Sensors and Actuators Reports 2 (2020) 100004

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Sensors and Actuators Reports

journal homepage: www.elsevier.com/locate/snr

Monolayer graphene chemiresistive biosensor for rapid bacteria detection

in a microchannel

Weiqi Zhao, Yunyi Xing, Yang Lin, Yuan Gao, Mengren Wu, Jie Xu

∗

Department of Mechanical and Industrial Engineering, University of Illinois at Chicago, Chicago, IL 60607, United States

a r t i c l e i n f o

Keywords:

Bacteria

Escherichia coli (E. coli)

Monolayer graphene

Chemiresistive biosensor

Concentration

Microchannel

a b s t r a c t

Bacteria, such as Escherichia coli (E. coli), can cause food poisoning and serious diseases. In fact, E. coli has

emerged multiple times as a severe public health concern in recent years. Therefore, it is important to detect

bacteria like E. coli in simple, rapid and sensitive approaches. In this study, a gate-free chemiresistive biosensor

based on monolayer graphene (MG) is proposed to detect E. coli , thanks to the high sensitivity stemming from

the anti-E. coli antibody-coated graphene. The immobilization of the antibodies is performed via streptavidin

and biotin conjugation on the graphene surface. Compared to conventional bacterial biosensors, the proposed

chemiresistive biosensor captures the E. coli bacteria on the surface of the sensor and performs the detection

through electric readouts, other than optical signals. That is the resistance measured in the biosensor increases

along with the increase of concentration of the bacteria. To further extend the capabilities of this biosensor and

provide a controllable test flow, a microchannel is integrated on the graphene surface. The results show that

the developed chemiresistive biosensor is able to detect E. coli in trace concentration down to 12 cfu mL − 1 ,

indicating an excellent sensing performance. The proposed monolayer graphene based chemiresistive biosensor

clearly manifest its advantages of low-cost, ease of fabrication, portability and good sensitivity, as well as the

potential for rapid in-situ bacterial detection.

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. Introduction

Bacteria accompany our lives, but sometimes they cause problems

n our health. For example, Escherichia coli (E. coli), which usually col-

nizes human intestines right after birth [1] , inevitably affects human

eings throughout their entire life spans. The discovery of E. coli can be

ated back to 1885 by Theodor Escherich, and it was originally consid-

red as a healthy mutualism bacterium helping human beings with syn-

hesis of vitamin K2 [2] . In 1996, scientists found that E. coli can cause

ood poisoning in human beings, as evidenced by a worldwide outbreak

f food poisoning occurred in Scotland [3] . In addition, E. coli infec-

ions can cause diarrhea and lead to life-threatening conditions, such

s kidney failure, high blood pressure, chronic kidney disease, and neu-

ologic problems. To minimize the impact of E. coli on human society,

fficient diagnosis of the bacteria is of utmost importance. At present,

variety of methods have been developed to detect and determine the

oncentration of E. coli . For instance, one of the most straightforward

ethods is through cell culture, in which E. coli is cultured on a petri

ish for 24 h at 37 °C in an incubator. By means of counting the number

f E. coli colonies, the concentration of E. coli can be determined [4] .

onetheless, this method usually requires long preparation time. There-

ore it has been gradually substituted by other approaches. For exam-

∗ Corresponding author.

E-mail address: jiexu@uic.edu (J. Xu).

ttps://doi.org/10.1016/j.snr.2020.100004

eceived 21 December 2019; Received in revised form 4 February 2020; Accepted 12

vailable online 18 February 2020

666-0539/© 2020 The Authors. Published by Elsevier B.V. This is an open access ar

http://creativecommons.org/licenses/by-nc-nd/4.0/ )

le, nucleic acid-based polymerase chain reaction (PCR) and enzyme-

inked immunosorbent assay (ELISA) technology could give rise to a

uch more accurate result in a shorter amount of time [5–7] . Neverthe-

ess, they are still not satisfactory for rapid detection as they are costly,

ime consuming and require immobile laboratory settings.

