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Droplet-based interfacial capacitive sensing†

Baoqing Nie,a Siyuan Xing,a James D. Brandtb and Tingrui Pan*a

Received 28th November 2011, Accepted 19th December 2011

DOI: 10.1039/c2lc21168h

This paper presented a novel droplet-based pressure sensor using elastic and capacitive electrode–

electrolyte interfaces to achieve ultrahigh mechanical-to-electrical sensitivity (1.58 mF kPa�1) and

resolution (1.8 Pa) with a simple device architecture. The miniature transparent droplet sensors,

fabricated by one-step laser micromachining, consisted of two flexible polymer membranes with

conductive coating and a separation layer hosting a sensing chamber for an electrolyte droplet. The

sensing principle primarily relied on high elasticity of the sensing droplet and large capacitance

presented at the electrode–electrolyte interface. A simple surface modification scheme was introduced

to the conductive coating, which reduced hysteresis of the droplet deformation without substantially

compromising the interfacial capacitance. Moreover, the major concern of liquid evaporation was

addressed by a mixture of glycerol and electrolyte with long-term stability in a laboratory environment.

Theoretical analyses and experimental investigations on several design parameters (i.e., the dimensions

of the sensing chamber and the droplet size) were thoroughly conducted to characterize and optimize

the overall sensitivity of the device. Moreover, the environmental influences (e.g., temperature and

humidity) on the capacitive measurement were further investigated. Finally, the simply constructed and

mechanically flexible droplet sensor was successfully applied to detect minute blood pressure variations

on the skin surface (with the maximum value less than 100 Pa) throughout cardiovascular cycles.

Introduction

Microfluidic-based sensors, enabled by emerging soft lithog-

raphy techniques,1,2 have been an active area of research, for

their excellent flexibility, high sensitivity, simple fabrication, and

wide adaptability.3–6 A variety of sensing and actuation mecha-

nisms have been incorporated in the development of microfluidic

sensing devices, most of which rely on changes in physical

properties (e.g., optical, electrical or mechanical) induced by

fluidic displacement7 and/or new material functionality intro-

duced to working fluids (e.g., as optical and electromagnetic

waveguides).8,9 For instance, an optical microfluidic pressure

sensor was devised to detect the cross-sectional diameter at

a focal plane, of which the elastic sensing chambers, loaded with

a suspension of fluorescently labeled nanoparticles, experienced

volumetric change under mechanical deformation and yielded

donut-shaped focal areas in different diameters under a fluores-

cence microscope. The mechanical-to-optical readout of the

aMicro-Nano Innovations (MiNI) Laboratory, Department of BiomedicalEngineering, University of California, Davis, USA. E-mail: tingrui@ucdavis.edubDepartment of Ophthalmology, University of California, Davis HealthSystem, USA

† Electronic supplementary information (ESI) available: Mathematicalderivation for eqn (1) along with the analytical expressions for the twoconstants related to mechanical sensitivities. Measurement method ofthe response time in Table 1. See DOI: 10.1039/c2lc21168h

1110 | Lab Chip, 2012, 12, 1110–1118

microfluidic system provided pressure sensing ranges up to 10 psi

with 80% of the diameter changes of the deflectable membrane.7

In addition, microbubbles were introduced as a regenerable

pressure sensing element. Built in an aqueous-filled parylene

chamber, a pair of platinum electrodes were utilized to produce

micrometre-size gas bubbles through electrolysis, and mean-

while, to assess the pressure-dependent bubble volume through

measurement of electrical impedance of the fluid in the confined

volume. The bubble-based microsensor offered a simple elec-

tromechanical sensing scheme (�11.8 U per psi) with on demand

bubble generation for a small pressure range (�2 to 4 psi).

However, it was highly sensitive to environmental influences

(e.g., change of position and temperature).10 Another group

reported an alternative scheme for inertial sensing using inter-

facial instability of aqueous droplets on superhydrophobic

surfaces, where an array of parallel electrodes detected the

positioning of the droplets in one dimension.11 More recently,

emerging conductive fluidic materials, such as liquid metals and

ionic liquids,8,12–14 were directly incorporated as sensing elements

for their mechanical flexibility and reversible deformability. In

one implementation, a mercury droplet was sandwiched between

two planar electrodes covered by an ultrahigh-permittivity

material. It offered high electrical sensitivity (2.24 mF kPa�1), and

more importantly, was compatible with the standard CMOS

processing, which enabled system integration and further

reduced overall device dimensions.14 Another example was to

include highly conductive ionic liquids as strain gauges in

This journal is ª The Royal Society of Chemistry 2012

Fig. 1 Droplet-based interfacial capacitive sensors (dyed with different

colors).

