RESEARCH PAPER

Miniaturization of immunoassay by using a novel module-levelimmunosensor with polyaniline-modified nanoprobesthat incorporate impedance sensing and paper-based sampling

Cheng-Hsin Chuang • Yuan-Chu Yu •

Da-Huei Lee • Ting-Feng Wu • Cheng-Ho Chen •

Shih-Min Chen • Hsun-Pei Wu • Yao-Wei Huang

Received: 25 September 2013 / Accepted: 13 February 2014

� Springer-Verlag Berlin Heidelberg 2014

Abstract This work presents a module-level impedance

measurement system integrated with a disposable immu-

nosensor for the immunoassay of bladder cancer cell lysate

(T24) to a specific antibody (galectin-1). The immuno-

sensor consisted of a flexible printed circuit patterned with

an interdigital microelectrode array which immobilized

polyaniline-modified nanoprobes on an electrode surface

by dielectrophoresis. A quantitative sampling of cell lysate

without a pump was made by using paper as the cell lysate

carrier and sweeping a moistened paper over the sensing

area of interdigital microelectrode array for sampling. In

this study, the impedance measurement results of the

module-level system were compared with those measured

by the precision LCR meter, in which the error is \2 %.

Additionally, the normalized impedance variation in im-

munosensing linearly increased with the cell lysate con-

centration. With a sensitivity based on a normalized

impedance variation of 124.4 % per mg/ml, the immuno-

sensor can rapidly detect the lowest concentration of cell

lysate for 0.0626 mg/ml in 10 min. Therefore, this work

has demonstrated the accuracy of the module-level

immunosensor as well as the reliability of impedance-

based sensing for bladder cancer cell lysate. The proposed

disposable sensor and portable impedance system module

are highly promising for use in point-of-care diagnostics.

Keywords Dielectrophoresis � Polyaniline �Immunosensor � Impedance sensing

1 Introduction

Bladder cancer ranks as the sixth most common cancer in

industrialized countries. According to the Urology Care

Foundation, in 2013, 72,570 Americans were diagnosed

with bladder cancer and 15,210 died of the disease. Recent

decades have witnessed a gradual increase in the incidence

of bladder cancer. Although up to 75–80 % of new cases

are diagnosed as a non-invasive (pathological stage Ta),

stroma invasive (T1) or carcinoma in situ (Tis) disease, the

remaining 20–25 % of tumors are presented as muscle

invasive or more advanced disease (T2–4) with a poor

prognosis. Moreover, although approximately 20 % of Ta

and T1 tumors are cured, after an initial removal, 60–70 %

of those tumors recur at least once in 5 years while

10–20 % progress to muscle invasive cancer (Jemal et al.

2010). During the diagnosis of bladder cancer, the grade

must be identified when deciding upon the cancer treat-

ment. Although cystoscopy is the most effective means of

examining the grade of bladder cancer from a biopsy of the

bladder lining, patients may need an anesthesia during this

procedure. Therefore, despite the need to routinely

administer cystoscopy in order to detect the recurrence of

bladder cancer in patients, the subsequent clinical burden

C.-H. Chuang (&) � Y.-C. Yu � S.-M. Chen � H.-P. Wu �Y.-W. Huang

Department of Mechanical Engineering, Southern Taiwan

University of Science and Technology, Tainan 71005, Taiwan

e-mail: chchuang@mail.stust.edu.tw

D.-H. Lee

Department of Electrical Engineering, Southern Taiwan

University of Science and Technology, Tainan 71005, Taiwan

T.-F. Wu

Department of Biotechnology, Southern Taiwan University of

Science and Technology, Tainan 71005, Taiwan

C.-H. Chen

Department of Chemical and Materials Engineering, Southern

Taiwan University of Science and Technology, Tainan 71005,

Taiwan

123

Microfluid Nanofluid

DOI 10.1007/s10404-014-1364-4

on patients is of priority concern. Besides cystoscopy, a

few urinary markers have been developed to reduce the

frequency of cystoscopy (Cheung et al. 2013). However,

their sensitivity and selectivity remain unsatisfactory,

necessitating the development of a highly accurate, non-

invasive and in vitro method for point-of-care diagnostics

of bladder cancer to help postoperative patients.