Recently, various electrochemical based biosensors for the detection

f E. coli pathogen has been developed [8] . For example, anodic particle

oulometry technique has been combined with cyclic voltammetry (CV),

nd the absorption spectra mean decreases as the concentration of the E.

oli decreases [9] . Moreover, nanomaterials such as gold nanorods and

old nanoparticles (AuNPs) have also been incorporated into sensing as-

ays to further enhance the specificity and sensitivity of detection [10] .

or example, Bhalla and coworkers have shown that gold nanoparticles

an increase the sensitivity of biosensors significantly. Specifically, the

old nanoparticles helped the detection of cardiac troponin I (cTnI) at a

oncentration as low as 0.2 ng mL − 1 , which is much better than ELISA

ests of 4.3 ng mL − 1 [11] . Depending on the function of gold nanoparti-

les, they can be linked to aptamer for capturing target bacteria. In order

o find the adaptor on the E. coli membrane for aptamer, the amplified

ingle-strand DNA need to be translated based on the DNA array of E.

oli , in which the concentration of E. coli can be obtained [12] . How-

ver, the DNA screening of aptamer is relatively difficult, and the speci-

February 2020

ticle under the CC BY-NC-ND license.

W. Zhao, Y. Xing and Y. Lin et al. Sensors and Actuators Reports 2 (2020) 100004

Fig. 1. (a) The structure of the PDMS microchannel. The PDMS microchannel was fixed on the monolayer-graphene with electrodes. Under the PDMS microchannel

is the structure of the biosensor in the center of the microchannel. (b) The sputter coater was used to coat two sides of electrodes. The mask was covered on the

upper surface of the graphene.

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Fig. 2. A photo of the proposed microfluidic chemiresistive biosensor.

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city provided by aptamers is often limited. In this study, a chemiresis-

ive biosensor based on monolayer graphene is developed with electrical

eadouts for E. coli detection. Instead of using aptamers, we choose to use

ntibody-streptavidin system together with AuNPs for high specificity

nd high sensitivity bacteria recognition. Streptavidin has been proved

o be effective when it is linked to gold nanoparticles for biosensing

13] . The sensing element in our system is monolayer graphene (MG).

raphene is widely used in biosensors because of its special electri-

al properties [14] . Graphene can transfer 15,000 cm

2 V s − 1 electrons

t room temperature while the resistance is only around 10 − 8 Ωm

− 1 .

oreover, multiple studies of monolayer-graphene have confirmed their

normous stiffness and ultrathin thickness [15–17] . Thereby monolayer

raphene with large contact area is naturally suitable for providing

bundant binding sites for sensing purpose. Graphene’s electric prop-

rties, such as conductivity, is very sensitive to its environment, not to

ention chemical bindings on its surface. Hence, we immobilize inte-

rated AuNPs and streptavidin-antibody system on the graphene surface

or E. coli capturing and monitor the electrical signals generated. Gold

anoparticles were spread over monolayer graphene to produce conduc-

ivity, meanwhile, the electrode was produced from gold nanoparticle

hich can increase the sensitivity of biosensors [18] . As observed in

he experiments, when the system binds with E. coli, the resistance of

he monolayer graphene is increased. Therefore, the bio-signals of E.

oli binding are converted into electrical signals, and the concentration

f E. coli is then verified. It is worth mentioning that both graphene

nd gold are antibacterial, which makes the device safer to operate for

athogen-related applications [19] .

With the development of microfluidic technology, different types of

ab-on-a-chip devices with in-channel chemical sensors have been de-

eloped. For instance, the microfluidic system based on single-walled

arbon nanotubes [20] and based on enzyme assays [21] have been inte-

rated for detecting glycerol concentration. In addition, a microchannel

ystem can also be used in glucose detection [22] . Research activities

nd advances in lab-on-a-chip devices have revolutionized the field of

iosensing research, since sensing in a microfluidic environment often

enefits from stable and precise sample control, parallel and multiplexed

ssay capabilities, and scaling-up potential. By nature, electrochemical

ensors are easily and effectively to integrate with microfluidic plat-

orms because of their simple structure. In this paper, the developed

iosensor is integrated into a lab-on-a-chip device for the detection of

ifferent concentration of E. coli solution.