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microfluidic channels and measure the corresponding change of

electric resistance in a standard Wheatstone bridge configura-

tion, which led to a semi-transparent and flexible sensing unit.

However, this design encountered a few critical sensing issues,

such as slow mechanical responses (response time of 4 seconds

due to the high fluidic viscosity), strong dependence on

thermal fluctuation, and more severely, low pressure sensitivity

(1.83 mV kPa�1).12

Capacitive sensing has gained increasing popularity in circuit

design for its high electrical sensitivity, low power consumption,

compact layout, simple device construction and immunity to

temperature fluctuation and thermal noises,14–17 in comparison

with the resistive counterparts. Recently, capacitive sensors have

been introduced to an array of biomedical applications,

including intraocular pressure monitoring,18 tactile sensing,19

and bio-analytical detection.20 For instance, capacitive pressure

sensors in a wireless configuration were introduced to monitor

intraocular pressure. The capacitive sensing module with

a sensitivity of 3.47 fF kPa�1, using a conventional silicon-based

microfabrication integrated with processing circuitry, was fully

encapsulated in a transparent glass diaphragm, which allowed

continuous recording of intraocular pressure once implanted.18

Another set of capacitive sensors were employed to detect

specific biomolecule binding events on the ligand-modified elec-

trodes.20–22 In a recent report, various concentrations of human

serum albumin (HSA) (from 10�16 M to 10�8 M) can be detected,

according to the interfacial charges presented on the anti-HSA-

coated gold electrode surfaces. By applying pulsed currents with

controlled interval and amplitude to the capacitive sensor, the

signal decays can be repetitively evoked, from which the solu-

tion–electrode interfacial capacitance can be measured in real

time.20 Similarly, an aptamer-based capacitive sensing system

was devised on an aluminium electrode surface, on which specific

DNA aptamers binding to bisphenol A (BPA) were deposited.

Applying the sensor to a BPA solution, the change of BPA

concentration as low as 100 pM can be detected in a perfusion

test with 2.9% capacitive change.22 Moreover, a tactile sensing

scheme was implemented on the capacitive platform with

improved sensitivity and sensing range. Specifically, dielectric

polydimethylsiloxane (PDMS) layers were structured with

micro- or nanopatterns as sensing elements between two indium-

tin-oxide (ITO)-coated plastic surfaces. The authors claimed the

highest sensitivity of 0.55 kPa�1 (DC/C0/DP) was achieved over

a large surface area among existing capacitive sensing designs, in

response to external mechanical loads.19

In this paper, we first report a novel droplet-based sensing

mechanism utilizing a highly capacitive electric double layer

(EDL) presented at an extremely flexible droplet–electrode

interface. In particular, implemented on a simply suspended

membrane structure, the EDL offers high charge density at the

nanoscopic ionic–electronic interface to establish an ultrahigh

interfacial capacitance, while the elastic deformability of elec-

trolyte droplets on hydrophobic-modified electrodes allows the

reversible fluidic expansion/contraction, in response to external

mechanical stimuli. The interfacial droplet sensor achieves an

ultrahigh pressure sensitivity of 1.58 mF kPa�1 along with an

ultrahigh resolution of 1.8 Pa, comparable to the highest value

reported in the literature.14,19 In addition, the response time of

the sensing device (of 260 ms) has been characterized under

This journal is ª The Royal Society of Chemistry 2012

a constant membrane deformation (of 50 mm). It has also been

shown that with change of the medium (less viscous) or modified

surface energy (more hydrophobic), the mechanical response can

be improved significantly. The droplet sensing device, fabricated

by one-step laser micromachining, is comprised of two flexible

polymer membranes with conductive coating and a separation

layer with a sensing chamber hosting the electrolyte droplet, all

optically transparent and mechanically flexible. Moreover, add-

ing glycerol to a highly conductive electrolyte droplet addresses

the primary evaporative concern with a long-term stability for

such a liquid-based sensor under room conditions (46% humidity

and 24 �C). Theoretical analyses and experimental investigations

on key design parameters (i.e., the radius and height of the

sensing chamber and droplet size) have been thoroughly con-

ducted to characterize and optimize the overall device perfor-

mance. Furthermore, the performance of the droplet sensors

under different temperatures and humidity levels has been

investigated once the thermodynamic equilibrium has been

reached. To demonstrate the utility of the simply constructed and

mechanically flexible droplet sensor with ultrahigh sensitivity, it

has been successfully applied to detect minute blood pressure

variations on the skin surface (with the maximum value less than

100 Pa) throughout cardiovascular cycles. Fig. 1 illustrates the

droplet-based interfacial capacitive sensors, in which the sensing

droplets are dyed with various colors for improved visibility.

