Wearable and flexible electronics for continuous molecular monitoring

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DOI: 10.1039/C7CS00730B
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Abstract
Wearable biosensors have received tremendous attention over the past decade owing to their great potential in predictive analytics and treatment toward personalized medicine. Flexible electronics could serve as an ideal platform for personalized wearable devices because of their unique properties such as light weight, low cost, high flexibility and great conformability. Unlike most reported flexible sensors that mainly track physical activities and vital signs, the new generation of wearable and flexible chemical sensors enables real-time, continuous and fast detection of accessible biomarkers from the human body, and allows for the collection of large-scale information about the individual's dynamic health status at the molecular level. In this article, we review and highlight recent advances in wearable and flexible sensors toward continuous and non-invasive molecular analysis in sweat, tears, saliva, interstitial fluid, blood, wound exudate as well as exhaled breath. The flexible platforms, sensing mechanisms, and device and system configurations employed for continuous monitoring are summarized. We also discuss the key challenges and opportunities of the wearable and flexible chemical sensors that lie ahead.
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Wearable and flexible sweat biosensors. (a) Epidermal temporary-transfer tattoo sweat biosensor for in situ continuous monitoring of lactate, pH, etc. Reproduced with permission from ref. 75. Copyright 2013 American Chemical Society. (b) Wireless sensor bandage for chronological monitoring of sweat Na + concentration. A RFID chip was incorporated for wireless communication with a smart phone. Reproduced with permission from ref. 81. Copyright 2015 IEEE. (c) Fully integrated ''smart wristband'' with a flexible sensor array and a flexible PCB for multiplexed in situ sweat analysis. Reproduced with permission from ref. 55. Copyright 2016 Nature Publishing Group. (d) Flexible sensor array for multiplexed heavy metal monitoring in sweat. Reproduced with permission from ref. 83. Copyright 2016 American Chemical Society. (e) Graphene-based flexible and stretchable diabetes patch for sweat glucose monitoring. Reproduced with permission from ref. 56. Copyright 2016 Nature Publishing Group. (f) A disposable strip for sweat glucose sensing. Reproduced with permission from ref. 84. Copyright 2017 American Association for the Advancement of Science. (g) Stretchable sensor with serpentine structure on porous polyurethane for sweat pH sensing. Reproduced with permission from ref. 86. Copyright 2014 Wiley. (h) Microfluidicbased device for sweat routing and colorimetric pH sensing. Reproduced with permission from ref. 87. Copyright 2012 Elsevier. (i) Stretchable microfluidic based sweat patch for spontaneous sweat routing through serpentine channels and the reservoir, with multiplexed colorimetric sensing. Reproduced with permission from ref. 57. Copyright 2016 American Association for the Advancement of Science.
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Cite this: DOI: 10.1039/c7cs00730b
Wearable and flexible electronics for continuous
molecular monitoring
Yiran Yang and Wei Gao *
Wearable biosensors have received tremendous attention over the past decade owing to their great
potential in predictive analytics and treatment toward personalized medicine. Flexible electronics could
serve as an ideal platform for personalized wearable devices because of their unique properties such as
light weight, low cost, high flexibility and great conformability. Unlike most reported flexible sensors that
mainly track physical activities and vital signs, the new generation of wearable and flexible chemical
sensors enables real-time, continuous and fast detection of accessible biomarkers from the human
body, and allows for the collection of large-scale information about the individual’s dynamic health
status at the molecular level. In this article, we review and highlight recent advances in wearable and
flexible sensors toward continuous and non-invasive molecular analysis in sweat, tears, saliva, interstitial
fluid, blood, wound exudate as well as exhaled breath. The flexible platforms, sensing mechanisms, and
device and system configurations employed for continuous monitoring are summarized. We also discuss
the key challenges and opportunities of the wearable and flexible chemical sensors that lie ahead.
1. Introduction
The increasing research interest in personalized medicine – an
innovative approach harnessing biomedical devices to deliver
tailored diagnostics and therapeutics according to the indivi-
dual characteristics of each patient – promises to revolutionize
traditional medical practices.