. Materials and methods

.1. Materials

Monolayer graphene coated on a square substrate of Si/SiO 2 was

urchased from Graphenea, Inc. Gold/palladium nano-particle target

as purchased from Ted Pella, Inc. Phosphate buffered saline (PBS,

.1 mol L − 1 , pH 7.4) with 0.1% Tween 20 was purchased from Sigma

ldrich Inc. Streptavidin was purchased from AnaSpec, Inc. E. coli K12

R2925 were purchased from New England Biolab, Inc. Biotinylated

nti - E. coli (from rabbit) antibody was purchased from abcam, Inc. Goat

nti-rabbit IgG (Heavy & Light Chain) antibody (Atto 488) was pur-

hased from Antibody-Online, Inc. The tryptic soy broth and tryptic soy

gar for medium were purchased from Thermo Fisher Scientific, Inc.

.2. Device fabrication

The overall detection system is illustrated in Fig. 1 (a). A monolayer-

raphene substrate with a size of 1 cm × 1 cm was first scrutinized to

nsure good quality. Afterwards, two electrodes were coated on the sur-

ace of the graphene substrate with gold/palladium alloy target (99.99%

u:Pd, 60:40 ratio) supported by a low pressure argon gas chamber

0.05 mbar) using a Au/Pd Polaron sputter coater E5100 series II. As

hown in Fig. 1 (b), a predesigned mask was fixed on the upper surface

f the substrate, after which the sputter coater was used to deposit two

5 nm thick Au/Pd continuous films as electrodes (for 10 min). After

tripping the mask, the sputtering coater was used for a second time,

W. Zhao, Y. Xing and Y. Lin et al. Sensors and Actuators Reports 2 (2020) 100004

Fig. 3. (a) A small area of the MG-based chemiresistive sensor in between the Au/Pd electrodes was characterized by SEM. (b) Raman spectra of the MG chemiresistive

sensor.

Fig. 4. Recording of the resistance change of anti-E. coli an-

tibody functionalized MG-based chemiresistive sensors when

interacting with E. coli solution (incubated with 10 7 cfu mL − 1

E. coli for 50 min).

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ut the second time coating was on the entire surface of the substrate

nd was only for 10 s. This way, the center of the graphene substrate

ad been coated with nanoparticle stratum granulosum as Au particle

inker.

After preparation of the electrodes, two PDMS layers were added

nto the substrate to create a microchannel. The first layer was the mi-

rochannel layer of PDMS which was used to cover the top of graphene

iece. The width, length, and the depth of the microchannel are 800 𝜇m,

mm and 1 mm, respectively. The dimensions are chosen because they

an be made conveniently using low-cost technology. Specifically, the

DMS layers were created via standard soft lithography [ 23 , 24 ]. The

old of the microchannel was fabricated on a polymethyl methacrylate

PMMA) plate by a micro-milling machine (CNC Mini-Mill/3 PRO). The

DMS pre-mixture was then mixed, degassed, and poured onto the mold,

nd kept in an oven for 8 min at 75 °C. While the PDMS was still a little

it sticky, it was adhered onto the graphene, and punched for the injec-

ion and extraction holes on the two sides of the channel. After that, the

raphene was placed with the PDMS layer down on the plate. Then the

econd layer PDMS liquid was poured onto the substrate and sealed the

raphene substrate into the PDMS. The whole device was placed in the

ven for 20 min at 75 °C. The two grooves were cut on the two sides

f the electrodes, which were used to gain electric contact on the elec-

rodes for electric measurements ( Fig. 2 ). Finally, the microchannel was

insed with deionized water (DI water), in order to remove impurities

n the microchannel. Afterwards, the microchannel was injected and in-

ubated overnight with 15 μL streptavidin (1 mg mL –1 ) at 4 °C. After

insing with buffer (PBS containing 0.005% Tween 20) and DI water

horoughly, the graphene was treated by 15 𝜇L of 0.5 mg mL − 1 biotiny-

ated anti-E. coli antibodies for 2 h, followed by rinsing with DI water.

fter the fabrication of devices, they can be used immediately for the

etection of bacteria. In our experiment, a series of E. coli K12 solu-

ions with different concentrations were injected into the microchannel

or 5 min consecutively for E. coli detection and measurements. If the

evices are not to be used immediately, it is recommended to seal an

ntibody stabilizer PBS buffer in the microchannels for preserving the

roteins and antibodies for a relatively long time, and the devices should

e kept in a refrigerator (2–8 °C).

