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A flux-adaptable pump-free microfluidics-based self-contained platform for
multiplex cancer biomarker detection
Article in Lab on a Chip · November 2020
DOI: 10.1039/D0LC00944J
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A flux-adaptable pump-free microfluidics-based
self-contained platform for multiplex cancer
biomarker detection†
Bo Dai, ‡a Cheng Yin,‡a Jiandong Wu,‡b Wei Li,a Lulu Zheng,a Francis Lin,c
Xiaodian Han,d Yongfeng Fu,*e Dawei Zhang *a and Songlin Zhuanga
Microfluidics drives technological advancement in point-of-care (POC) bioanalytical diagnostics towards
portability, fast response and low cost. In most microfluidic bioanalytical applications, flowing antigen/
antibody reacts with immobilized antibody/antigen at a constant flux; it is difficult to reach a compromise
to simultaneously realize sufficient time for the antigen–antibody interaction and short time for the entire
assay. Here, we present a pump-free microfluidic chip, in which flow is self-initialized by capillary pumping
and continued by imbibition of a filter paper. Microfluidic units in teardrop shape ensure that flow passes
through the reaction areas at a reduced flux to facilitate the association between antigen and antibody and
speeds up after the reaction areas. By spotting different antibodies into the reaction area, four types of
biomarkers can be measured simultaneously in one microfluidic chip. Moreover, a small-sized instrument
Received 17th September 2020,
Accepted 3rd November 2020
was developed for chemiluminescence detection and signal analysis. The system was validated by testing
four biomarkers of colorectal cancer using plasma samples from patients. The assay took about 20
minutes. The limit of detection is 0.89 ng mL−1, 1.72 ng mL−1, 3.62 U mL−1 and 1.05 U mL−1 for the assays of
DOI: 10.1039/d0lc00944j
carcinoembryonic antigen, alpha-fetoprotein, carbohydrate antigen 125 and carbohydrate antigen 19-9,
respectively. This flux-adaptable and self-contained microfluidic platform is expected to be useful in various
rsc.li/loc
POC disease-monitoring applications.
Introduction
The high incidence and mortality rate of colorectal cancer
(CRC) makes it a huge burden worldwide.1–3 The mortality of
CRC ranks 2nd in the United States and the incidence rate
increased by 1% annually among individuals aged 50 to 64
years and approximately 2% among those aged younger than
50 years, as reported by the American Cancer Society.4 In
China, CRC is the 3rd most commonly diagnosed carcinoma
a
Engineering Research Center of Optical Instrument and System, The Ministry of
Education, Shanghai Key Laboratory of Modern Optical System, University of
Shanghai for Science and Technology, Shanghai, 200093, China.
E-mail: dwzhang@usst.edu.cn
b
Bionic Sensing and Intelligence Center, Institute of Biomedical and Health
Engineering, Shenzhen Institutes of Advanced Technology, Chinese Academy of
Sciences, Shenzhen 518055, China
c
Department of Physics and Astronomy, University of Manitoba, Winnipeg, MB,
R3T 2N2, Canada
d
Department of Laboratory Medicine, Shanghai Cancer Center, Fudan University,
Shanghai, 200032, China
e
Department of Medical Microbiology and Parasitology, School of Basic Medical
Sciences, Fudan University, Shanghai, 200032, China. E-mail: yffu@fudan.edu.cn
† Electronic supplementary information (ESI) available. See DOI: 10.1039/
d0lc00944j
‡ These authors contributed equally to this work.