1,2
This presents a tremendous
opportunity for developing wearable devices toward predictive
analytics and treatment. On the other hand, the Internet of
Things (IoT) – sensors and actuators connected by networks –
has received enormous attention in the past decade.
3,4
The IoT
is expected to revolutionize future medicine by enabling highly
personalized and accessible healthcare and will have an eco-
nomic impact on the healthcare of over 1 trillion dollars in 2020.
3
As healthcare cost and the world’s aging population increase,
there has been a need for personalized wearable devices to
continuously monitor the health status of patients while patients
Division of Engineering and Applied Science, California Institute of Technology,
1200 E California Blvd, Pasadena, CA 91125, USA. E-mail: weigao@caltech.edu
Yiran Yang
Yiran Yang received her BS degree
in Bioengineering (magna cum
laude) from Rice University. She
then joined Dr Wei Gao’s
research group and is currently
pursuing her PhD degree in
Medical Engineering at Caltech.
Her research interests include
wearable electronics, biosensors
and nanomedicine. She is also a
member of the Tau Beta Pi
Engineering Honor Society.
Wei Gao
Wei Gao is currently an Assistant
Professor of Medical Engineering
at the California Institute of
Technology. He received his PhD
in chemical engineering from the
University of California, San
Diego in 2014 (with Professor
Joseph Wang) as a Jacobs Fellow
and HHMI International Student
Research Fellow. He worked as a
postdoctoral fellow in Electrical
Engineering and Computer
Sciences at the University of
California, Berkeley (with Professor
Ali Javey) during 2014–2017. His current research interests include
flexible electronics, wearable biosensors, micro/nanorobotics, and
nanomedicine.
Received 7th January 2018
DOI: 10.1039/c7cs00730b
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are out of hospital. In this case, wearable biosensors can sample
physiological signals conveniently and noninvasively, and thus
provide sufficient information for health monitoring and even
preliminary medical diagnosis.
4–6
Owing to their unique properties such as light weight, low
cost, high flexibility, excellent stretchability, and great conform-
ability, flexible electronics could serve as an ideal platform for
personalized wearable devices.
7–17
Recent advances in the fabri-
cation and development of flexible electronic devices coupled
with novel micro/nanostructured materials have shown tremen-
dous potential in a number of practical applications, including
physiological monitoring, intelligent robotics, smart displays,
and energy harvesting and storage.
18–28
In particular, wearable
and flexible sensors have been demonstrated for tracking con-
ventional physical signals in the past decade. These wearable
devices are able to perform continuous physiological monitoring
via real-time measurements of body motion, blood pressure,
body and skin temperature, heart rate as well as electrophysio-
logical activities including EEG (electroencephalography), ECG
(electrocardiography) and EMG (electromyography).
29–42
Several
key review articles on wearable and flexible electronic devices
have been published in the past years with emphases on
materials selection, sensor fabrication, device platforms, or their
physical health monitoring applications.
43–48
Despite the rapid growth in wearable and flexible sensing
technologies, the commercially available wearable devices at
present fail to provide more insightful information on users’
health state at the molecular level. Biomarkers from the human
body are defined as ‘‘molecules that can be objectively measured
and evaluated as indicators of normal or disease processes and
pharmacological responses to therapeutics’’ and can provide a
dynamic and powerful access to understand a broad spectrum of
health conditions, and will aid in the prediction, screening,
diagnosis and therapy of diseases.
49,50
Biomarker monitoring
has been used by generations of epidemiologists, physicians,
engineers and scientists to study human diseases. However,
many studies using biomarkers never achieve their full potential
due to the lack of continuous monitoring technologies that
could identify the roles of biomarkers in a timely manner. The
instantaneous and continuous detection of relevant biomarkers
in the human body using wearable chemical sensors can provide
more in-depth personal healthcare monitoring and medical
diagnosis, as compared to the detection of physical activities
and vital signs. The limited availability of wearable chemical
sensors has hindered further progress towards continuous
personalized health monitoring. Up to now, there have been
few single-platform, cost-effective, portable and fully-functional
systems available in the market due to their respective inherent
limitations.
