Novel hybrid optical sensor materials for in-breath O2 analysis

Novel hybrid optical sensor materials for in-breath O2 analysis

Clare Higgins, Dorota Wencel, Conor S. Burke, Brian D. MacCraith and Colette McDonagh*

Received 19th October 2007, Accepted 5th November 2007

First published as an Advance Article on the web 28th November 2007

DOI: 10.1039/b716197b

This study focuses on the optimisation and characterisation of novel, ORganically MOdified

SILicate (ORMOSIL)-based, hybrid sensor films for use in the detection of O2 on a breath-by-

breath basis in human health monitoring applications. The sensing principle is based on the

luminescence quenching of the O2-sensitive ruthenium complex [Ru(II)-tris(4,7-diphenyl-1,10-

phenanthroline)], which has been entrapped in a porous sol–gel film. The detection method

employed is that of phase fluorometry using blue LED excitation and photodiode detection.

Candidate sensor films include those based on the organosilicon precursors,

methyltriethoxysilane, ethyltriethoxysilane, n-propyltriethoxysilane and phenyltriethoxysilane.

While it has been established previously by the authors that these films exhibit a stable, highly

sensitive response to O2, this study focuses on selecting the material most suited for use in a breath

monitor, based on the sensitivity, response time and humidity sensitivity of these films. Key

parameters to be optimised include the O2 sensitivity of the film and the film polarity, i.e. the

degree of hydrophobicity. These parameters are directly linked to the precursors used. In this

study a n-propyltriethoxysilane-derived O2 sensor platform was selected as the optimum material

for in-breath O2 analysis due to its short response time, negligible humidity interference and

suitable O2 sensitivity in the relevant range in addition to its compatibility with a single-point

calibration strategy.

Introduction

The diagnostic value of O2-in-breath analysis is well estab-

lished.1 Breath analysis is attractive as it permits non-invasive

monitoring of a patient’s state of health.2 Despite the benefits

of breath analysis, the use of breath gas analysers is typically

restricted to a laboratory environment due to the expensive,

cumbersome nature of most breath analysis systems. The

development of a portable, low-cost system is attractive as

such a system would be compatible with home healthcare and

advanced athletic performance monitoring applications. A key

requirement for such a breath monitoring device is a

disposable sensor element, capable of operating at a range of

breath rates in high humidity, which is incorporated into a

lightweight, compact sensing module.

Luminescence-based O2 sensors have been extensively

researched in recent years due to their advantages over the

Clark electrode.3 These sensors combine the intrinsic sensitiv-

ity of the luminescence process with the wide availability of

low-cost, discrete optoelectronic components, thereby enabling

a range of sensor configurations and facilitating design

features that are desirable in biomedical applications, such as

miniaturisation and disposability. Sol–gel materials are an

ideal sensor matrix, due to their ease of fabrication and the

versatility of the process. The sol–gel process is compatible

with the use of deposition techniques such as ink-jet- and pin-

printing,4–6 which enable deposition onto a range of planar

and non-planar substrates. This paper addresses some key

issues relevant to the successful implementation of a portable

in-breath O2 sensor platform. These issues include response

time, influence of humidity on sensor response and enhanced

sensitivity in the range of interest for O2 sensing. A range of

ORganically MOdified SILicate (ORMOSIL)-based O2

sensor materials are reported where the O2-sensitive complex

[Ru(II)-tris(4,7-diphenyl-1,10-phenanthroline)], [Ru(dpp)3]2+,

is entrapped in the ORMOSIL xerogel. The O2 sensitivity of

such systems is dependent on the choice of precursor, which

influences the relative hydrophobicity of the resulting material,

thereby impacting on O2 transport into the material. This

paper reports on the effect of various organosilicon precursors

on the response time, relative hydrophobicity and O2

sensitivity of the resulting sensor elements.