Operating principle

The electric double layer has been long known to present large

capacitance per unit area at the nanoscopic interface between an

electrode and electrolyte (as high as tens of mF cm�2),23,24 and has

been widely explored in the fields of energy storage and digital

microfluidics.25,26 Several theoretic models have been proposed to

explain the remarkable interfacial capacitance, established by

mobile electrons in a conductive solid phase and counter-ions

immigrating in the adjacent liquid environment.27–29 The inter-

facial capacitance is influenced by the surface charge density and

Debye length, which is used to describe the thickness of the EDL

when electrostatic and thermodynamic activities reach equilib-

rium in the solution phase. Surface charge density is influenced

Lab Chip, 2012, 12, 1110–1118 | 1111

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by the physiochemical property of the interface, kinetic energy of

the ionic species, electric potential as well as permittivity and

concentration of the solution.30 In this paper, the proposed

droplet-based interfacial sensing employs the ultra-large inter-

facial EDL capacitance at the highly elastic droplet–electrode

contact. The interfacial EDL capacitance of the sensor is

proportional to the area of the contact interface, unlike the solid-

state strain gauges, which measure the change of bulk resistance

under mechanical deformation.

Fig. 2 illustrates the principle of the droplet-based interfacial

capacitive sensing. As can be seen, a sensing chamber within the

elastic separation layer hosts an electrolyte droplet, which is

sandwiched between two polymeric membranes coated with

a transparent conductive material (Fig. 2a). The EDL forms

immediately upon the droplet–electrode contact, with mobile

electrons migrated from the conductive membrane surface and

a counter-ion layer accumulated from the electrolyte solution.

Under external mechanical loads, the suspended polymer

membranes and the separation layer deform elastically, and as

a result, the contact area of the droplet–electrode interface

experiences circumferential expansion (assuming incompressible

fluid with unaltered volume of the droplet). Given a relatively

constant charge density, the variation in the contact area will

lead to a proportional change in the interfacial capacitance. An

equivalent circuit diagram of the sensing structure has been

presented in Fig. 2b, in which the EDL capacitances at the

droplet–electrode interfaces are modeled as two capacitors con-

nected through a resistive element of the conductive droplet. The

overall deformation of the sensing chamber, which includes

the deflection of the flexible membrane and the compression of

the elastic separation layer, will lead to the change of the inter-

facial contact area, and therefore, result in variation of the

capacitive measurements. More specifically, small deflections of

circular membranes can be mathematically predicted according

to the classic thin-plate theory,31 while elastic deformation of the

separation layer is well adapted to the linear strain–stress rela-

tionship. As aforementioned, the unit area capacitance (co) of the

EDL can be considered as an experimentally determined

constant. Therefore, the overall interfacial EDL capacitance

(CEDL) can be simply calculated from the product of the unit area

capacitance and the droplet–electrode contact area. The

mechanical-to-electrical sensitivity of the sensing device can be

analytically expressed as:

DCEDL

P¼ co

�aR2 þ bH

�VdR2

H2(1)

Fig. 2 (a) Illustration of the droplet-based interfacial capacitive sensing

principle and (b) the equivalent electrical circuit diagram.