Using a few biomarkers for bladder cancer, recent works

have verified the relevance of their expression levels to the

prognosis of recurrence, such as annexin 1 (Li et al. 2010),

lactate dehydrogenase B (LDH-B) (Liao et al. 2011),

galectin-1 (Memon et al. 2005). Based on the expression

level, doctors can decide whether cystoscopy is necessary

as a routine check for postoperative patients. However, the

protein expression is usually determined by immunohisto-

chemistry (IHC), Western blotting or enzyme-linked

immunosorbent assay (ELISA). Despite the widespread use

of these analytical methods to detect specific proteins in the

given sample of tissue homogenate or extract, their pro-

cedures require both complex sample preparation and

instrumentation. Alternatively, microfluidic devices have

received considerable attention for use in immunoassay,

owing to their desirable features such as a low quantity of

the required sample, rapid response, low fabrication cost

and addressable sensing area for array manner. The most

widely used detection method is fluorescence, followed by

electrochemistry. Theoretically, the immunosensor sur-

faces are immobilized with antibodies; the specific proteins

thus bind with antibodies based on antibody–antigen

interaction. Once the proteins are labeled with fluorescent

dye, the fluorescence response can accurately reflect the

expression level (Song et al. 2011; Seo et al. 2011).

However, for label-free detection, the immunoreaction can

be measured by cyclic voltammetry (CV) and electro-

chemical impedance spectroscopy (EIS) (Samanman et al.

2012; Moreira et al. 2013; Lin et al. 2013). Despite mini-

aturization of the immunosensor, the measurement instru-

ments (e.g., fluorescent microscopy, electrochemical

analyzer and impedance analyzer) are still bulky and

expensive for point-of-care purpose. This work develops a

novel module system of impedance measurement to

achieve a portable size and reliable signal processing for

immunoassay. An attempt is also made to immobilize

antibodies on the immunosensor by using dielectrophoresis

(DEP) in order to condense the nanoprobes on the surface

of interdigital microelectrodes array (IDMA). The nanop-

robes are coated with conductive polymer (polyaniline) and

conjugated with specific antibodies, galectine-1, for

immunoassay of bladder cancer cell lysate (T24).

Polyaniline (PANI) is an important member of intrin-

sically conducting polymer. Chemical oxidization and

electrochemical synthesis are two major routes for syn-

thesizing PANI (Pouget et al. 1992). Among the

advantages of PANI over other conducting polymers

include ease of synthesis, low cost, high environmental

stability, a unique doping/dedoping mechanism, and

physical properties that can be controlled by both oxida-

tion and protonation states. Correspondingly, related

works have demonstrated the feasibility of using NPs

coated with PANI in immunoassay (Yuk et al. 2009; Gu

et al. 2009). The synthesis of PANI typically requires a

strong protonic acid as the dopant to yield a moderately

high conductivity of up to 100 s/cm. However, residual

strong protonic acid in PANI can decrease the pH value

and degrade protein during immunoassay. Hence, in this

work, a polyaniline called PANDB is synthesized using a

weak protonic acid, dodecyl benzene sulfonic acid

(DBSA) as the dopant (Chuang et al. 2013). As a bulky

molecule containing a hydrophilic head (sulfonic acid) and

a long hydrophobic chain (–C12H25), DBSA functions both

as a surfactant and dopant. The sulfonic acid group

interacts strongly with the polyaniline backbone. The

strong interaction is sufficient to prevent the DBSA mol-

ecule washed out from the polyaniline backbone by the

aqueous solution. Therefore, PANDB is a promising

electrical material for biomedical applications, owing to

its thermal and chemical stability in the doped form. As is

well known, the conductive polymer coated on the nano-

particles yields a porous microstructure as nanoprobes

immobilized on the electrode surface and reduces the base

line of initial impedance of IDMA. Therefore, as is

expected, the modification scheme increases the sensitivity

of immunosensor. Additionally, an attempt is made to

achieve an inexpensive and disposable immunosensor

without a pumping device by designing an open sensing

area in order to avoid the complex fabrication processes of

a microchannel and flow chamber. Although micropipette

is the most common sampling procedure in the laboratory,

it is infeasible for patient use at home. Therefore, based on

a relatively easy approach in which paper is used as the

sample carrier, this work presents a novel immunosensor

based on DEP trapping of nanoprobes. Moreover, accu-

racy of the proposed module-level impedance measure-

ment system is verified using the precision LCR meter.

Furthermore, the immunoassay sample is a clinical cell

lysate instead of commercial biomarker kits. Thus, the

feasibility of using point-of-care diagnostics to treat

bladder cancer is demonstrated as well.