2. Wearable and flexible chemical
sensors
According to the International Union of Pure and Applied
Chemistry (IUPAC), a chemical sensor is ‘‘a device that transforms
chemical information, ranging from the concentration of a
specific sample component to total composition analysis, into an
analytically useful signal’’.
51
A typical chemical sensor usually
consists of a recognition element (receptor) and a physicochemical
transducer. The function of the receptor is to provide high
selectivity towards the target analyte in the presence of poten-
tially interfering chemicals while the transducer is the key com-
ponent that converts the chemical information to a measurable
analytical signal.
Wearable and flexible chemical sensors could be used as
attractive alternatives to the bulky and expensive analytical
instruments used in the healthcare sector. In traditional clin-
ical settings, urine and blood samples are routinely analyzed
through standard analytical techniques, which are expensive,
time-consuming and unable to provide continuous measure-
ments of the concentration of an analyte of interest. In addi-
tion, although the current gold-standard fluid for diagnostics is
blood, it requires invasive sampling that poses a major hurdle
and is unsuitable for long-term continuous use. Recognizing
the significance of non-invasive wearable and flexible sensors
for continuous molecular monitoring, researchers have focused
tremendous effort on wearable and flexible techniques that can
sample and analyze the major electrolytes, metabolites, heavy
metals, and toxic gases directly in alternative body fluids, such
as sweat, tears, interstitial fluid and saliva, as well as exhaustion
breath, particularly in the past 5 years; researchers have also
developed wearable epidermal sensors that can monitor wound
healing (Fig. 1). The transition from blood to other body fluids
and breath provides a noninvasive means of in situ sensing,
which is more attractive toward continuous health monitoring
in daily life. Based on the transduction technique, the wearable
and flexible chemical sensors are mainly either electrochemical
or optical in nature.
At present, electrochemical biosensing is the most common
wearable and flexible sensing strategy owing to its unique advan-
tages of sensor miniaturization, high sensitivity and label-free
Fig. 1 Wearable and flexible chemical sensors for non-invasive health
monitoring.
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direct measurement.
52–56
Some classic examples of electro-
chemical sensors have been developed and demonstrated on
wearable or flexible platforms over the years: (1) amperometric
sensors. Amperometric sensors measure current generated
from the oxidation or reduction of an electroactive analyte in a
chemical reaction. Enzymatic amperometric sensors have been
used for continuous monitoring of glucose, lactate, ethanol and
uric acid, where chemical reactions of the target metabolite
catalyzed by a specific enzyme (e.g. glucose oxidase, lactate
oxidase and urate oxidase) generate electrical current propor-
tional to the target concentration. (2) Potentiometric sensors.
Potentiometry is usually defined as a zero-current technique that
measures the potential appearing between the working electrode
and the reference electrode in an electrochemical cell.
58
The ion
selective electrode based potentiometric sensors have been
widely used for selective ion quantification.
59
They usually
contain a permselective membrane where the target ions interact
with the corresponding ionophore (e.g. a sodium ionophore X for
Na
+
, valinomycin for K
+
and a calcium ionophore ETH 129 for
Ca
2+
) in the sensing membrane and cause voltage changes. For
potentiometric sensors, the relationship between the analyte
concentration and voltage output can be described by the Nernst
equation: E=RT/zFln(A/A
1
) = const + Slog(A), where Aand A
1
are
ion activities/concentrations outside and inside the membrane,
zis the charge of the ion, and Sis the sensor sensitivity. For
sodium or potassium sensing at room temperature, theoretically
an ion selective sensor should have a sensitivity of 59.16 mV per
decade of concentration. (3) Voltammetric sensors. Voltammetry
is a very versatile and well-explored electroanalytical method and
a transduction principle for deriving information about one or
multiple analytes dynamically by measuring the current as a
function of the varied potential. For example, stripping voltam-
metry is suitable for monitoring the heavy metals in body
fluids.