This journal is © The Royal Society of Chemistry 2021
and the 5th most common cause of cancer-related death.5
Early diagnosis and personalized treatment of CRC can
improve the survival rate.6,7 Serological testing is particularly
helpful in the evaluation and treatment of CRC, aiding in the
assessment of the degree of malignancy and monitoring of
the advances of therapy as well as providing valuable
prognostic information. A strong relationship between the
clinical and pathological parameters of CRC and a group of
glycoprotein and carbohydrate biomarkers, such as
carcinoembryonic antigen (CEA), carbohydrate antigen 19-9
(CA19-9), carbohydrate antigen 125 (CA125) and carbohydrate
antigen 724 (CA724), has been confirmed in many research
studies.8–10 CEA and CA19-9 have been suggested in the
clinical management for monitoring patients who might have
CRC and have been diagnosed with CRC.11–13
Although each biomarker is correlated with a large
proportion of patients with specific disease characteristics, it
would be less meaningful when used alone as a screening tool
owing to its limited specificity. Not one of the cancer
biomarkers has presented good utility as an independent
predictive biomarker.14 In practice, the use of a single
biomarker is not recommended because the genome
instability of cancers in mutually dependent and highly
complex biological systems would make the assay not
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Paper
reliable.15 Instead, it is widely accepted that multiplex
biomarkers are evaluated simultaneously for a comprehensive
assessment.16
Microfluidics technology drives the development of
bioassays in many aspects, including reduced volume of
reaction system, shortened reaction time, enhanced detection
limit, improved throughput and simplified operation.17–22
Microfluidic chips can be flexibly designed to realize the
detection of multiple biomarkers in immunoassays.23–26
Chemiluminescence immunoassay (CLIA), in which a
chemical probe generates light when reacted with an enzyme
label, is a sensitive, simple approach; no excitation light
sources or emission filters are needed. A microfluidic chip
containing 1536 chambers was demonstrated to quantify 384
biomarkers by measuring luminescence in a microplate
reader.27 To improve the efficiency of the immunoassay for
detecting two prostate cancer biomarkers, a threedimensional (3D) printed chip was fabricated, in which the
sample and reagents could be preloaded, and a
programmable syringe pump was employed for driving the
sample and reagents into a 3D mixer and a detection
chamber according to the assay protocol.28 Furthermore, an
automated microfluidic platform was developed for both
direct sandwich immunoassay (assay for C-reactive protein)
and competitive immunoassay (assay for testosterone). The
injection of preloaded reagents was controlled by six on-chip
valves, a valve actuator and a peristaltic pump.29 In previous
research work, rapid, sensitive and high-accuracy CLIAs for
detecting multiple biomarkers have been successfully
demonstrated. Nevertheless, in these prototypes, a group of
valve and pump systems is required for loading reagents,
sample and wash buffer, making the whole system
complicated.
Passive pumping microfluidics has the potential to
address the problem of the utilization of external bulky
pumping systems.30,31 Usually, passive fluid infusion is based
on gravity-driven or capillary-driven mechanisms. In the
gravity-driven microfluidic chip, the fluid was driven into the
chip by setting a hydraulic head difference between a
reservoir and a drainage.32–35 Viewed in terms of energy, the
infusion is attributed to gravitational potential energy that
requires enough difference in height. Gravity-driven
microfluidic chips are suitable for cell culture applications in
which a large amount of culture medium could be stored in
the reservoirs for long-term cell culture. Microfluidic chips
harnessing capillary action are appropriate for low-volume
applications because capillary force plays a dominant role
and could drive fluid to flow in a tiny space with dimensions
much smaller than the capillary length. Capillary pumpbased microfluidic manipulation has been realized by
designing proper microstructures and controlling the surface
wettability.36–38 Capillary-driven microfluidics have been used
in immunoassays.39–42 However, only a few demonstrated
quantitative multiplex detection of cancer biomarkers.
Besides, microchannels in these microfluidic chips had a
uniform size; thus, the flux of the flow through the entire
144 | Lab Chip, 2021, 21, 143–153
Lab on a Chip
microchannel was fixed. It was not possible to
simultaneously satisfy both strong antigen–antibody
interaction, in which low flux was preferred, and rapid assay,
in which high flux was required.
The behavior of microfluid in the microchannels is
dependent on the geometric structures of the microchannels.
Microchannels with a special design can be developed into a
variety of functional devices. Teardrop-shaped micro-cavities,
which have a simple structure, have been used in many
applications. In a microfluidic reflective multicolor display, a
dot pixel matrix was formed by a sequence of teardropshaped micro-cavities.43 Since the teardrop-shaped microcavities ensured that the water could be retained in the
specific dot pixels, an image could be maintained in the
display using dyed water droplets and air gaps with no energy
consumption. In addition, teardrop-shaped cavities could be
used to trap microbubbles. By oscillating microbubbles
confined in a group of teardrop-shaped cavities upon
acoustic actuation, complex flow patterns could be
generated.44 Teardrop-shaped cavities which were developed
for cell trapping could be used to assemble precise
constellations of cell clusters.45
In this paper, we present a pump-free microfluidic chip
and a CLIA serological analysis platform for measuring
multiple CRC-related biomarkers, including CEA, alphafetoprotein (AFP), CA125 and CA19-9. The initial selftriggered start-up and the following continuous flow of
plasma sample and reagents without using any peripheral
pumping apparatuses make the entire assay simple and
straightforward. The teardrop-shaped design of the reaction
units allows the flow to slow down around the reaction areas
and speed up in the non-reaction areas, which is beneficial
to achieve sufficient antigen–antibody reaction within a short
assay time. The microfluidic platform is validated and
evaluated using clinical samples of patients.