2 Theory of dielectrophoresis

Dielectrophoresis was first described by Pohl (1951, 1978).

A dielectric object with a permittivity different from the

surrounding medium situated in a non-uniform electric

field experiences a net force given as follows:

Microfluid Nanofluid

123

FDEP ¼ 2peR3pRe½KðxÞ�rE2

rms ð1Þ

where em is the electrical permittivity of the surrounding

medium; Rp is the radius of the particle; rE2rms

��

�� ¼

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

rE2x þrE2

y þrE2z

q

is the gradient of the square of

applied electric field magnitude; and K(x) is the frequency-

dependent Claussius–Mosotti (CM) factor for a dielectric

uniform sphere. For instance, for a bead, it is expressed as

follows:

KðxÞ ¼e�p � e�me�p þ 2e�m

ð2Þ

where e* is the complex permittivity of the medium (m) or

particle (p) and is defined (by OR as follows:)

e� ¼ e� jrx

ð3Þ

where e is the permittivity of the medium or particle; r is

the conductivity of the medium or particle; x is the

angular frequency; and j is H(-1). Therefore, the CM

factor can be viewed as the ratio of electrical conduc-

tivities between the particle and the medium at a low

frequency; this factor can also be regarded as the ratio of

permittivities between the particle and the medium at a

high frequency. The sign of the CM factor signifies

whether it is a positive DEP (p-DEP) or negative DEP

(n-DEP). When the real part of the CM factor is a positive

value, Re e�p � e�m

.

e�p þ 2e�m

h i

[ 0; particles suspended in

the medium are moved toward the region possessing a

high-intensity electric field by a p-DEP force. Conversely,

when Re e�p � e�m

.

e�p þ 2e�m

h i

\0, the DEP force moves

the particles toward the region possessing a low-intensity

electric field, which is the so-called n-DEP. Furthermore,

when Re e�p � e�m

.

e�p þ 2e�m

h i

¼ 0, the DEP force is equal

to zero, implying that the suspended particles are unaf-

fected by the DEP force; corresponding frequency of the

AC signal is called the cross-over frequency. Thus, in a

non-uniform electric field, the direction of particle

movement depends on the CM factor, and the magnitude

of the DEP force is determined by the imposed electric

field at the particle position, as well as the particle size.

Owing to that the DEP force decreases with particle size,

other factors should be considered when the particle size

shrinks to a nanoscale size such as Brownian motion and

Stokes force. Chuang and Huang (2012, 2013) numeri-

cally and experimentally demonstrated the feasibility of

manipulating nanoparticles. In this work, the positive

DEP force is applied to immobilized biomodified nano-

particles onto the electrode surface.

3 Materials

3.1 Synthesis of PANDB/Al2O3 NPs

PANDB surrounding the Al2O3 NPs was synthesized, owing

to not only the ability of the amino groups (–NH2) of PANDB

to bind to the aldehyde groups (–CHO) of antibodies

(Fig. 1a), but also the ability of the conductivity of PANDB to

reduce the initial impedance of IDMA after PANDB-modi-

fied Al2O3 NPs immobilized on an electrode. The synthesis of

PANDB/Al2O3 NPs is described as follows. In this work,

commercially available aluminum oxide nanoparticles (Al2O3

NPs) from Evonik Degussa Taiwan Ltd. (AEROXIDE� Alu

C, Evonik Degussa), were used, in which their average size

ranges from approximately 20 to 30 nm. PANDB–Al2O3 NPs

were first synthesized using 3 g of aniline monomer, 8.4115 g

of DBSA, 7.35 g of ammonium persulfate (APS) and 2 g of

Al2O3 NPs. Al2O3 NPs were then added to the aniline solution

for 10 min and ultrasonically agitated, followed by the

introduction of the DBSA solution and subsequent stirring for

10 min. Finally, the APS solution was added and the system

was allowed to react for 2 h. PANDB was polymerized by a

standard oxidative polymerization of the aniline monomer,

with APS used as the oxidizing agent. Following polymeri-

zation for 2 h, darkish green PANDB/Al2O3 NPs suspension

was filtered and washed with acetone and reverse deionized

(DI) water until the filtrate became as colorless as the RO

water. Finally, 20 ml of DI water was mixed with powders of

PANDB/Al2O3 to yield an NP suspension. According to

Fig. 1b, average size of the PANDB-coated Al2O3 NPs ran-

ges from 120 to 200 nm, indicating that PANDB encloses and

links Al2O3 NPs into an irregular shape. This morphology can

cause the stacking of porous microstructures on the electrode

surface by DEP force, as well as reduce the impedance due to

the interconnection of conductive PANDB. Hence, the

modification scheme benefits impedance sensing approaches.