.3. Bacteria preparation

The original E. coli K12 was cultured in tryptic soy broth solution

nd incubated in an incubator for 16 h at 37 °C with 246 RPM stir

W. Zhao, Y. Xing and Y. Lin et al. Sensors and Actuators Reports 2 (2020) 100004

Fig. 5. (a) Resistance spectra of anti-E. coli antibody func-

tionalized MG-based chemiresistive sensor injected with dif-

ferent concentrations of E. coli solution (every concentration

solution was injected and sit in the microchannel for 5 min).

(b) Resistance spectra of anti-E. coli antibody functionalized

MG-based chemiresistive sensor injected with a constant flow

(2 𝜇L min − 1 ) of different concentrations of E. coli K12 solutions

(every concentration solution was injected and maintained at

the constant flowrate for 5 min).

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peed. The bacterial population was maximized, and then the initial

oncentrations of bacteria samples were obtained by serial dilution. The

ryptic soy agar was used in a petri dish to culture E. coli , in order to

erify the initial count of the E. coli , using the so-called plate count

ethod.

.4. Electrical measurement

Electrical measurements were performed on monolayer-graphene

iosensor using a micropositioner with a Keithley 4200 source me-

er system at room temperature. The sensing signal of the device was

ecorded by LabView software monitoring the change between two elec-

rodes in the current (I) for a given source voltage (U) from 0 V to 2 V

hen the device was exposed to different concentrations of E. coli solu-

ion. The scanning electron microscopy (SEM) images were obtained

sing a JSM 6320F SEM. Raman spectra was obtained on a micro-

aman spectrometer with 532 nm laser excitation. Fig. 2 shows the

ctual graphene biosensor integrated in a PDMS microchannel. Two

rooves on the two sides of the electrodes were used to get access to the

lectrodes.

. Results and discussion

.1. Characterization of monolayer graphene

The SEM image of the MG is depicted in Fig. 3 (a), which shows

n area in between the two electrodes of the sensor. As observed, the

raphene film is continuous, uniform, and dominantly single-layered

tructure across the entire surface. Raman spectra measurements were

erformed to further confirm the presence of the high-quality MG layer.

s shown in Fig. 3 (b), the MG displays at two prominent peaks at

580 cm

− 1 and 2676 cm

− 1 , corresponding to the well-documented G

nd 2D peaks, respectively [ 25 , 26 ]. The fact that the 2D peak is greater

han the G peak indicates the presence of a defect-free MG layer on the

hemiresistor sensor, which is also in accordance with the evidence from

he SEM image ( Fig. 3 (a)).

.2. Characterization of linker

The linkers were applied on the graphene substrate to capture the

. coli , which then causes a change in resistance in MG for detection.

he primary linkers are Au nanoparticle, streptavidin and antibody on

W. Zhao, Y. Xing and Y. Lin et al. Sensors and Actuators Reports 2 (2020) 100004

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he graphene. Streptavidin is the most widely used analogue of avidin

27] . The avidin-biotin affinity interaction was approximately 10 3 to

0 6 times higher than an antibody-antigen interaction [28] . Therefore,

pplication of streptavidin in this biosensor enables very stable linking

compared to antibody-antigen). Moreover, streptavidin can protect the

iotinyl esters from hydrolysis, while avidin augments this hydrolysis

29] . Au nanoparticles and streptavidin are biologically specific bind-

ng. These functional groups of water-soluble nanoparticles are usually

arboxylic acids, which stabilize nanoparticles by electrostatic repulsion

nd can be used to conjugate other molecules to particles [30] .

In order to verify that these linkers (Au nanoparticle, streptavidin

nd anti-E. coli antibody) can be linked together very well, we used la-

elled secondary antibody to verify that the anti-E. coli had fixed on the

urface. The Au nanoparticles were coated on a glass slide. Meanwhile

circle was drawn on the glass slide, and the streptavidin and anti-E.

oli solution were dropped on this area respectively. After that, PBS so-

ution was applied to rinse the slide surface three times. The secondary

ntibody with fluorescence under a laser microscope was observed in

he circled area only. Therefore, it was confirmed that these linkers had

een jointed together and the other linkers which were not linked had

een rinsed out.