Materials and methods
Chemicals
CEA antigen, AFP antigen, CA125 cancer antigen and CA19-9
cancer antigen, which were used as the standards, were
purchased from BiosPacific Inc., USA. Human sera were
purchased from Sigma-Aldrich, USA (S7023). Capture
antibodies including mouse anti-CEA monoclonal antibody,
mouse anti-AFP monoclonal antibody, mouse anti-CA125
monoclonal antibody and mouse anti-CA19-9 monoclonal
antibody were obtained from BiosPacific Inc. Goat anti-AFP
polyclonal antibody and goat anti-CEA polyclonal antibody as
the detection antibodies were purchased from BiosPacific
Inc. Then, horseradish peroxidase (HRP)-conjugated donkey
anti-goat antibody (Jackson ImmunoResearch Laboratories,
Inc., USA) was used as the secondary antibody against the
detection antibodies of CEA and AFP biomarkers. HRPconjugated mouse anti-CA19-9 monoclonal antibody and
HRP-conjugated mouse anti-CA125 monoclonal antibody as
the detection antibodies against CA19-9 and CA125 antigens
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were purchased from Wason Biotech Inc. Phosphate-buffered
saline (PBS) with 0.05% Tween-20 (PBST) used as the wash
buffer was purchased from Thermo Fisher Scientific. The
response of human serum albumin (HSA) was used as a
positive control in the assay. Mouse monoclonal antibody
against HSA (Santa Cruz Biotechnology, Inc., USA) was used
as the capture antibody. Goat anti-albumin polyclonal
antibody (Sigma-Aldrich, USA) was used as the detection
antibody. The secondary antibody for HSA detection was also
HRP-conjugated donkey anti-goat antibody. Bovine serum
albumin (BSA) was purchased from Solarbia, Co., China.
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Preparation of the standards
Standard samples were prepared by spiking four kinds of
antigens into the human sera (S7023, Sigma-Aldrich, USA)
respectively with known concentration. The gradient
concentration for CEA and AFP standards was from 1 ng
mL−1 to 1 μg mL−1 and the concentration for CA125 and
CA19-9 standards was from 2 U mL−1 to 1250 U mL−1.
Collection of blood samples
Blood samples were collected from 30 patients with colorectal
cancer who were under treatment in Shanghai Cancer Center,
Fudan University. All patients have signed informed consent,
and samples were collected under ethical approval (certificate
no.: 050432-4-1212B). Plasma was collected using procoagulation tubes.
Paper
Pump-free microfluidic chip
Fig. 1 shows the schematic diagram and photo of the pumpfree microfluidic chip. The details of the fabrication process
of the microfluidic chip are described in the ESI† and
illustrated in Fig. S1. 0.5 μL of 0.2 mg mL−1 monoclonal
antibody in protein spotting buffer A (CapitalBio Technology,
China) for each biomarker was first coated on a substrate
whose surface was covalently modified to contain an
aldehyde functional group (OPPolymerSlide™ D, CapitalBio
Technology, China) by using CapitalBio SmartArrayer™ 136
(CapitalBio Technology, China). The CEA, AFP and HSA
capture antibodies were coated on one side which would be
later covered by a microchannel, while the CA125 and CA19-9
antibodies were coated on the other side which would be
covered by another microchannel. The coating area for the
each capture antibody was a spot of 1 mm diameter. Then,
the substrate was soaked in 5% BSA-PBS for 1 hour at room
temperature. The substrate coated with the capture
antibodies was stored at 4 °C for at least 3 months.
Meanwhile, two microchannels were patterned on a silicon
wafer by photolithography. Polydimethylsiloxane (PDMS) with a
mixture ratio of 10 : 1 (elastomer versus curing agent) was used
to cast the pattern and solidified at 80 °C for 4 hours. 4 mm
radius inlet holes were created by punching through the PDMS
at the end of the microchannels. The thickness of the PDMS
layer is 2 mm. The volume of the inlet hole is 100.48 μL. In
each microchannel, there were three successive teardrop-shaped
units. The inlet and the outlet were connected to the wide end
of the first unit and the tapered end of the last unit,
respectively, through two straight channels. Fig. S2† shows the
structure of the microchannel in detail. Finally, the PDMS layer
was placed on the substrate. Each unit was used as a reaction
cavity for one biomarker and the wide end of every unit had a
coating area for each antibody. The outlet hung over the edge of
the substrate and opened downward.