1112 | Lab Chip, 2012, 12, 1110–1118

where a and b represent the membrane deflection and the elastic

deformation of the separation layer, and can be determined by

the geometrical and mechanical properties of the sensing

membrane and the separation layer, respectively. (The detailed

explanations and derivations are included in the ESI†.) R and H

represent the radius and height of the sensing chamber, while Vd

indicates the volume of the electrolyte droplet. As can be seen,

both the membrane deflection (the 1st term) and the separation

layer deformation (the 2nd term) contribute to the overall

mechanical-to-capacitive sensitivity. It is also worth noting that

both the hydraulic pressure of the droplet and the radius of the

curvature at the droplet–electrode interface have been ignored in

the simplified mathematical expression. In addition, the gravi-

tational effect has been neglected, since the droplet volume is

typically confined to the ones smaller than the cube of capillary

length.32

Materials and methods

Surface modification

Indium-tin-oxide (ITO)-coated (100 nm thick) polyethylene

terephthalate (PET) films (of 125 mm in thickness, Southwall

Technology) were used to serve as the pressure sensing

membranes as well as to establish the interfacial droplet–elec-

trode contact in the device. The ITO coated substrate exhibited

excellent optical clarity (the transmission coefficient of 80.9% at

visible wavelengths), high electrical conductance (the sheet

resistance of 50 U ,�1), and strong mechanical properties

(Young’s modulus of 3–4 GPa).33 In order to reduce the

hysteresis of the droplet deformation, simple surface modifica-

tion was introduced to the conductive coating. In brief, the ITO-

coated PET sheets were exposed to oxygen plasma at 30W for 30

s for surface hydroxylation. Following the plasma activation,

a PDMS stamp, made from a mixture of a base and a curing

agent at 10 : 1 weight ratio (Sylgard 184, Dow Corning) and

thermally cured at 80 �C for 2 hours, was brought into physical

contact with the conductive coating activated with hydroxyl

groups for 2 hours, during which a nanometre-thick layer of

PDMS oligomers was transferred and immobilized onto the

electrode surface.34–36 In addition, an indented hole (of 200 mm in

diameter) was included in the center of the PDMS stamp, leaving

a hydrophilic spot surrounded by the hydrophobic oligomer

layer on the electrode surface, by which the sensing droplet could

be anchored and stabilized in the middle of the sensing chamber.

For the sake of enhancement of the surface hydrophobicity and

thereby reduction of the droplet response time, a super-

hydrophobic nanocomposite material (PFC M1604V, Fluo-

roPel�) was spray-coated onto the PDMS-oligomer-modified

ITO surface. The commercial superhydrophobic coating was

measured with water contact angle (CA) of 155�.

Device fabrication

Direct laser micromachining was employed to fabricate the

droplet-based sensors, which provided a facile approach to form

the geometrical shape of the sensing membranes as well as

the PDMS separation layer containing a sensing chamber in

one single step. Specifically, a desktop CO2 pulsed laser

engraver (VersaLaser, Universal Laser) with graphic user

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interface (e.g., CorelDraw or Photoshop) was used to conduct

the laser-etching process. Various power intensities from 0.3 W

to 3 W were applied to trim ITO-coated PET or PDMS

substrates of different thicknesses, from which a minimal feature

resolution of 100 mm can be reliably achieved.37 Two ventilation

channels were engraved through the PDMS separation layer to

maintain the pneumatic pressure balance during the chamber

deformation. In the subsequent step, the laser-trimmed PET

substrates with ITO coating were imprinted with PDMS oligo-

mers as aforementioned and bonded to the PDMS separation

layer through an oxygen plasma-assisted bonding (at 90 W for

30 s). Prior to the final assembly of the device, an electrolyte

droplet with a desired volume was dispensed by a micropipette in

the centre of the sensing chamber.

Electrical and mechanical characterization

Interfacial capacitance of the EDL layer can be assessed elec-

trically by using an LCR meter (4284A, Agilent). At an AC

excitation voltage of 0.5 V, the interfacial capacitor was con-

nected to the LCR meter in a bipolar configuration, sweeping

from 20 Hz to 100 kHz for acquiring sub-MHz responses of the

EDL layer. A high-speed digital camera (EXILIM EX-F1,

Casio) was used to capture the shape change of the droplet in

order to measure its transient mechanical response (with up to

1200 frames per second). Snapshots of the dynamic droplet

responses to mechanical stimuli are shown in Fig. S1 in the ESI†.

External mechanical point loads were applied onto the centre of

the sensing membrane through a custom-built motorized force

gauge with 1 mN resolution (DFS, Chatillon), driven by

a computer-controlled step motor (VT-80, Micos) with a spatial

resolution of 0.2 mm. The mechanical-to-capacitive responses

were evaluated twice on two identical devices for each parameter.

For the resolution measurement, minute droplets (of 50 mL

volume) dispensed by a micropipette were applied directly onto

the sensing membrane till a noticeable capacitive change

appeared in the LCR meter, which was evaluated three times for

the same device.

Results

Sensing droplets

As a vital part of the interfacial sensing device, physical prop-

erties of the sensing droplet have to satisfy the following criteria:

high ionic concentration (ensuring high electrical conductance

and interfacial capacitance), polarized molecular structure

Table 1 Summary of physical properties of the electrolyte/glycerol mixture asurfaces (given the electrolyte solution of NaCl at the concentration of 1.1 m

Mixing ratio ofelectrolyte/glycerol (v/v%)

