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How to embed three-dimensional flexible electrodes in microfluidic devices forcell culture applications†

Andrea Pavesi,ab Francesco Piraino,a Gianfranco B. Fiore,a Kevin M. Farino,a Matteo Morettic

and Marco Rasponi*a

Received 31st January 2011, Accepted 7th March 2011

DOI: 10.1039/c1lc20084d

This communication describes a simple, rapid and cost effective

method of embedding a conductive and flexible material within

microfluidic devices as a means to realize uniform electric fields within

cellular microenvironments. Fluidic channels and electrodes are

fabricated by traditional soft-lithography in conjunction with chem-

ical etching of PDMS. Devices can be deformable (thus allowing for

a combination of electro-mechanical stimulation), they are made

from inexpensive materials and easily assembled by hand; this method

is thus accessible to a wide range of laboratories and budgets.

Advances in microfluidics have been significantly enhanced by the

development of simple, low cost fabrication techniques. Perhaps

the most influential of these techniques is the use of poly

(dimethylsiloxane), PDMS, for molding microchannels from silicon

master molds,1 now routinely employed. However, next generation

microfluidic devices, which incorporate electrodes,2 internal struc-

tures3–5 and other functionalities,6 demand new fabrication methods.

Several methods of incorporating electrodes into microfluidic

devices have already been proposed. The most practical approach is

the metal deposition of patterns on a glass substrate, and their further

alignment to a PDMS-based microfluidic device.7 However, although

widely used, this solution suffers from two major drawbacks: (i) the

electrodesare essentiallyplanar, therefore auniformdistributionof the

electric field across the channel cross-section is difficult to achieve; (ii) it

requires expensive metal deposition facilities. In addition, this solution

is not suited when device flexibility is required; indeed, direct deposi-

tion methods have been reported,8–10 but metals poorly adhere to

PDMS due to its low surface energy, which causes failure during the

fabrication processes. Furthermore thin metallic layers suffer from

potential cracking when bent or stressed.9

The generation of uniform electric fields, which is desirable in

dielectrophoretic and electro-orientation applications,11,12 can be

aBioengineering Department, Politecnico di Milano, Piazza Leonardo daVinci 32, 20133 Milano, Italy. E-mail: marco.rasponi@polimi.it; Fax:+39 02 2399-3360; Tel: +39 02 2399-3377bDepartment of Mechanics, Politecnico di Torino, Corso Duca degliAbruzzi 24, 10129 Torino, ItalycGruppo Ospedaliero San Donato Foundation, Corso di Porta Vigentina 18,20122 Milano, Italy

† Electronic supplementary information (ESI) available. See DOI:10.1039/c1lc20084d

This journal is ª The Royal Society of Chemistry 2011

achieved through three-dimensional (3D) electrodes; however their

realization is still challenging. Vertical electrodes in the sidewall of

rectangular micro- or nano-fluidic channels were obtained by means

of techniques (multi-step optical lithography involving SU-8 with

metal deposition and electroplating12,13) which proved to be expen-

sive. On the contrary, cost effective solutions to inject conductive

materials in PDMS devices were recently proposed.14–16 Although in

some works a direct contact between electrodes and fluidic channels

was achieved,14,16 these methods result limited as: (i) the injection

process is manual and hard to control (hence, only simple electrode

configurations can be reliably achieved) and (ii) the final shape of the

electrodes cannot be predetermined (thus losing the electric field

uniformity feature).

In this work, we present a novel and simple method to reliably

obtain 3D flexible electrodes with vertical walls through injection, thus

overcoming previous limitations. The electrodesconsistofa mixture of

PDMS and carbon nanotubes (CNTs), and the devices are fabricated

by single-step soft lithography in conjunction with chemical etching of

PDMS.17 Devices with complex electrode patterns were fabricated and

tested in terms of electrical response. Possible cytotoxicity effects were

also addressed by culturing H9c2 cells for seven days. This technique is

a potential breakthrough in the context of electrical stimulation of

cells; moreover, the inherent flexibility of the electrodes allows for the

concomitant application of mechanical stimulations, which proved to

drive key cellular processes, such as apoptosis, matrix deposition, gene

and protein expression,18,19 and stem cell differentiation.20 Finally,

being easy to automate, the proposed method is well-suited for fast

turnaround, low-cost microfluidics research and development.

Device fabrication

The fluidic layout consisted of a simplified geometry, with a single cell

culture region (500 mm wide and 5 mm long) connected to inlet and

outlet channels (200mmwideand5mmlong).Moreover,16additional

channels (electrode channels) were designed so as to nearly intersect

thecell culturechamber,butkept separatedfromthatbyathinwall (30

mm thick). The widthof each electrode channel was set equal to200 mm

nearby the culture chamber, and linearly increased moving away from

that (up to a final width of 1 mm).