Serological analysis platform (SAP)
Fig. 1 (a) Fabrication process of the pump-free microfluidic chip. (b)
Photo of the microfluidic chip. (c) The flow is driven by the capillary
force and the gravitational force in the initialization stage. (d) The
continuous flow resulted from the suction of the filter paper during
the post-initialization stage.
This journal is © The Royal Society of Chemistry 2021
The design and the photo of the SAP are shown in Fig. 2. An
imaging system was inside the SAP to capture an image of the
reaction products. The image was captured by a monochrome
charge-coupled device (CCD) camera (FLIR Grasshopper®3,
Edmund Optics, USA) via an aspherical lens (effective focal
length = 50 mm) (Edmund Optics, USA) and an aluminum
plane mirror (Edmund Optics, USA). The field of view was 16.2
mm × 20.3 mm. A chip holder was above the imaging system
for holding the microfluidic chip. A disposable tray, well fitting
into the chip holder, was designed to avoid crosscontamination. On one end of the tray, there was a groove filled
with a stack of filter paper (GB002, Whatman, England). When
the microfluidic chip was placed on the tray, the outlets, which
were hung over the edge of the chip, were attached to the filter
paper. The chip holder could be tightly closed by magnetic
attraction. After the chip holder was closed, the imaging system
aimed at the area of six teardrop-shaped units in the dark. The
frame of the SAP was made of aluminum alloy and produced by
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Paper
Fig. 2 (a) Schematic diagram of the serological analysis platform
(SAP). (b) Photo of the platform with the microfluidic chip on the
disposable tray placed on the chip holder.
computer numerical control (CNC) machining. The disposable
tray was produced by 3D printing. The size of the platform was
183 mm × 123 mm × 70 mm and the weight was 1.2 kg.
Assay protocol for detecting cancer biomarkers
The procedure of the assay is shown in Fig. 3 and
demonstrated in Video S1.† In the assay, the microfluidic
chip was placed on the disposable tray, whose groove was
Lab on a Chip
filled with filter paper. The disposable tray was placed on the
chip holder. When loading sample and reagents, the chip
holder was pulled out from the SAP. 20 μL of sample (10 μL
plasma diluted with 10 μL PBS) was added into the inlet of
each microchannel. After about 10 minutes and before the
sample drains away in the inlet, 20 μL of mixed CEA, AFP
and HSA detection antibodies and secondary antibody with a
mixture ratio of 1 : 1 was added to the microchannel for the
assay of CEA, AFP and HSA biomarkers, and 20 μL of CA125
and CA19-9 detection antibody was added to the other
microchannel for the assay of CA125 and CA19-9 biomarkers.
4 minutes later, 10 μL of PBST was added per microchannel
to flush the unbound antibodies away. After 4 minutes, 5 μL
of chemiluminescent peroxidase substrate, i.e. luminol-based
solution (catalog number: CPSOC, Sigma-Aldrich, USA), was
added per microchannel. The imaging system was triggered
after 2 minutes to acquire the image of the reaction products.
The exposure time is set as 6 seconds. Six 20 × 20 pixel blocks
were extracted from the reaction areas in the six units. The
average chemiluminescence intensity was calculated for each
pixel block. The values for cancer biomarkers of interest
(AFP, CEA, CA125 and CA19-9), IBiomarker, were calibrated with
respect to the response of the HSA, IAlbumin, and the
background intensity of the blank unit, IBlank. Whatever the
albumin level varies among patients, the response of the HSA
measured in the microfluidic chip would be similar (Fig.
S3†), because the dose of HSA in the plasma is high46 and
significantly in excess of the amount of the HSA capture
antibody. Thus, the response of the albumin could be used
as a positive control. The resultant response of a specific
cancer biomarker can be expressed as
I Biomarker
Calib
¼ ðI Biomarker − I Blank Þ ×
I Albumin Control
I Albumin − I Blank
(1)
where IAlbumin_Control is the reference intensity obtained in the
preliminary experiment (Fig. S3†). Finally, the concentration
of the biomarkers could be quantitatively figured out based
on standard curves, which were established by the relation of
a set of biomarker concentrations to the corresponding
Fig. 3 Procedure of the assay for detecting multiple biomarkers.
146 | Lab Chip, 2021, 21, 143–153
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chemiluminescence intensity (additional details provided in
the ESI†).