Electricalconductivity

Unit-areacapacitance c0/mF cm�2

100/0 Y 4.5 � 0.175/25 Y 4.7 � 0.250/50 Y 4.4 � 0.325/75 Y 4.4 � 0.20/100 N —

a Experiments were conducted at 46% humidity and 24 �C. b Pure glycerol is

This journal is ª The Royal Society of Chemistry 2012

(reversible elasticity on the hydrophobic surfaces) and low fluidic

viscosity (allowing rapid mechanical response). Aqueous-based

electrolyte solution (e.g., NaCl) with high ionic concentration

can be a natural choice, satisfying all except for moderate

evaporation under room conditions. According to a previous

study, mixing an aqueous solution with glycerol can effectively

reduce evaporation due to decreased vapor pressure.38 However,

glycerol is electrically non-conductive and has substantially

higher viscosity than that of water (about 1400 times). Therefore,

an optimal mixing ratio of glycerol and electrolyte solution needs

to be determined. Table 1 summarizes the physical and chemical

properties of the electrolyte/glycerol mixture (given a NaCl

solution at the concentration of 1.1 mol L�1) at various mixing

ratios (v/v%), which includes electrical conductivity, unit-area

capacitance (c0) at 20 Hz, evaporation, viscosity (m), relaxation

time (t) and contact angle (q) (all measured on PDMS oligomer-

coated ITO surfaces). As can be seen, the mixing ratio has

a minimal influence on electrical conductivity and unit-area

capacitance, except for pure glycerol. Importantly, the 25/75%

mixture of electrolyte/glycerol presents no appreciable evapora-

tion under a regular laboratory environment (46% humidity and

24 �C) after 24 hours, as compared to a pure glycerol sample,

which becomes hygroscopic under the same condition. More-

over, adding more glycerol to the mixture drastically increases

the dynamic viscosity as reported previously,39 from 1.0 to 1412

Pa s, and the 25/75% mixture becomes 60 times more viscous

than water, which considerably affects the relaxation time

(increasing 0.02 s to 0.28 s).40 In addition, glycerol and water

have similar contact angles on the hydrophobic coating (i.e.,

PDMS oligomer-coated ITO surface), and as a consequence,

various mixtures have similar contact angles on the hydrophobic

surface. In summary, considering the physical properties and

evaporative stability of the mixtures, the 25/75% electrolyte/

glycerol solution is chosen as the working fluid for the sensing

droplet, which is non-evaporative under laboratory conditions.

Surface modification

As aforementioned, surface hydrophobic treatment has been

applied to the conductive coating, which ensures the reversible

and elastic deformability of the sensing interface. As previously

reported, anultrathin layer ofPDMSoligomers can be universally

transferred and immobilized on a hydroxylated surface. The

interfacial oligomer layer has reportedly possessed a nanometre

thickness (of 1–2 nm),35,41 which would not significantly alter the

EDL upon electrolyte–electrode contact. And meanwhile, it

t various mixing ratios (v/v%), measured on PDMS oligomer-coated ITOol L�1)

EvaporationaViscositym28/Pa s

Relaxationtime t/s

Contactangle q/�

Y 1.0 0.02 90.5 � 0.7Y 2.5 0.06 89.5 � 0.7Y 8.4 0.18 � 0.02 86.0 � 2.8N 60.1 0.28 � 0.02 89.5 � 3.5Nb 1412.0 2.29 � 0.02 93.0 � 1.4

hygroscopic.

Lab Chip, 2012, 12, 1110–1118 | 1113

Fig. 3 Frequency responses of the electrolyte/glycerol mixture on the

unmodified (ITO), the hydroxylated (ITO-OH), and PDMS oligomer-

bonded ITO surfaces (ITO-PDMS) with a control experiment of a pure

electrolyte solution on the same oligomer-coated ITO surfaces. In addi-

tion, the contact angle measurements on various ITO surfaces with

different treatments are illustrated in the inset photos.

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presents substantially reduced interfacial energy.42 The inset in

Fig. 3 shows the variation in surface energy during the oligomer

transfer process, by measuring the contact angle of a mixed elec-

trolyte/glycerol solution (25/75%).As canbe seen,with anoxygen-

plasma treatment, the ITO coating renders apparent hydrophi-

licity (the contact angle reduces from 66� to 15�). Following the

oligomer transfer step, the PDMS oligomer-coated ITO surface

switches the polarity from hydrophilicity to hydrophobicity with

the contact angle of 90�, which further enhances the elasticity and

reversibility of the droplet–electrode contact by reducing adhesive

energy of the liquid to the substrate.

Interfacial capacitance

The EDL capacitance, established by mobile electrons at the

conductive solid surface and a counter-ion layer accumulated in

the electrolyte solution, is frequency-dependent in nature with

several mechanisms associated (e.g., electrophoresis and inter-

facial polarization).24,43 Fig. 3 compares the frequency responses

of the electrolyte mixture on the original ITO surfaces and on the

following modifications. As expected, the surface modifications

have marginal influence on the interfacial EDL capacitance, and

it is most likely that neither surface hydroxylation nor oligomer

transfer alters the physical separation between the electrolyte and

electrode interface. Moreover, in a control experiment, a pure

electrolyte solution (given NaCl at the concentration of 1.1 mol

L�1) exhibits a similar frequency response to that of the elec-

trolyte mixture on the PDMS oligomer modified surface within

the sub-MHz range, which reconfirms that the mixture maintains

reasonably good electrical property at the interface. We have also

noticed that the unit area capacitance measured in this study is

consistently lower than that previously reported, this is possibly

due to the molecular structure of the ITO layer prepared by

different coating techniques.