3.2 The preparation of the nanoprobes

Antibodies (galectin-1) were conjugated with PANDB/

Al2O3 NPs oxidizing 2 lg of the antibody against galectin-

1 protein for 5 min at room temperature with 1 mM

sodium metaperiodate with the pH value adjusted to 5.5

using 0.1 M sodium acetate. During the oxidizing reaction,

the hydroxyl groups (–OH) of carbohydrate moieties of

antibody were oxidized to aldehyde groups (–CHO), which

subsequently reacted with the amino groups exposed on the

surface of PANDB/Al2O3 NPs. The oxidized antibody was

then mixed with PANDB/Al2O3 NPs with rotation at room

temperature for 30 min. Following conjugation reaction,

antibody-modified PANDB–Al2O3 NPs were collected by

centrifugation at 18,0009g for 5 min; the nanoprobes are

referred to as galectin-1/PANDB/Al2O3 NP.

Microfluid Nanofluid

123

3.3 T24 cell lysate

Widely recognized for its high grade and invasivenes,

human urinary bladder urothelial carcinoma T24 cell was

purchased from Bioresource Collection and Research

Center, Hsinchu, Taiwan, and cultured at 37 �C in

McCoy’s5A [GIBCO (Life Technologies Corporation),

Grand Island, NY, USA], supplemented with 10 % fetal

bovine serum. The T24 cell lysate was harvested and lysed

using mammalian protein extraction buffer (GE Health-

care) according to the manufacturer’s suggestions. The

protein concentration was determined using the Bio-Rad

DC protein assay kit. The original concentration of T24

cell lysate was 0.25 mg/ml, which was collected from T75

flask with *90 % cell confluent. Furthermore, the detec-

tion limit was evaluated using three dilutions of the original

lysate for the assay, e.g., 0.25, 0.125 and 0.0626 mg/ml.

4 Module-level immunosensor

4.1 Immunosensor

The fabrication process is illustrated in Fig. 2, an inter-

digital microelectrode array (IDMA) was patterned on a

3.5 9 2.5 cm2 flexible printed circuit (FPC) by standard

photolithography and wet etching. In order to keep the

flatness of FPC and easy for handling, the FPC was finally

laminated with a polyethylene terephthalate (PET) film for

stiffening the substrate, as shown in Fig. 2a. The number of

IDMA finger pair was 28, and the total area of IDMA was

about 4 by 4 mm square, defined as the sensing area. The

width, gap and length of each electrode were 30, 60 and

4 mm, respectively, as shown in the optical micrograph in

Fig. 2b. A planar non-uniform electric field can be con-

structed by applying an AC signal to the left/right elec-

trode. The nanoprobes therefore move to the electrode edge

where possesses highest density of electric field due to a

positive DEP force. Consequently, a low-cost, light-weight,

disposable immunosensor immobilized with nanoprobes on

IDMA surface was fabricated for immunoassay.

4.2 Module-level system

This work develops a portable, inexpensive and highly

accurate impedance measurement system module for point-

of-care diagnostics based on electrical detection. The pro-

posed system module integrates three components, i.e.,

impedance measurement platform, microcontrol unit

(MCU) and wireless communication. Based on the imple-

mentation of TI CC2530 (Texas Instruments Inc.), an

attempt is also made to control the highly accurate

impedance measurement platform, Analog Device AD5933

(Analog Devices, Inc.), and to construct a wireless com-

munication architecture based on the Zigbee network sys-

tem. Additionally, the proposed our measurement flow

scheme can filter the undesired noise and human influence

by the far-end control (Fig. 3a).