Results and discussion
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Initialization stage: flow triggering
In the pump-free microfluidic chip, the flow experiences two
stages. Initially, the microchannel is empty and the fluid in
the inlet slowly moves towards the outlet through the three
teardrop-shaped units. The motion is driven by the
gravitational force, FGravity, generated by the fluid itself in the
inlet and the capillary force, FCapillary, due to the interaction
between the fluid and the microchannel walls. Once the fluid
reaches the outlet where the filter paper is placed, the fluid is
driven by the capillary suction, FSuction, of the filter paper in
addition to the gravitational force, as shown in Fig. 1c and d.
Fig. 4a shows the self-initialized flow in the successive
teardrop-shaped units. The units are designed with a fixed
volume, i.e. area of 5 mm2 and height of 20 μm, and variable
taper angle, θ. The fluid in the microchannel with a large
taper angle flows faster and reaches the outlet earlier.
In the initialization stage, the fluid is driven by its own
gravitational force as well as the capillary force. The pressure
exerted by gravitational force, i.e. hydrostatic pressure, is
Fig. 4 Analysis of the flow initialization in the microchannel. (a)
Montage of the flow in the microchannel. (b) Contact angles of the
plasma on the PDMS and the substrate. (c) The calculated capillary
pressure along the flow direction in a single teardrop-shaped unit
when the taper angle is different. Inset: the structure of a single
teardrop-shaped unit. (d) The calculated hydraulic resistance of the
microchannels. (e) Time duration for filling up the three units in the
initialization stage. Red curve: theoretical calculation. Dots:
experimental measurement.
This journal is © The Royal Society of Chemistry 2021
Paper
PG = ρgh
(2)
where ρ is the density of the plasma, g is the gravitational
acceleration, and h is the height of the fluid in the inlet.
Furthermore, the Young–Laplace equation can be applied
to describe the capillary pressure over the interface of plasma
and air in the microchannel. The pressure drop in the tearshaped unit with a rectangular cross section can be expressed
as36,47,48
cosðφPDMS Þ þ cosðφSubstrate Þ 2 cosðφPDMS − α ðzÞÞ
þ
PC ¼ γ
h
W ðzÞ
(3)
where γ is the surface tension of the plasma, h is the height
of the microchannel, and W(z) is the position-dependent
width of the microchannel along the flow direction, as
follows:
(
pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi
2 2zr þ z2 þ d; 0 ≤ z ≤ r ð1 þ sinðθÞÞ
W ðz Þ ¼
2ðr − zÞ tanðθÞ þ 2r secðθÞ þ d; r ð1 þ sinðθÞÞ < z ≤ L
(4)
φPDMS and φSubstrate are the contact angles of the plasma on
the PDMS and the substrate. PDMS has a hydrophobic
surface, forming a contact angle, φPDMS, of about 107°, as
shown in Fig. 4b. On the glass substrate, an aldehyde group
is introduced to activate the surface. Therefore, the glass
substrate has a hydrophilic surface and the contact angle,
φSubstrate, is 41°. α(z) is related to the tangential angle of the
side walls, which is
8
< − arcsin r − z ; 0 ≤ z ≤ r ð1 þ sinðθÞÞ
r
α ðz Þ ¼
(5)
:
θ; r ð1 þ sinðθÞÞ < z ≤ L
Fig. 4c shows the capillary pressure along the flow
direction in the unit. The high wettability of the glass
substrate allows the fluid to spread over the bottom surface,
while the hydrophobic characteristics of the PDMS top and
side walls adversely affect the flow of the fluid. Since the
microchannel has a low aspect ratio (height-to-width ratio),
the interaction of the fluid and the top and bottom walls
dominates the capillary pressure. Moreover, the surface
hydrophilicity of the glass substrate is relatively stronger than
the hydrophobic effect on the top PDMS wall. Thus, the
hydrophilicity of the glass substrate provides sufficient
driving force to pull the fluid forward.
Considering the capillary effect as well as hydrostatic
pressure in Hagen–Poiseuille's law, the volumetric flow rate,
Q, of the flow overcoming hydraulic resistance, R, can be
written as
QInitial ¼
P G þ PC̅ ̅
R
(6)
where PC̅ ̅ is the average capillary pressure over the
microchannel, and the hydraulic resistance of the three
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teardrop-shaped units can be estimated as
ðL
12μz
dz
R¼3
3
z¼0 W ðz Þh ð1 − 0:63h=W ðzÞÞ
Lab on a Chip
(7)
where μ is the viscosity of the plasma.