1114 | Lab Chip, 2012, 12, 1110–1118

Mechanical-to-capacitive sensitivity

As presented in the Operating principle, the overall mechanical-

to-electrical sensitivity (DCEDL/P) can be determined by the

geometrical confinements (the radius R and height H of the

sensing chamber and the thickness of the polymeric membrane),

droplet volume (Vd), and the material properties of the construct

(Young’s modulus and Poisson’s ratio), given a fixed unit-area

capacitance (co). Among those, the radius of the sensing chamber

plays the most important role (4th power in membrane deflection)

as the theory predicted, followed by the height of the chamber

(inverse 2nd power in membrane deflection). In addition, the

system sensitivity is directly proportional to the volume of the

sensing droplet. Experimental investigations have been con-

ducted to verify the above theoretical predications. All the

sensitivity measurements are plotted as a function of capacitive

change, measured by a digital LCR meter, versus the external

pressure applied by a motorized force gauge, as described in

Methods.

Influence of chamber radius

Fig. 4a and b illustrate the capacitive changes over a wide spec-

trum of pressures loaded on the droplet sensors with different

sizes of the sensing chamber, of which the radius varies from 1.5

to 9.0 mm, given the chamber height of 200 mm and the droplet

volume of 0.3 mL. The experimental measurements (dots) are

plotted in comparison with the values calculated from eqn (1)

(curves), and the slope rate of each device measurement defines

the corresponding device sensitivity. As predicted by the sensing

theory, the radius exhibits 4th power in membrane deflection and

2nd power in elastic deformation, which has a significant effect on

the system sensitivity. Within the small deflection limit, the

capacitive charges change linearly with the external load as one

would expect. In the devices with the largest sensing chamber (of

9.0 mm radius), the highest sensitivity of 90.2 nF kPa�1 has been

achieved, as shown in Fig. 4a. In comparison, as the radius is

reduced to two-thirds of the former one (6.0 mm), the system

sensitivity drops drastically to less than 20% of that

(17.2 nF kPa�1) and closely correlates with the theoretical

prediction, i.e., the sensitivity is directly proportional to the 4th

power of radius, as the membrane deflection dominates the

overall mechanical deformation. Those data suggest that

the theoretical model fits the experiments reasonably well under

the assumptions. In both cases, the contact area between the

droplet and the deformed membrane is smaller than 5% of the

whole chamber area, and therefore can be approximated as a flat

interface instead of curved. However, in the smallest devices

(with a radius of 1.5 mm), the measurements deviate consider-

ably (more than 40%) from the stimulated values. In this case it is

highly possible that the small deformation limit, which is asso-

ciated with the ratio of the maximum deflection to the radius of

the membrane, has been exceeded.44 In addition, it has been

observed that the radius of the curvature of the deflected

membrane can no longer be ignored in the smallest sensing unit.

Overall, the radius of the chamber is a determinant factor for the

overall mechanical-to-electrical sensitivity, given its 4th power

relation, as the membrane deflection serves as the primary

mechanical deformation mechanism under external load.

This journal is ª The Royal Society of Chemistry 2012

Fig. 4 Influence of geometrical confinements on the device sensitivity: (a) different radii of the sensing chambers of 6.0 and 9.0 mm; (b) different radii of

the sensing chambers of 1.5 and 3.0 mm; (c) different heights of sensing chambers (from 140 mm to 350 mm) and (d) different volumes of droplets (from

0.3 mL to 3.6 mL).