Fig. 1 a Synthesis of

nanoprobes of galectin-1/

PANDB/Al2O3 NPs, b TEM

picture of PANDB/Al2O3 NPs,

c FT-IR analysis of (1) PANDB/

Al2O3 NPs, (2) galectin-1/

PANDB/Al2O3 nanoprobes in

the absence and in the presence

of C–N, C=O and O(=)C–H

Microfluid Nanofluid

123

AD5933, which contains a direct digital synthesis (DDS)

module and a digital Fourier transfer (DFT) module, pro-

vides a highly accurate impedance measurement perfor-

mance. An attempt is also made to provide high control

flexibility of AD5933 using the improved 8-bit MCU in

CC2530 to achieve measurement control on single frequency

and sweep frequency modes through an I2C interface. First,

AD5933 initiates the stimulus signal by the internal 27-bit

DDS module. The impendence response signal is then

sampled by an internal 12-bit analog-to-digital converter

(ADC) module. Additionally, the embedded 1024-point FFT

module calculates all of the impedance results with respect to

each frequency pin. Furthermore, to communicate the final

results through a wireless system, CC2530, which contains

Fig. 2 a Fabrication of

immunosensor with interdigital

microelectrode array (IDMA);

b schematic picture of non-

uniform electrode field of

IDMA and the finished

disposable immunosensor

Fig. 3 a Experimental setup of

immunosensing based on

impedance system module and

paper sampling; b photograph

of module-level impedance

system

Microfluid Nanofluid

123

8-bit MCU and the Zigbee physical layer, is responsible for

the control of impedance measurements (AD5933) and far-

end communication with personal computer (PC). In the

proposed Zigbee network topology, the impedance is mea-

sured by the terminal of an end device, which controls the

AD5933 through I2C interface. As a serial communication

protocol, the I2C standard contains two bidirectional open

drain wires, i.e., SDA and SCL. This standard is a master and

slave-based architecture. Based on the I2C standard, the

master (i.e., end device of CC2530) sends commands to I2C

bus including the start frequency command, incremental

frequency command, cycle numbers command and sweep

frequency number command. The device (AD5933) may

then issue the impedance measurement tasks. Finally, the

frequency responses of impedance results may be sent to an

end device in the form of real and imaginary data parts. When

the terminal of coordinator obtains the impedance results

from the terminal of end device through Zigbee wireless

communication, the coordinator forwards the impedance

results to a PC by the RS-232 interface. Using the LabVIEW-

based platform, the proposed measurement system achieves

the far-end control and computation system in the PC, which

can complete the final frequency response. Importantly, the

large amount of final data can be stored on a far-end server

for further accounting analysis. Therefore, the proposed

measurement system can communicate over an extremely

long distance by CC2530, possibly around 1 K meter, thus

resolving the signal decade problem on a wire-based com-

munication system. Furthermore, the highly accurate results

achieved by AD5933 can be stored in the far-end server for

advance analysis. Above features facilitate point-of-care

diagnosis for elderly individuals, whom require frequent

monitoring and telehealthcare at home.

5 Experimental method

Figure 4 illustrates the five steps of the experimental

procedure. First, 10 ll of nanoprobes suspension with a

measured conductivity of about 500 ls/cm (SC-170,

Suntex Instruments Co., Ltd.) was dripped on the sensing

area of an immunosensor by using a micropipette. The

AC signals were then applied to IDMA for trapping

nanoprobes on IDMA surface by a positive DEP force;

the AC signals were 10 Vpp at 50 kHz through a function

generator (AFG3022, Tektronix). Owing to that the DEP

trapping was performed in the atmospheric environment

at room temperature, AC signals were applied to IDMA

for only 10 min before the nanoprobes suspension dried

out. Following DEP trapping, the immunosensor was

washed by DI water and dried by a N2 gun in the second

step. Notably, although the DEP force was not applied

during the washing procedure, the immobilized

nanoprobes were not removed in the experimental

observation. This formation can be attributed to the layer-

like microstructure of PANDB nanoparticles and surface

charge of proteins. Therefore, based on van der Waals’

force and electrostatic force, physisorption between the

immobilized nanoprobes and electrode surface was

improved. Meanwhile, the initial impedance at 50 kHz of

immunosensor was measured by the proposed impedance

system module and denoted as Zprobe. To obtain a stable

value of Zprobe, the above procedures were iterated for six

times to investigate the variation of Zprobe that corre-

sponded to DEP trapping within 60 min. In the third step

of the experimental procedure, a chartula (i.e., a folded

paper for containing medicinal powder) with a size of

7 mm 9 25 mm was immersed into the T24 cell lysate

for 2 s, and then, the moistened paper was swept across

the sensing area. Chartula rather than conventional copy

paper was chosen owing to its smoother texture of char-

tula and greater density. Therefore, the absorbed protein

remains on the chartula surface rather than in the texture.