In the calculation, the density and surface tension of the
plasma are 1.06 × 103 kg m−3 and 55.89 × 10−3 N m−1,49
respectively. 20 μL plasma is in the 4 mm radius inlet. The
slight change of the liquid level in the inlet is neglected.
Fig. 4d depicts the hydraulic resistance of the microchannel
with different taper angles. The hydraulic resistance
decreases with the increase of the taper angle because the
microchannel becomes wide and the total length is
shortened.
Then, the volumetric flow rate can be calculated using eqn
(6). Furthermore, the time for filling up the three units from
the moment when the plasma was added into the inlet to the
moment it reaches the filter paper can be estimated. The
estimated time and the measured results are shown in
Fig. 4e. The measured time has the same tendency as that in
the calculation. The deviation between the theoretical
estimation and the measured results could be attributed to
the fabrication error of the microchannels. Fig. S4† shows
the scanning electron microscope (SEM) images of the
microchannels. The height of the microchannel is 16 μm,
which is slightly lower than that in the design. The reduced
height leads to higher hydraulic resistance and lower
volumetric flow rate. As a result, it requires a little bit more
time to fill up the microchannels. Since all the
microchannels were replicated from the patterns on the same
silicon wafer, the height is identical and the consistency can
be ensured.
In the initialization stage, the flow which is free from any
power-driven pump could be self-triggered and fill up the
microchannel rapidly. The time duration for the initialization
in the unit with a large taper angle is short because of high
capillary pressure and low hydraulic resistance. In the
following analysis and the assay, microfluidic chips
consisting of two microchannels with a taper angle of 7° are
used. The dimensions of the microfluidic chips are
illustrated in Fig. S5.† Fig. S4† shows the SEM images of the
chip.
diminishing the attraction of liquid molecules to each other,
i.e. cohesion. As a result, when the cohesive force was weaker
than the adhesive force, the capillary suction of the filter
paper became significant and thus the filter paper imbibed
PBST quickly.
Furthermore, the volumetric flow rate of the plasma and
reagents was evaluated by measuring the time duration for
the imbibition of 20 μL of fluid in the microchannels. 20 μL
of plasma, chemiluminescent peroxidase substrate and PBST
were added into the microchannels. The time duration was
counted from the moment when the flow reached the filter
paper to the moment when they completely drained away in
the microchannels, as depicted in Fig. 5a. The measurement
was repeated four times and plasma samples from six
patients were used.
The volumetric flow rate is proportional to the driving
force of the flow and inversely proportional to the viscosity of
the liquid. Since the plasmas were slightly viscous, the
volumetric flow rate was relatively low, about 1.8 μL min−1.
The volumetric flow rate of the chemiluminescent peroxidase
substrate was 2.3 μL min−1. Low-viscosity PBST, which
underwent a strong capillary suction of the filter paper, had a
high volumetric flow rate of 3.8 μL min−1.
Each reaction cavity was in a teardrop shape, whose nonuniform cross section leads to a variation of flux along the
flow direction, i.e. J(z) = QF S −1(z), where the subscript F
stands for various liquids, i.e. plasma, chemiluminescent
peroxidase substrate and PBST, and S(z) is the cross-sectional
area of the microchannel. The calculated flux for plasma,
chemiluminescent peroxidase substrate and PBST is shown
in Fig. 5b.
The flux is low when the flow passes through a wide end
where reaction happens, while the flux around the tapered
end after the reaction area increases. It has been reported
that the association rate of flowing antigen/antibody with
immobilized antibody/antigen could increase with the
decrease of the volumetric flow rate.50,51 In these previous
studies, the microchannel was straight and had a uniform
cross section along the flow direction. Thus, the flow had a
Post-initialization stage: continuous flow with adaptable flux
When the fluid reaches the outlet, the capillary suction of the
filter paper, FSuction, and the gravitational force, FGravity, of the
fluid in the inlets work together to drive the fluid and the
capillary suction plays a dominant role. The capillary suction
of the filter paper was investigated by dripping 20 μL of
plasma, chemiluminescent peroxidase substrate and PBST on
three pieces of filter paper and measuring the spreading
diameters after 30 seconds. The process of the spreading is
recorded in Video S2.† It is obvious that PBST spread much
faster. It could be attributed to the presence of the surfactant,
i.e. Tween-20. The surfactant reduced the surface tension,
148 | Lab Chip, 2021, 21, 143–153
Fig. 5 (a) Time duration for blotting up 20 μL of plasma,
chemiluminescent peroxidase substrate and PBST from the moment
when the flow reached the filter paper to the moment when they
completely drained away in the microchannels. (b) The calculated flux
along the flow direction in a single teardrop-shaped unit. The shadow
area represents the antibody-coating area.
