-care test for the detection of P. aeruginosa and its toxin genes.