Theory

Oxygen-sensitive materials comprised of an O2-sensitive

luminophore immobilised in an optically transparent

O2-permeable host matrix have been reported widely in the

literature.7–16 Sensing using these elements is attractive as the

sensor does not consume O2 and is not prone to electrical

interference. Such sensor elements are reported here. The

sensing mechanism centres on the dynamic quenching of

the excited-state luminescence lifetime (and, correspondingly,

the luminescence intensity) of the luminophore by O2

molecules. If the luminescence quenching is purely dynamic,

the excited-state lifetime and the intensity are related to the O2

concentration. This process is governed by the Stern–Volmer

relation,17 given in eqn (1),

Optical Sensors Laboratory, National Centre for Sensor Research,School of Physical Sciences, Dublin City University, Dublin 9, Ireland.E-mail: [email protected]; Fax: +353 1 700 8221;Tel: +353 1 700 5301

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This journal is � The Royal Society of Chemistry 2008 Analyst, 2008, 133, 241–247 | 241

I0

I~

t0

t~1zKSV O2½ � (1)

where I is the luminescence intensity, t is the excited-state

lifetime of the luminophore, [O2] is the O2 concentration, KSV

is the Stern–Volmer constant and the subscript 0 denotes the

absence of O2.

KSV is a measure of the sensitivity of a sensor element and is

related to t0 by eqn (2) (where kq is the diffusion-dependent

bimolecular quenching constant) and to the diffusion coeffi-

cient, D, by eqn (3):

KSV = t0kq (2)

kq = 4pgRND (3)

In eqn (3), g is the spin statistical factor, R is the collision

radius, and N is Avogadro’s number. From eqn (2) and eqn

(3), it is clear that KSV is proportional to D, all other factors

being constant. Consequently, the sensitivity of the sensor

layer is increased by improving O2 transport through a

material.

From eqn (1), for an ideal, homogeneous environment, a

plot of the ratio of t0/t as a function of [O2] yields a straight

line with an intercept at 1 and a slope of KSV. In the case of

luminophores entrapped in a solid matrix, these plots deviate

from linearity.15,18–22 This behaviour is associated with the

distribution of the luminophore within the solid matrix. The

host micro-heterogeneity causes luminophore populations in

different sites to be quenched differently with a resultant

downward curve in the Stern–Volmer plot. Various models

have been developed to describe quenching data that deviate

from the Stern–Volmer ideal.23–25 For the sensor systems in

this work, the Stern–Volmer [eqn (1)] and two-site Demas

model [eqn (4)]7 were employed to analyse the O2 data. The

Demas model assumes that the luminophore population

distributed in the solid material resides at two different site

types, with each site exhibiting a particular quenching

constant.

I0

I~

t0

t~

f1

1zKSV1 O2½ �z

f2

1zKSV2 O2½ �

� �{1

(4)

In eqn (4), fi represents the fractional contribution of the

total emission from the luminophores located at site i (under

unquenched conditions) that exhibit a discrete Stern–Volmer

constant given by KSVi.

Excited-state lifetime measurements offer certain

advantages with respect to the performance of the sensors.

As lifetime is an intrinsic property of the luminophore,

instrumental fluctuations, leaching and photobleaching of

the luminophore do not affect the sensor performance.

Therefore, the long-term stability of such sensors is much

improved compared with that of luminescence intensity-based

sensors.

In this work, phase fluorometry is employed to indirectly

monitor t. The excitation signal is sinusoidally modulated and

the luminophore’s emission is also modulated but is time-

delayed or phase-shifted relative to the excitation signal. The

relationship between t and the corresponding phase shift, w,

for a single exponential decay is given by eqn (5),

tan w = 2pft (5)

where f is the modulation frequency of the LED excitation

source. Therefore, phase fluorometry facilitates the indirect

monitoring of the luminescence excited-state lifetime, thereby

avoiding the high cost and time-consuming data processing

issues associated with direct lifetime measurements.26,27 This

capability can be provided using low-cost instrumentation and

straightforward data analysis, making it a more attractive

candidate for sensor development compared to direct lifetime

measurements.