.3. Sensing results and discussion

In the bacteria-antibody binding system, the amount of the bacteria-

ntibody bonds being produced during the sensing process directly af-

ects the resistance of the MG. It is reported that the membrane of natu-

al biological cells shows a resistance of 10 2 –10 5 Ω cm

− 1 [ 18 , 31 ]. When

he bacteria cells attach to the surface of the graphene film, the forma-

ion of bacteria-antibody conjugation could produce kinetics barrier for

he electron transfer process, which may induce more distinct changes

n the carrier hole density in graphene film through polarization of cell-

urface charges, intracellular bioactivity. In addition, more and more

lectron clouds were caused by steric hindrance, which restricts the pas-

age of electrons to the electrode surface, resulting in the resistance in-

rease of the chemiresistive sensors. Therefore, the resistance change of

he system can be directly used to indicate the concentration of E. coli

n the sample solution. In order to determine the kinetics of bacteria-

ntibody binding event, the MG-based chemiresistive sensor was made

n contact with 10 7 cfu mL − 1 of E. coli bacteria solution. As shown in

ig. 4 , the graphene resistance was measured in every 5 min from 0

o 50 min, which shows a trend of resistance increase in time. This is

ttributed to the gradual increase in the number of E. coli captured by

he antibodies on the graphene surface. The resistance values increased

harply and reached a plateau as the incubation time increases. These

esults suggest that the binding reaction finished in about 10–15 min

ue to the saturation of bacteria-antibody binding. In order to avoid

he exhaustion of antibodies for sensing purpose, we decided to choose

min for the time lag in all the following sensing experiments.

The resistance spectra was obtained in Fig. 5 (a) for the MG-based

hemiresistive sensor when bacteria solution was injected into the mi-

rochannel and let sit for 5 min. The tested concentrations of the E. coli

12 solutions ranged from as low as 2.4 × 5 to 2.4 × 5 7 cfu mL − 1 . In

ig. 5 , ∆R is defined as the absolute change of resistance

R = 𝑅 𝑥 − 𝑅 0 (1)

here R 0 is the initial resistance of the non-bacteria chemiresistive sen-

or and R x is the actual resistance value of the chemiresistor after 5 min

f injection and sitting of E. coli K12 samples with various concentra-

ions of 2.4 × 5 X cfu mL − 1 ( x = 0–8). The result shows that the resistance

f the MG-based chemiresistive sensors tend to increase with logarith-

ically increasing adsorption of E. coli on the sensor surface. The ( ∆R)

as finally plateaued at about 260 ohms, which means that the anti-

ody was probably completely consumed by E. coli in 5 min at high

oncentrations. These results confirm that the combination of MG-based

hemiresistive sensor with the real-time measurement of the resistance

hange is appropriate for the development of highly sensitive pathogen

acteria detection sensors for a wide range of concentration.

We also explored another sampling mode, where the sample flow

ate was maintained at 2 𝜇L min − 1 after a series of E. coli K12 solutions

ith different concentrations were injected in the microchannel. The

esults are shown in Fig. 5 (b), and the increasing trend is very similar to

he injection-and-stop sampling method. This means that the MG-based

hemiresistive biosensor is suitable for being integrated into continuous-

ow microfluidic systems for more sophisticated sensing and diagnostics

asks.

. Conclusions

In this work, we developed a novel microfluidic approach to detect

. coli K12 using a monolayer graphene chemiresistive biosensor. The

icrochannel was fabricated using PDMS, and Au nanoparticles and

treptavidin combination on the graphene were applied to conjugate

he antibody. The solution samples were injected into the microchannel

or testing, which was achieved by monitoring the graphene resistance

hanges during the adsorption of E. coli by the antibody. A linear in-

reasing trend of ∆R was observed with logarithmically increasing E.

oli concentrations from 2.4 × 5 to 2.4 × 5 6 cfu mL − 1 . Both injection-

nd-stop and constant-injection modes of the samples were tested and

he results indicate a similar trend. Overall, the developed microfluidic

ensor demonstrates rapid and sensitive bacterial detection. One major

imitation here is that the graphene sensors cannot be recycled for reuse.

herefore, these sensors will have to be disposable in real-life applica-

ions. But the portability, low-cost, and ease of fabrication of the sensor

rovides a promising approach for future development of multiplexed

icrofluidic biosensor platforms for a wide variety of applications.

eclaration of Competing Interest

The authors declare no conflict of interest.

cknowledgments

This work was supported by an Early Career Faculty grant

80NSSC17K0522) from NASA’s Space Technology Research Grants Pro-

ram.

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