Microfluidic devices were realized in PDMS by means of standard

soft lithography techniques.1 Master molds were fabricated by

replicating the photomask layouts with a patterned 170 mm thick

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layer of SU-8 50 (MicroChem Corp). PDMS (Sylgard 184, Dow

Corning Corp) was prepared by mixing prepolymer base solution

and curing agent at a ratio of 10 : 1 (w/w). After degassing, PDMS

was poured directly onto the master molds and cured in an oven at 80�C for 3 hours. The PDMS was gently peeled from the mold, and the

electrode ports were punched with a 1 mm diameter puncher, while

an 800 mm puncher was used to obtain input and output fluidic ports.

After a 30 seconds air plasma treatment (Harrick Plasma Inc), each

device was permanently bonded to a histology glass slide.

Electrodes consisted of microfluidic blind-end channels filled with

a conductive nanocomposite material. In order to prepare the elec-

trically conductive mixture, the following steps were followed: (1)

multi-walled carbon nanotubes (MWCNT) were suspended in chlo-

roform in a ratio of 1 : 30 (w/v) by sustained ultrasonic mixing in an

evaporation flask; (2) PDMS pre-polymer was added to the nanotubes

suspension, in a ratio of 1 : 3 (w/w) with respect to the MWCNT

content, and mixed both mechanically and through further sonication;

(3) to achieve complete chloroform evaporation, the flask was placed

in a rotary evaporator for 40 minutes; (4) before use, the PDMS curing

agent was added to the pre-polymer/nanotube mixture, in a ratio of

1 : 10 (w/w) with respect to the prepolymer, and manually mixed.

The process of electrode embedding is schematically depicted in

Fig. 1. A 2.5 ml plastic syringe was back-filled with the conductive

nanocomposite, and used to manually inject the electrodes into the 16

channels (Fig. 1a), taking advantage of the PDMS permeability to

gases (Fig. 1b). In order to establish physical connection between the

electrodes and the cell culture chamber, the separation membrane was

removed by means of chemical etching following a previously

Fig. 1 A simple method to embed 3D electrodes in microfluidic devices

through injection. On the left, a schematic representation of a PDMS

device bonded to a glass substrate. The top of the channels is sketched as

open to better illustrate the steps of the process. Initially, the fluidic

channel (top to bottom) and the electrode channels (lateral blind-end

ones) are empty (a). A conductive material is manually injected into the

lateral channels, which can be completely filled by taking advantage of

the permeability of PDMS to gases (b). Subsequent to a curing step in

oven, a PDMS etchant solution, composed of tetrabutylammonium

fluoride (TBAF) and N-methylpyrrolidinone (NMP) in a ratio of 1 : 3 (w/

w), is pumped through the fluidic channel (c) until complete removal of

lateral walls is achieved (d). On the right, pictures of the cell culture device

are presented. The top view (f) shows in blue the fluidic path (inlet and

outlet channels connecting the culture chamber), and in black the 16

electrode channels. Cross-section and top view magnification of the

electrode/culture chamber region are also shown before (e) and after (g)

the etching step.

1594 | Lab Chip, 2011, 11, 1593–1595

published protocol.18 A PDMS etching solution composed of tetra-

butylammonium fluoride (TBAF) and N-methylpyrrolidinone

(NMP), in a ratio of 1 : 3 (w/w), was pumped through the cell culture

channel (Fig. 1c). After 20 minutes, the PDMS etchant completely

dissolved the 16 separationwalls, enlarging at the same time the culture

chamber to a final width of 560 mm (Fig. 1d). The channel was then

washed with NMP and ethanol.

Electrode characterization

Prototype devices were designed and fabricated in order to charac-

terize the electrical behavior of the electrodes. The devices, consisting

of independent and parallel straight channels (10 mm long, 200 mm

wide and 150 mm high), were molded out of PDMS. Input and output

ports were punched at each channel end, right before plasma bonding

on a glass substrate. Steel couplers were connected to the device ports

and used to inject the electrodes; successively they served to perform

electrical measurements on the nanocomposite electrodes. The

measured resistance value (4.87 � 1.46 kU) yielded an electrical

resistivity of the mixture of 14.6� 4.4� 10�3 U m.

Cell culture

In order to rule out potential cytotoxic effects related to the electrode

fabrication technique, myoblast cell line H9c2, derived from embry-

onic rat heart,20 was used as an in vitro cellular model for cardiac cells.

H9c2 cells were cultured in Dulbecco’s Modified Eagle Medium/F12

(Invitrogen) supplemented with 5% Fetal Bovine Serum and 1%

penicillin and streptomycin at 37 �C and 5% CO2 hereby called

completemedium(CM).Atsemi-confluence, cellsweredetachedusing

0.25% trypsin/1 mM EDTA (Gibco), centrifuged at 800 rpm for 5 min

and resuspended in CM at 2� 106 cells per ml for seeding. Cells were

manipulated under sterile tissue culture hoods and maintained in a 5%

CO2 humidified incubator at 37 �C.