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Lab on a Chip
not least, the tapered end takes only a small portion of the
reaction cavity, reducing to some extent the size of the units
and avoiding the waste of sample/reagents on the nonreaction areas. The design of the teardrop-shaped unit
ensures a proper condition for the reaction and an efficient
way for the entire assay.
Validation of the assay for multiple-biomarker detection
First, standards of gradient concentrations from 1 ng mL−1 to
1 μg mL−1 (CEA and AFP antigens) and 2 U mL−1 to 1250 U
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constant flux through the entire microchannel. Herein, it is
worth clarifying that it is flux that affects antigen–antibody
interaction. Precisely speaking, low flux is beneficial for
association of flowing antigen/antibody with immobilized
antibody/antigen. Therefore, in the teardrop-shaped unit, the
association between antigen and antibody would be high at
the wide end where the flux drops to the minimum.
In addition, if the width of the reaction area (the wide
end) is fixed, the tapered end, contributing to a relatively
high hydraulic resistance, can be used to control the
volumetric flow rate over the entire microchannel. Last but
Paper
Fig. 6 (a)–(d) Established standard curves, (e)–(h) linear regression of the concentration measured by the microfluidic platform and the
commercial testing and (i)–(l) B&A plots of the difference between the microfluidic platform and the commercial testing. (a), (e) and (i) Assay for
CEA biomarker. (b), (f) and (j) Assay for AFP. (c), (g) and (k) Assay for CA125. (d), (h) and (l) Assay for CA19-9.
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Paper
mL−1 (CA125 and CA19-9 antigens) were tested. The
intensity of the reaction products was read out. The raw
data and the calculated values for the 5 samples are listed
in Table S1.† Standard curves for the four biomarkers were
established by fitting the data to a four-parameter logistic
(4PL) equation, as shown in Fig. 6a–d. The data used for
establishing standard curves were obtained from three
parallel experiments.
Then, as a proof-of-principle test for clinical applications,
30 plasma samples from CRC patients were tested using the
pump-free microfluidic chip and the SAP. Fig. S6† shows the
images of the reaction products captured by the SAP. The
units for albumin present identical brightness, while the
blank units have a dark background. The coating areas in the
units for CEA, AFP, CA125 and CA19-9 emit light with
different brightness. The intensity was recorded and the
concentration of the antigens in the plasma samples could
be derived from the intensity based on the inverse functions
of the corresponding 4PL equations.
The testing results measured by our microfluidic platform
was compared with those obtained in the clinical testing, in
which electrochemiluminescence immunoassays (ECLIAs)
based on commercial testing kits (Elecsys CEA, Elecsys AFP,
Elecsys CA125II and Elecsys CA19-9, Roche Diagnostics
GmbH, Germany) were adopted and measured using a
commercial immunoassay analyzer (Cobas e801, Roche
Diagnostics GmbH, Germany), as shown in Fig. 6e–h. The
coefficient of determination (R 2) for the curve fitting is
higher than 0.9990, implying a highly linear correlation
between the SAP and the commercial immunoassay analyzer.
The results for 5 out of 30 samples are listed in Table S1.†
The difference percentage, defined as the ratio between the
difference and the average of the concentration obtained in
our measurement and the clinical testing,52 was calculated
based on the results for all the 30 samples. The low
difference percentage (<6.6%) confirmed high consistency
between the results obtained in our scheme and the clinical
testing.
Furthermore, Bland–Altman (B&A) analysis was conducted.
All difference data points, except for the case of AFP
biomarker testing (96.67%), are within the limits of
agreement, i.e. δ ± 1.96s, where δ and s are the mean and the
standard deviation of the differences, respectively, as shown
in Fig. 6i–l. The bias between the SAP and the commercial
immunoassay analyzer is less than 2.23%. The B&A analysis
indicates that the CLIAs for CEA, AFP, CA125 and CA19-9
multiple biomarkers conducted in our microfluidic platform
have a perfect agreement with the ECLIAs in the clinical
testing.
Reproducibility and sensitivity of the assay
The reproducibility of the assay was evaluated by testing
three samples repeatedly. Each sample was tested in the
three microfluidic chips three times. The standard deviation
was no more than 0.034, 0.025, 0.034, and 0.031 for CEA,
150 | Lab Chip, 2021, 21, 143–153
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AFP, CA125 and CA19-9, respectively (Table S2†), indicating
that the reproducibility of the assay could be guaranteed.