The support medium in this work is produced via the sol–gel

process. Materials prepared using sol–gel technology can range

from simple inorganic glasses (e.g. when using tetraethoxysi-

lane, TEOS) to more complex inorganic–organic materials

called ORMOSILs.28–30 These materials facilitate the develop-

ment of a broad range of innovative materials and are of the

general form R4 2 xSi(OR9)x, where R represents an organic

group, such as –CH3, –C2H5, –C3H7, –C6H5, (–CH2)nNH2. R

is generally bound to the silicon via a Si–C bond that is not

hydrolysable and these silanes are often used as network

modifiers and network formers. They can combine the

advantages of the organic and inorganic constituents within

the same matrix and the physico-chemical properties can be

tailored through considered choice of R. Therefore, one can

fabricate materials with the desired hydrophobicity, flexibility

and stability. Sensor technology is an area in which the use of

these sensor matrices is becoming increasingly prevalent due to

the reliable, versatile, low-temperature nature of the sol–gel

technique.31,32 The organosilicon precursors selected for use

here have the effect of increasing the relative hydrophobicity of

the surface of the resulting material and also reducing the

connectivity of its microstructure. Both of these factors

contribute to improved O2 transport within the resulting

material, increasing D and thereby contributing to an

increased KSV. We will show that the increased hydrophobicity

also influences humidity sensitivity. This is of particular

relevance when selecting a material for use in a breath

monitoring application, due to the high level of humidity in

breath.

Materials and methods

Chemical reagents

RuCl3?3H2O, 4,7-diphenyl-1,10-phenanthroline ligand, abso-

lute ethanol (EtOH), 0.1 M HCl and the organosilicon

precursors methyltriethoxysilane (MTEOS), ethyltriethoxysi-

lane (ETEOS), n-propyltriethoxysilane (PTEOS) and phenyl-

triethoxysilane (PhTEOS), were purchased from Aldrich

Chemicals. All chemicals were used as received.

[Ru(dpp)3]2+was synthesised as described in the literature.33

Glass microscope slides were soaked for 24 h in HNO3, and

then rinsed with deionised water and EtOH before drying

under a N2 flow. All experiments were performed at room

temperature.

242 | Analyst, 2008, 133, 241–247 This journal is � The Royal Society of Chemistry 2008

Fabrication of sol–gel-based sensor platforms

All sol–gel-based films were prepared from sols containing the

relevant precursor, EtOH as co-solvent, HCl and the

[Ru(dpp)3]2+ complex at a concentration of 2.5 g L21 with

respect to the total volume of sol. For all sols, the final molar

ratio of silane : EtOH : water : HCl was 1 : 6.25 : 4 : 0.007. All

films used in this study were dip-coated in a controlled

environment using a computer-controlled dipping apparatus.

After deposition onto microscope slides, the films were dried at

110 uC for 18 h. Profilometry measurements showed that the

dip-coated sensor films were 400–500 nm in thickness.

Phase fluorometry instrumentation

The principle of phase fluorometry and the experimental

system used to examine the performance of the O2 sensors

have been described elsewhere.34,35 Briefly, the characterisa-

tion system consisted of a blue LED (Nichia, NSPE590,

Japan), which was modulated at a frequency of 20 kHz and

provides excitation of the luminescent sensors. A silicon

photodiode (Radionics, 194-290, Ireland) was used for the

detection of the O2-sensitive luminescence signal. In order to

examine their performance, O2-sensitive films were placed in a

flow cell into which controlled mixtures of O2 and N2 were

flowed using mass flow controllers (Celerity, Ireland). In order

to test the effect of humidity on O2 sensitivity, two gas wash

bottles containing deionised water were included in the gas line

before the gas mixture entered the flow cell, and a commercial

humidity probe (Testo 625, Germany) was employed to

monitor the humidity in the cell. To determine the effect of

temperature on the O2 sensitivity, a gas heater (Radionics, 200-

2496, Ireland) was included in the gas delivery setup.