Prior tocell seeding, devices were preconditionedwith ethanol 100%

(10 minutes) and phosphate buffer solution (PBS) and coated with

a solution of 1.1% rat tail collagen type I (BD Bioscience) at 0.02 N

acetic acid (injection of 15 ml of coating solution followed by an incu-

bation of 60 minutes). Afterwards, a complete rinse was performed

three times with PBS, right before filling the device with 100 ml of CM

and incubating overnight. A 15 ml volume of cellular suspension was

injected, through the inlet port, directly into the microfluidic channel.

Devices were incubated in a humidified incubator at 37 �C with 5%

CO2 during the following seven days, changing the medium daily.

After the first three days of culture, H9c2 cells were electrically

stimulated for the remaining 4 days using a biphasic square wave21,22

(1 ms at +3 V, 1 ms at �3 V at 1 Hz). The imposed electrical stim-

ulation was assessed through inlet and outlet pins by recording the

signal with a multifunction acquisition system (USB-6009, National

Instruments Corp).

Cell viability was evaluated using a LIVE/DEAD kit (Molecular

Probes). At day 7 of culture, the assay solution (1 mM calcein AM

and 2 mM EthD-1a diluted in 5 ml of sterile PBS solution) was

prepared and used to replace the CM contained in the PDMS device.

After 20 min of incubation, the samples were visualized under

a fluorescence microscope (IX70, Olympus Corp). Fig. 2 shows

a typical field of view within the culture chamber, evidencing

a marked presence of vital cells (fluorescing green) and a substantial

absence of dead cells (emitting red).

This journal is ª The Royal Society of Chemistry 2011

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Conclusions

This work presents a simple method to embed flexible electrodes in

PDMS-based microfluidic devices, which does not require further

microfabrication steps other than those related to fluidic channel

realization. Multi-walled carbon nanotubes, in a controlled mixture

with PDMS, are used as a conductive filler of specifically designed

blind-end channels; successively, direct contact between electrodes and

fluidic channels is achieved through PDMS etching. Such a technique

intrinsically enables the realization of precise 3D electrodes, even in

complex configurations, overcoming previous limitations.14–16 Indeed,

vertical walled 3D electrodes produce uniform electric fields within

microfluidic channels, which can be advantageous in many micro-

fluidic designs.

The devices used in this study had a geometry consisting of a single

culture channel and sixteen lateral electrodes. The lateral resolution of

the electrodes is only determined by the specific channel dimensions.

In this work, electrodes were vertical, facing the lateral walls of the

culture chamber, with dimensions of 200 � 170 mm. However, the

same nanocomposite material was successfully tested for injection in

channels as small as 50 � 100 mm. Despite its complex design, the

device resulted in an easy-to-use platform for biological operations,

demonstrated by the fact that H9c2 cells were successfully condi-

tioned for seven days, without increasing the complexity of routine

cell culture operations.

The nanocomposite material yielded relatively high resistivity

values, as compared to standard electrode materials, which could

represent a limitation in those applications involving sensing or

detections. Indeed, the increase of the MWCNT content also

increased the electrical conductivity, but limited the overall

Fig. 2 (a) A device consisting of inlet and outlet channels, connecting a 5

mm long, 500 mm wide and 170 mm high cell culture chamber (blue color),

was developed. Each lateral wall of the chamber was provided with 170�200 mm vertical electrodes, spaced 400 mm one to the other (black color).

Potential cytotoxic effects related to the electrode fabrication technique

were evaluated through seven days myoblast cell line H9c2 cultures. After

the first three days of culture, H9c2 cells were electrically stimulated for

the remaining 4 days using a biphasic square wave (1 ms at +3 V, 1 ms at

�3 V at 1 Hz). At day 7, the cell viability was evaluated using a LIVE/

DEAD kit (Molecular Probes). Scale bar: 5 mm. (b) A typical field of

view within the culture chamber is shown, evidencing a marked presence

of vital cells (fluorescing green) and a substantial absence of dead cells

(emitting red). Scale bar: 500 mm.

This journal is ª The Royal Society of Chemistry 2011

injectability. Although the mixture injectability in open micro-

channels was still possible (at CNT contents above 25% as w/w), the

pressure required for the injection in closed channels did not allow to

reliably obtain electrodes. For this purpose, whenever electrode

flexibility and biocompatibility are not a strict requirement, the

injected material can be easily substituted (e.g. with low melting-point

metals). The same technology can be also used to fabricate fully

deformable devices, by simply adopting a PDMS substrate. Hence,

the proposed technique allows, to standard microfluidic laboratories,

for manufacturing devices suited for biological studies involving

combinations of electrical and mechanical stimulations.

Acknowledgements

We acknowledge a Cariplo Foundation grant (# 2008-2531) and the

Rocca Foundation for financial support.

Notes and references

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