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Influence of chamber height

As predicted by the theoretical analysis, the device sensitivity is

inversely proportional to the 2nd power of the height of the

sensing chamber as the membrane deflection dominates, and

thus, a lower sensing chamber will lead to a larger contact area at

a given droplet volume. Fig. 4c presents the variations of the

sensitivity over four different chamber heights ranging from

140 mm to 350 mm, provided the chamber radius of 6.0 mm and

droplet volume of 3 mL. As can be seen, the lowest chamber

(140 mm in height) shows the highest sensitivity (374 nF kPa�1)

due to the presence of the largest contact area at the droplet–

electrode interface. As the height of the chamber rises from

140 mm to 210 mm, the device sensitivity decreases from

374 nF kPa�1 to 144 nF kPa�1, closely following the negative

quadratic relationship between the sensitivity and the chamber

height. The highest chamber (of 350 mm) exhibits the minimal

sensitivity of 54.5 nF kPa�1. Moreover, as the chamber height

increases, the separation layer becomes increasingly influential

on the overall sensitivity. In other words, the thicker separation

This journal is ª The Royal Society of Chemistry 2012

layer produces more substantial deformation under the same

load according to the linear strain–stress relationship.45

However, in our cases, the maximal deformation of the separa-

tion layer is still one order of magnitude smaller than that of the

membrane deflection as expected. In brief, all four sets of

the sensitivity measurements are in a good agreement with the

theoretical analysis presented above.

Influence of droplet volume

Furthermore, the droplet volume is directly proportional to the

system sensitivity. In general, a larger droplet covers more

interfacial area with a longer circumferential periphery along the

contact, which leads to a higher sensitivity given constant

dimensions of a sensing chamber. However, the increased droplet

size can also lead to non-linear response as the droplet–electrode

interface can no longer be approximated as a planar surface.

Four different droplet sensing units (of 0.3 mL, 0.6 mL, 1.2 mL

and, 3.6 mL, respectively) have been characterized and compared

Lab Chip, 2012, 12, 1110–1118 | 1115

Fig. 5 (a) The response times of the droplet sensors with decreased droplet viscosity (using pure electrolyte) and improved surface hydrophobicity

(using a superhydrophobic coating) in comparison with a control group (electrolyte/glycerol 25/75% (v/v) mixture on the oligomer-coated ITO surface);

(b) the humidity influence on the initial electrolyte–electrode capacitance; and (c) the temperature influence on the unit-area interfacial capacitance.

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using identical dimensions of the sensing chamber (6.0 mm

radius and 200 mm height), as shown in Fig. 4d. As the droplet

volume increases from 0.3 mL to 0.6 mL, and then to 1.2 mL, the

corresponding sensitivity rises in a linearly proportional manner

from 17.2 nF kPa�1 to 32.1 nF kPa�1, and then to 65.7 nF kPa�1.

The highest sensitivity (218 nF kPa�1) has been achieved on the

sensing droplet of 3.6 mL in volume. Again, this matched finding

reconfirms the applicability of eqn (1).

Optimization of device sensitivity

Finally, we have combined the optimal conditions from the

above analysis (of 9.0 mm in radius, 140 mm in height and 3 mL in

droplet volume) and characterized the highest device sensitivity

of 1.58 mF kPa�1, in comparison to the highest value reported in

the literature (2.24 mF kPa�1, using a mercury droplet and an

ultrahigh permittivity material), to the best of our knowledge.14

In addition to the design parameters discussed above, the

thickness of the polymeric membrane can be crucial in deter-

mining the sensing performance. As shown in eqn (S1) in the

ESI†, the sensitivity caused by membrane deformation is

inversely proportional to the cubic power of its thickness. In this

study, we have been using the off-the-shelf ITO-coated substrate

(125 mm in thickness); however, further reducing the membrane

thickness could lead to drastic increase of the device sensitivity.

Moreover, in the optimal sensing design, the device shows an

extremely high pressure resolution of 1.8 � 0.4 Pa.

Response time

Response time serves as another critical measure for mechanical

sensing devices in addition to the sensitivity and resolution. In

general, the response time of such a droplet-based sensor can be

influenced by the viscosity of the droplet medium, the hydro-

phobicity of the electrode surface, and the elasticity of the PDMS

separation layer and the PET membrane, of which the liquid

viscosity and surface hydrophobicity dominate. Accordingly, we

have further characterized the response time of the geometrically

identical sensing devices with varied droplet viscosity and surface

1116 | Lab Chip, 2012, 12, 1110–1118

hydrophobicity under the same level of membrane deformation

of 50 mm. The findings are plotted in Fig. 5a. As can be seen, the

response time is around 260 ms in an optimized sensing design

with a chamber of 6 mm in radius and 140 mm in height and the

volume of the droplet is 3 mL as described previously; whereas the

decreased droplet viscosity (from 60 Pa s of the 25/75% electro-

lyte/glycerol mixture to 1 Pa s of a pure electrolyte solution) and

the enhanced surface hydrophobicity (from a CA of 90–155�) canlead to more than two-fold reduction in the response time (to