The limit of detection of the microfluidic platform for the
assays of CEA, AFP, CA125 and CA19-9 is 0.89 ng mL−1, 1.72
ng mL−1, 3.62 U mL−1 and 1.05 U mL−1, respectively.
Additional details about the calculation of the 4PL equation
and limit of detection are provided in the ESI.†
Influence of the coating order on the assay
In the pump-free microfluidic chip, a multiplex assay for
different biomarkers is conducted simultaneously in the
connected units. The flow is unidirectional from the inlet to
the outlet via the three reaction areas in series. Since the
capture antibodies immobilized on the chip are all
monoclonal antibodies, they are solely specific to the
antigens of interest. To verify the irrelevance of the
sequential order to the assay, the coating positions for
immobilizing
capture
antibodies
were
swapped.
Differentiated from the coating sequence as illustrated in
Fig. 3, two microfluidic chips were specially prepared with
capture antibodies for AFP and CA19-9 biomarkers
immobilized in the units next to the inlets and capture
antibodies for CEA and CA125 biomarkers immobilized in
the middle units. The coating position for HSA was
unchanged and the amount of each capture antibody
immobilized on the chip remained the same. Two plasma
samples were tested. The procedure of the entire assay
followed the same protocol. Fig. S7† shows the captured
images of the reaction products for the two samples tested in
the microfluidic chips using different coating strategies and
Table S3† lists the concentration measured in the assays. The
almost identical results obtained from the microfluidic chips
with different coating sequence indicate that the antigen–
antibody reaction is specific and not affected by the
coexistence of other irrelevant antibody molecules in the
reagents. Therefore, the simultaneous multiplex assay of
cancer biomarkers in the connected units is feasible and the
sequential order has no influence on the assay.
Conclusions
We have developed a flux-adaptable and self-contained
microfluidic platform including a pump-free microfluidic
chip on which CLIAs for multiple biomarkers can be
conducted simultaneously and a SAP in which the
chemiluminescence emission of reaction products can be
detected. In the immunoassay, the flow of the samples and
reagents can be triggered by gravitational force and capillary
force and continued by the suction of the filter paper. The
design of the teardrop-shaped units contributes to retardant
flow with low flux through the reaction area and a timesaving process for the entire assay. We have demonstrated
the CLIAs of the four representative CRC biomarkers by using
the microfluidic platform. Only 20 μL plasma per clinical
sample is used. The entire assay is easy to operate and costeffective. The assay can be simply accomplished within 20
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Lab on a Chip
minutes. The total assay cost per sample testing for the
quantitative multiplex detection of the four cancer
biomarkers is US$ 8.84 (Table S4†). The outcomes of the
testing for 30 patient samples have a good agreement
with those obtained in the clinical testing that was based
on commercial testing kits and instruments. The platform
can be further developed to include enhanced features.
Clocked control schemes for reagent loading and image
acquisition are desired in order to realize an automated
assay. Besides, in an improved version, a dense
arrangement of the reaction units and a large field of
view of the imaging system are necessary to achieve
multiple-sample assay. In addition, the testing cost could
be further reduced by bulk purchasing in the mass
production. The high-efficiency microfluidic platform has
proven to be a promising candidate in the field of
serological immunoassays.
Paper
5
6
7
8
9
Author contributions
Conceptualization: B. D., L. Z., Y. F., D. Z.; project
administration: B. D., F. L., Y. F., D. Z., S. Z.; methodology: B.
D., J. W., L. Z., Y. F. ; data curation: B. D., C. Y., W. L., X. H.,
Y. F.; formal analysis: B. D., C. Y., W. L., L. Z., Y. F.;
investigation: B. D., J. W., L. Z., Y. F.; visualization: C. Y., W.
L.; validation: B. D., J. W., L. Z., Y. F.; writing – original draft:
B. D., Y. F.; writing – review and editing: B. D., J. W., L. Z., F.
L., Y. F., D. Z., S. Z.; funding acquisition: B. D., Y. F., D. Z.
10
11
Conflicts of interest
There are no conflicts to declare.
Acknowledgements
The work is financially funded by the National Key Research
and Development Program of China (2016YFD0500604,
2016YFD0500603), the National Natural Science Foundation
of China (61775140), the Shanghai Science and Technology
Commission (18142200800) and the Shanghai Rising-Star
Program (20QA1407000).
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