Other characterisation techniques

In order to measure the intrinsic response time of the sensor

films and not that of the gas handling system (i.e. the time

taken to fill the gas cell and lines), the gas exchange times must

be minimised. To this end, measurement instrumentation was

developed that incorporated a fast solenoid valve capable of

5 ms switching times. Details of this characterisation system has

been reported previously.36

The oxygen diffusion coefficient within the films was

determined in order to establish the origin of the observed

oxygen sensitivity. This was achieved by measuring the

response time for films of known thickness. Sensing layers

with a thickness of 1 mm were used as this was the minimum

thickness compatible with the detection efficiency of the

response time measurement apparatus. Diffusion coefficients

were then obtained during a step change in oxygen concentra-

tion, using the technique reported previously in ref. 36.

The excited-state lifetime of [Ru(dpp)3]2+ immobilised in

each ORMOSIL film was determined using a pulsed laser

system.37 The samples were excited with 15 ns pulses from a

Nd:Yag laser (l = 355 nm). Samples were degassed with N2

prior to measurements. The emitted decay was detected using a

photomultiplier tube and captured using a digital oscilloscope.

The excited-state luminescence lifetime was calculated by

analysing the decay trace using MicroCal Origin.

Water contact angle measurements were used to quantify

the relative hydrophobicity of the surface of ORMOSIL films.

These measurements were carried out using a static contact

angle analyser (FTA 200, USA).

The sensor film thickness was measured using a white light

interferometer (WYKO, NT1100, Veeco, USA).

A commercially available metabolic calibration unit

(VacuMed, USA) was employed to characterise the response

of the sensor film to changes in O2 concentration on a breath-

by-breath basis. This device is essentially a lung simulator,

which mimics O2 consumption through the dilution of an

inspired volume of ambient air with a mixture of carbon

dioxide in nitrogen (referred to as the calibration gas). This

apparatus is described in more detail in ref. 38.

For the work described, a module containing a sensor chip

coated with a PTEOS-based sensor film38 was attached to the

outlet of the lung simulator and its response to changes in O2

concentration was recorded.

With the exception of the sensor platforms used for response

time measurements, all other sensor film characterisation was

carried out using dip-coated films with a thickness in the range

of 400–500 nm.

Results and discussion

All MTEOS-, ETEOS-, PTEOS- and PhTEOS-based O2

sensor platforms exhibit excellent repeatability, reversibility

and short response times. Fig. 1 shows typical phase response

for all sensor elements tested in this study.

The key requirements in the successful implementation of a

sensor platform for in-breath O2 analysis are discussed in the

following sections. Table 1 is a compilation of all character-

isation data of the O2 sensor elements produced in this work.

Response time

Rapid sensor response is an important consideration in

designing a device for breath-by-breath monitoring. To

maximise the diagnostic possibilities and ensure the breath

monitor is suited to many medical applications, a range of

breath rates must be accommodated. A shorter response time

allows for the detection of fast breath rates, thereby increasing

the breathing range over which the sensor can be used. The

response time of a sensor membrane has been defined here as

t90, the time taken to achieve 90% of the optical signal change,

as the environment switches from vacuum to atmospheric

pressure. The sensor response time is dependent on the

diffusion path length of the O2 molecules to the luminophore,

which is directly related to film thickness. Sensing layers with a

thickness of 1 mm were used for the measurement of sensor

response time.

From Table 1 the shortest t90 was found for ETEOS- and

PTEOS-based sensors. The response times of these sensors

were 232 and 223 ms, respectively. The required response time

for this breath monitoring application is 50 ms and can be

obtained by reducing the sensor layer thickness to approxi-

mately 500 nm, meaning that the sensor films produced for all

other characterisation stages reported here were adequately

thin for this application.


The longest t90 was found for the PhTEOS-based sensor

platforms (t90 = 2687 ms) due to the relatively low value of D

for this film. This response time was an order of magnitude

greater than that of any of the other materials tested.

Accordingly, PhTEOS-based sensors were deemed to be

unsuited for this application and were not investigated further.

Influence of relative humidity

Fig. 2 shows the humidity dependence for MTEOS-, ETEOS-

and PTEOS-based O2-sensitive films. Clearly, the more

hydrophobic ETEOS- and PTEOS-based films exhibit negli-

gible sensitivity to humidity. This is consistent with the water

contact angle (CA) data in Table 1. The most hydrophobic

films in this work are ETEOS- and PTEOS-based sensor

elements with a water CA of 97 ¡ 2u and 100 ¡ 1u,respectively. As such, these coatings should be less prone to

humidity interference than, for example, MTEOS-based coat-

ings. These hydrophobic sensor films have a reduced sensitivity

to ambient moisture compared to hydrophilic films.