Additionally, the quantity of T24 cell lysate adsorbed on

the paper was measured about 0.01 g. In the fourth step

of the experimental procedure, T24 cell lysate which

contained various proteins was incubated for 10 min of

immunoreaction. However, only the galectin-1 protein in

T24 cell lysate conjugated with nanoprobes, owing to the

specific binding between the antibody and antigen. In the

fifth step of the experimental procedure, an attempt was

made to eliminate the non-binding proteins on the

immunosensor by washing the sensor surface with DI

water and then drying it by a N2 gun. The impedance of

IDMA at 50 kHz after immunoassay was then measured

by the proposed impedance system module and denoted

as ZT24. The variation of impedance due to specific

binding of galectin-1 protein can be attributed to

DZ = ZT24 - Zprobe. Moreover, an attempt was made to

eliminate the difference of initial impedance (Zprobe)

between different immunosensors by normalizing DZ with

Zprobe, i.e., DZ/Zprobe, which is denoted as a normalized

variation of impedance. In our experiments, three T24 cell

lysate concentrations were evaluated with respect to the

sensitivity of an immunosensor based on DZ/Zprobe.

6 Results and discussion

6.1 Characterization of nanoprobes

This work also demonstrated the feasibility of imple-

menting the final step of nanoprobes preparation, i.e.,

binding antibodies on PANDB/Al2O3 nanoparticles, by

comparing two FT-IR spectra of PANDB/Al2O3 nano-

particles before and after binding with antibodies in

Microfluid Nanofluid

123

Fig. 1c. According to this figure, three regions in the

spectra changed after binding with antibodies, implying

that these functional groups explicitly stretched or

changed after the modification. As a result, two regions

appearing around 1,780 and 3,000 cm-1 are, respec-

tively, assigned to C=O and O(=)C–H; another peak

appearing at 1,300 cm-1 reveals the presence of C–N

bond, especially for the NH2 group. The functional

group C=O may originate from the carboxylic acid

(–COOH) groups of galectin-1. Therefore, we can infer

that the free amino groups on PANDB surface (NH2)

bond with the carbohydrate groups (O(=)C–H) of oxi-

dized antibodies, as shown in Fig. 1a. Consequently, the

PANDB nanoprobes were successfully synthesized for

protein immobilization in the immunosensing system.

Additionally, the antigen binding region (Fab region, the

Y-shape arms) is only 10 % of the antibody structure,

explaining why most of the carbohydrate group bridges

the amino groups; the Fab region can be exposed as the

receptor of antigen as well. This modification scheme

thus can assure the receptor of antibody exposed on the

NP surface for immunosensing.

6.2 Characterization of impedance system module

Accuracy of the proposed impedance system module

was verified using a precision LCR meter (WK 6420,

Wayne Kerr Electronics). The impedance of an immu-

nosensor with IDMA before coating was measured by

the proposed impedance system module and precision

LCR meter at 50 kHz for 10 times, respectively. The

average results of impedance were 577 and 588 kXfrom impedance system module and precision LCR

meter, respectively. Notably, the error is only 1.89 % if

the value from the LCR meter is used as the reference,

allowing us to achieve a highly accurate and reliable

impedance measurement based on the proposed imped-

ance system module.

6.3 Time effects of DEP trapping and immunosensing

Using the DEP force, this work also attempted to immo-

bilize nanoprobes onto the IDMA surface in order to

achieve a lower baseline of Zprobe after DEP trapping, in

which the impedance change was measured for 60 min

with 6 iterations of 10 min DEP trapping. According to the

blue curve in Fig. 5, the impedance decreased slightly

within 20 min, yet it descended a large range when DEP

trapping was performed for 30 min. The impedance then

remained stable in a low level until 60 min. Therefore, the

PANDB-coated nanoprobes effectively reduce the initial

impedance of IDMA after immobilization on the electrode

surface by DEP trapping. Thus, 30-min DEP trapping is

Fig. 4 Experimental procedure

for DEP trapping of nanoprobes

and immunosensing by

impendence sensing

Fig. 5 Impedance variation during trapping nanoprobes and immu-

nosensing for 60 min, blue line represents the decrease of impedance

when nanoprobes are immobilized on the electrode surface by DEP

force, where the red line represents the increase in impedance when

proteins binding with nanoprobes (color figure online)

Microfluid Nanofluid

123

sufficient for fabricating an immunosensor with a low

Zprobe. The SEM pictures in Fig. 6 display the microelec-

trodes before and after DEP trapping of nanoprobes for

60 min, indicating that the DEP force can immobilize the

nanoprobes on the IDMA surface, even when the washing

process is applied after DEP trapping.