around 100 ms). However, the pure electrolyte solution is subject

to evaporative effects, while converting the electrolyte surface

from hydrophobic to superhydrophobic introduces a micro-

metre-thick (around 10 mm) dielectric layer, which substantially

lowers the interfacial capacitance.46

Influences of temperature and humidity

Importantly, the performance of the droplet sensors can be

highly subject to the environmental humidity level. Specifically,

the electrolyte/glycerol mixture establishes thermodynamic

equilibrium with different stable mixing ratios at different

humidity levels by either losing (evaporation) or gaining

(condensation) aqueous contents to/from the environment.47,48

As shown in Fig. 5b, when the humidity level alters, the volume

and capacitance of a droplet sensor can substantially deviate

from the initial condition (of 25/75% electrolyte/glycerol) as the

mixture compositions change. In other words, as the humidity

level rises, the volume of the droplet would likely increase as

additional moisture condensates into the mixture. However, once

a new thermodynamic equilibrium is established, the capacitance

will remain stable. Therefore, we can either employ sensing

elements with multiple mixing ratios for different humidity levels

or let the sensing element equilibrate with the environment

thermodynamically prior to its usage.

In contrast, the environmental temperature fluctuation has

only posed minor influence on the interfacial capacitance, in

comparison with the resistance-based devices.12,49 As shown in

Fig. 5c, the electrolyte–electrode interfacial capacitance has been

This journal is ª The Royal Society of Chemistry 2012

Fig. 6 (a) Applying the droplet-based ultrahigh sensitive sensor (b) to

continuously record the blood pressure variations on the skin surface

(above the carotid artery).

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measured at three separate temperatures, 4 �C (reduced), 25 �C(regular), and 45 �C (elevated), respectively, from which less than

10% variation in the unit-area capacitance has been observed at

the steady state. In comparison, within the same range of

temperature fluctuation, the resistance measurement can change

up to 55% according to a previous report.49

Demonstration

To demonstrate the utility of the simply constructed and

mechanically flexible droplet sensor with ultrahigh sensitivity

and resolution, we applied the device to record blood pressure

variations on the skin surface. In this dynamic measurement,

a droplet sensor was devised with a built-in chamber size of 6 mm

in radius and 140 mm in height. An electrolyte droplet of 3 mL

acted as the sensing element to achieve a facilitated response

(with the response time of 100 ms).

The sensor was attached to the skin surface above the carotid

artery (Fig. 6a) to record the blood pressure wave under a gentle

contact force being applied. As shown in Fig. 6b, the minute

blood pressure variations were recorded with the maximal value

less than 100 Pa at around 1 Hz. As demonstrated, the ultrahigh

sensitivity/resolution, simple configuration and flexible construct

of the droplet sensing device could be highly attractive to a wide

range of biomedical applications (e.g., ocular, pulmonary, and

cardiovascular systems), where the maximal pressure variations

typically range from 2 kPa to 20 kPa and meanwhile, the body

conformability and comfortableness are of particular importance

for continuous recording.50–52

Conclusions

In this paper, we have presented a novel droplet-based interfacial

capacitive sensor with a very simple architecture and tunable

sensitivities. The interfacial sensing principle utilizes the presence

of a large unit-area EDL capacitance at the elastic droplet–

electrode contact. A theoretical model has been proposed to

establish the linkage between the geometrical parameters (the

radius and height of the sensing chamber and the droplet

volume) and the device sensitivity. The interfacial droplet-based

sensing device offers several distinct advantages over the existing

counterparts: (1) ultrahigh mechanical-to-capacitive sensitivity

(1.58 mF kPa�1) and resolution (1.8 Pa) (as comparable to the

This journal is ª The Royal Society of Chemistry 2012

highest reported value),14 (2) simple fabrication (one-step laser

micromachining), (3) mechanical flexibility, (4) optical trans-

parency (PDMS, ITO/PET and electrolyte/glycerol), (5) insen-

sitivity to evaporation (using electrolyte/glycerol mixture) and (6)

thermal noises (capacitive sensing instead of resistive-based

schemes). Moreover, the environmental influences (e.g.,

temperature and humidity) on the capacitive measurement have

been further investigated. The proposed interfacial capacitive

sensing scheme is expected to be employed in a wide range of

biomedical applications, where ultrahigh sensitivity and extreme

flexibility are concurrently required.

Acknowledgements

This work is in part supported the NSF CAREER Award

(ECCS-0846502) and EFRI Award (EFRI-0937997) to TP. BN

and SX acknowledge the fellowship support from China Schol-

arship Council (CSC). Authors would also like to acknowledge

the samples of ITO-coated membranes generously provided by

Southwall Technology. In addition, we would like to thank Mr

Yuzhe Ding for the valuable discussion on the blood pulse

measurement.

Notes and references

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This journal is ª The Royal Society of Chemistry 2012