Since considerable humidity interference was observed for

the MTEOS-based sensor films, this material was consequently

eliminated as a possible sensor matrix for this application.

Sensitivity

It has been demonstrated previously that films with increased

hydrophobicity exhibit higher O2 sensitivity, both in gas and

dissolved phase.39 Table 1 shows that there is a good

correlation between the Stern–Volmer constant (KSV, O2

sensitivity) and both the measured D and water CA. The

parameters KSV, D and CA increase in magnitude as the alkyl

chain length increases. The sensitivity of the MTEOS-based

sensor element is less than that of either the ETEOS- or

Fig. 1 Phase response of MTEOS-, ETEOS-, PTEOS- and PhTEOS-based sensor platforms.

Table 1 Effect of organosilicon precursor on the characterisation data of the O2 sensor platforms

Organosilcon precursor t0 (accuracy ¡ 10%)/ms KSV (0–30% O2)/[O2 %]21 Dw (= w25 2 w15) D(61026)/cm2 s21 t90a/ms Contact angle/u

MTEOS 4.92 0.062 ¡ 0.003 2.7 9.86 432 ¡ 17 85 ¡ 1ETEOS 4.91 0.072 ¡ 0.002 2.4 62.1 232 ¡ 9 97 ¡ 2PTEOS 5.11 0.108 ¡ 0.003 2.6 67.3 223 ¡ 16 100 ¡ 1PhTEOS 5.81 0.023 ¡ 0.001 2.9 0.03 2687 ¡ 161 90 ¡ 1a Normalised response time, representative of a sensor layer thickness of 1 mm.


PTEOS-based sensor elements, with the sensitivity of the

PTEOS-based sensor layer being twice that of the MTEOS-

based sample. As mentioned earlier, O2 sensitivity, represented

by KSV, is dependent on both t0 and D. From Table 1, t0 was

found to remain essentially constant for MTEOS-, ETEOS-

and PTEOS-based sensor materials. While all samples were

stored in a N2 environment, minor deviations in the lifetime

data are thought to be due to slight fluctuations in O2

concentration within the measuring environment due to leaks

in the sample cell. Therefore, we can attribute the enhanced

sensitivity to the increase of O2 transport within the ETEOS-

and PTEOS-based O2 films, as evidenced by the diffusion

coefficients recorded for these films. PhTEOS-based sensor

elements exhibit longer excited-state lifetime (ca. 18%) than

other materials tested but exhibit ca. 56 lower sensitivity than

the most sensitive PTEOS-based films, which is due to the

steric effect of the bulky phenyl group.

While KSV data are presented here over the full O2 range to

facilitate comparison with other O2 work in the literature, for

in-breath O2 analysis, the concentration range of interest is 16–

21%. Therefore, it was necessary to compare the sensitivity of

the films studied over this range. The parameter Dw, was used

as a relative sensitivity parameter to denote the change in

phase response as O2 concentration changed from 15 to 25%

(Dw = w25 2 w15). From the Dw data in Table 1, the most

sensitive film, in this range, was the PhTEOS-based sensor

element. However, this material had been eliminated as a

possible candidate for the application due to its unacceptably

long t90. Bearing this in mind, the most suitable matrix was

clearly PTEOS-based with a Dw value of 2.6 compared with 2.4

for an ETEOS-based sensor. Having identified the optimum

sensor matrix, all further analysis was limited to this matrix

only.

Fig. 3 presents the best fit to the Stern–Volmer and Demas

models for PTEOS-based sensor elements. The fit to the

Demas model is superior (r2 = 0.999). The Stern–Volmer plot

for PTEOS-derived sensor platforms are also very well

described by eqn (1) (r2PTEOS = 0.997), which would allow

for a simple one-point calibration of these sensors if desired.