However, an attempt was also made to confirm the

sufficient time for immunosensing by recording the

impedance change during immunosensing for 60 min, as

shown in the red curve in Fig. 5. Notably, the immuno-

sensor was prepared with 30 min of DEP trapping, in

which each iteration of immunosensing was 10 min; the

impedance was also measured after washing and drying.

Therefore, T24 cell lysate was swept six times for 60 min

of immunosensing. As a result, the impedance increased

rapidly during the first 10 min after the T24 cell lysate was

coated on the IDMA surface; the impedance then increased

slightly from 20 to 60 min after the coating. Ten minutes of

incubation time after sweeping T24 cell lysate for immu-

nosensing is thus sufficient. The proposed impedance sys-

tem module demonstrates the ability to achieve rapid

detection by paper sampling without a pump or syringe.

6.4 Immunosensing for different concentrations of T24

cell lysate

Three concentrations of T24 cell lysate ranging from 0.06

to 1.25 mg/ml were prepared to evaluate the sensitivity of

an immunosensor with galectin-1/PANDB/Al2O3 nanop-

robes. According to Fig. 4, all of the measurements fol-

lowed the standard procedures. The initial impedance of

Zprobe can be obtained after 30 min of DEP trapping

and ZT24 after 10 min of immunosensing, respectively.

Then, Fig. 7 plots the normalized variation of impedance,

DZ/Zprobe, for different concentrations of cell lysates.

According to this figure, the normalized variations of

impedance increase linearly with the concentration of T24

cell lysate. Additionally, sensitivity is defined as the per-

centage of normalized variations of impedance per

concentration of T24 cell lysate with mg/ml unit; the

sensitivity of the present immunosensor is about 124.4 %

per (mg/ml). For the lowest concentration of cell lysate

(0.0626 mg/ml), the normalized impedance change is

approximately 5 %, i.e., roughly five times greater than the

error bar (*1 %) measured by the module-level system.

However, the error bar in the case of 0.25 mg/ml is about

2 %, explaining why the concentration was not lowered

further to achieve an adequate signal/noise ratio. Restated,

0.0626 mg/ml is the detection limit based on the proposed

impedance module system. Figure 7 shows the impedances

of the immunosensor, which were also simultaneously

measured by the precision LCR meter. Comparing the

results from the proposed impedance system module and

LCR meter reveals that the error is \2 %. Therefore, the

reliability of proposed impedance system module is satis-

factory. Additionally, owing to that a cell lysate contains

Fig. 6 a SEM pictures of the

microelectrodes before DEP

trapping; b after DEP trapping

of nanoprobes for 60 min and

following wash process,

nanoprobes have been

immobilized on the interdigital

microelectrodes array (IDMA)

surface by DEP force

Fig. 7 Normalized variations of impedance (DZ/Zprobe,

DZ = ZT24 - Zprobe) for different concentrations of T24 cell lysate

Microfluid Nanofluid

123

over tens of thousands proteins; the impedance-based

immunosensor can still achieve a high sensitivity for a

specific galectin-1 protein. Therefore, the proposed por-

table system module and disposable immunosensor are

highly promising for use in point-of-care diagnostics.

7 Conclusions

This work has developed a module-level impedance system

integrated with an impedance-based immunosensor in

order to detect a specific protein level from bladder cancer

cell lysate. By applying the DEP method to immobilized

conductive-polymer-coated nanoprobes, galectin-1/

PANDB/Al2O3 NPs, on the microelectrode surface, the

immunosensor can achieve a high sensitivity and rapid

response within 10 min. Additionally, the proposed por-

table impedance system module demonstrates its high

reliability and accuracy over that of a sophisticated

impedance analyzer. In particular, a bladder cancer cell

lysate is used to evaluate this miniaturized immunosensing

system rather than using commercial biomarker kits. Those

results demonstrate the feasibility of using the proposed

module-level immunosensing system in point-of-care

diagnostics.

Acknowledgments The authors would like to thank the National

Science Council of the Republic of China, Taiwan, for financially

supporting this research under Contract No. NSC 102-2218-E-218-

001. The Optoelectronics Research Center at Southern Taiwan Uni-

versity of Science and Technology is appreciated for use of its MEMS

fabrication facilities. Ted Knoy is appreciated for his editorial

assistance.

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