Influence of temperature on O2 sensor response

Temperature effects on the response of the PTEOS-sensor

films have been tested. The phase response was recorded as

Fig. 2 Influence of humidity on O2 response for MTEOS-, ETEOS- and PTEOS-based sensor platforms. The data were recorded at 0% and 95%

RH.


described earlier, using gas at temperatures of 20, 30 and 40 uC.

This temperature range was selected as it reflects the working

range required of a breath monitor. A change in temperature

results in a change in t0 of the [Ru(dpp)3]2+ complex. The

diffusion coefficient of the sol–gel matrix is also temperature

dependent, with higher temperatures resulting in faster

diffusion, leading to an increase in collisional quenching. The

results in Fig. 4 show that sensor performance does not deviate

greatly and the sensors still operate sensitively within the tested

temperature range. However, as expected, temperature cali-

bration will be required for these sensor platforms.

Optimum material for O2-in-breath sensor platform

PTEOS-based sensor platforms were selected as the optimum

material for use in an O2-in-breath monitor, as they exhibited

sufficient O2 response in the relevant O2 concentration range,

from 15 to 25%, coupled with negligible humidity interference

and short response time. The error bars associated with the

films shown in Fig. 3 are very small (0.03u at 30% O2) and

demonstrate that film-to-film reproducibility is excellent. A

good fit to the linear Stern–Volmer model facilitates one-point

calibration procedures. The O2 response of such a PTEOS-

based sensor in the flow-cell-based instrumentation in this

laboratory yielded an LOD of 0.09% (of full scale) and a

resolution of 0.09% (in the range 10–20% O2). Fig. 5 presents

the phase data obtained using a PTEOS-based sensor element

to monitor the changing O2 levels delivered by the lung

simulator described previously. The simulator was configured

to deliver a 2% reduction in the expired O2 concentration

relative to ambient O2 levels at a breathing rate of 10 breaths

per minute. The ability of the sensor to respond to changes in

O2 concentrations on a breath-by-breath basis is clearly

demonstrated and a resolution of approximately 0.03% O2

was achieved in this concentration range.

While PTEOS has been selected here as the optimum

organosilicon precursor for the production of materials for an

O2-in-breath monitor, it is worth noting that the criteria for

that application are similar to those essential for O2-monitors

in bioprocessing applications. As such, PTEOS-based materi-

als are suited to a broad range of sensor solutions in many

industrial applications.

Conclusions

Oxygen-sensitive xerogel systems based on hydrophobic

ORMOSILs have been fabricated and tested. These sensor

platforms exhibit increased O2 sensitivity with increasing

length of alkyl group and there is a strong correlation between

O2 sensitivity and film hydrophobicity. The sensor response

has been measured using phase fluorometry. The humidity

sensitivity of the O2 sensor films was investigated and it was

shown that the influence of humidity can be virtually

eliminated by increasing film hydrophobicity. Temperature

interference has been examined and we conclude that, while

temperature correction must certainly be implemented for

Fig. 3 Best fit to the Stern–Volmer and Demas model for PTEOS-

based sensor platforms. The calibration is based on the performance of

four sensor films. The slope represents KSV, O2 sensitivity.

Fig. 4 Influence of the temperature of the gas on O2 response for

PTEOS-based sensor platforms. The data were recorded at gas

temperatures of 20, 30 and 40 uC.

Fig. 5 Response of PTEOS-based sensor element. DO2 = 2%,

10 breaths per minute.


ORMOSIL films, it is clear that these sensor films operate to a

sufficiently high sensitivity within the working temperature

range required of a breath monitor. Overall, we demonstrated

that PTEOS-based O2 sensor platforms exhibit characteristics

such as: excellent reproducibility, short response time,

humidity insensitivity, enhanced O2 sensitivity and compat-

ibility with single-point calibration procedures, making them

an ideal choice for breath monitoring applications. Future

work will address the evaluation of sensor performance in

clinical trials.

Acknowledgements

The authors wish to acknowledge Dr Mary Pryce, School of

Chemical Sciences, Dublin City University for assistance with

lifetime measurements.

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