58
(4) Electrochemical biosensors including affinity-based
immunosensors and DNA sensors which have widely been used
for analyzing proteins, peptides and DNAs/RNAs.
Optical sensing represents another attractive way for develop-
ing integrated chemical sensors owing to its low cost and
simplicity.
60
Optical chemical sensors employ optical transduc-
tion techniques (e.g. colorimetric, fluorescence, and lumines-
cence) to retrieve analyte information. The most widely used
optical sensing technique in wearable and flexible electronics is
colorimetry which refers to sensing elements undergoing a
simple color change in the presence of a target analyte through
a chemical reaction. Such a color change is typically quantified
through a simple absorbance measurement where the sensing
element is illuminated and the reflected or transmitted light is
recorded. Specificity and sensitivity are typically dictated by the
modified substrates or colorimetric chemical reactions. This
technique has been explored for detecting metabolites (e.g.
glucose and lactic acid) and electrolytes (e.g. H
+
,Na
+
and Cl
)
in body fluids.
57
When used for in situ analysis in physiological samples,
wearable chemical sensors are subjected to many environ-
mental effects that can affect their stability, reproducibility,
and sensitivity. The challenges of on-body chemical analysis of
body fluids and the early efforts of developing wearable
chemical sensors for monitoring sweat have been summarized
by Coyle et al. in previous reviews and book chapters.
52,61–63
In
another review by Bandodkar and Wang, technological gaps
impeding the successful realization of effective wearable chemical
sensor systems were discussed, including material selection,
power source and consumption, analytical procedure, wireless
communication, data acquisition and processing, and protec-
tion of private information.
54
The emerging nanotechnology,
materials science and flexible electronics have recently led to a
remarkably rapid development in the field of wearable chemical
and biochemical sensors, particularly in the past few years.
52–57
This review focuses on the very recent advances of wearable
and flexible electronics for non-invasive health monitoring
through the in situ analysis of body fluids. In the following
sections, we will mainly review separately the main research
activities of continuous and non-invasive biomarker monitoring
of human sweat, tears, saliva, blood, interstitial fluids, wound
fluids as well as exhaled breath reported over the past few years.
We will cover wearable and flexible substrates, conducting
electrodes, sensing materials, device configurations and working
principles employed for continuous biomarker monitoring.
Finally, we will also discuss the overall challenges, opportunities
and commercialization perspectives of wearable and flexible
chemical sensors that lie ahead toward personalized health
monitoring.
3. Wearable and flexible electronics
for continuous molecular analysis
3.1. Sweat analysis
Sweat is a very important body fluid that contains rich infor-
mation about our physiological state. The wide distribution of
sweat glands in the human body and the abundant biochemical
compounds in sweat have made sweat a feasible and ideal
biofluid for non-invasive biosensing.
64–66
Eccrine sweat, which
is easily accessible non-invasively, is excreted directly onto the
surface of the skin and is composed of water and various electro-
lytes (e.g. sodium, potassium, calcium, and chloride), nitrogenous
compounds (e.g. urea and amino acids), and metabolites such as
glucose, lactate and uric acid. Xenobiotics such as drugs and
ethanol can also be found in sweat.
67
Abnormal health conditions
(e.g. electrolyte imbalance and physical stress) and diseases can
alter sweat composition by either varying the concentration of
common components or leading to the presence of new compo-
nents. Sweat ethanol concentration is highly correlated with blood
ethanol concentration; elevated sweat urea concentration is linked
to kidney failure.
68–72
Abnormally high sweat chloride concen-
tration is observed in cystic fibrosis (CF) patients and sweat
chloride analysis has been adopted as the gold standard for cystic
fibrosis diagnosis.
67,73,74
Despite many advantages, sweat analysis
remains an underrepresented solution for health monitoring and
clinical diagnosis compared to blood and urine analysis due to the
challenges of contamination, evaporation as well as the lack of real
time sweat sampling and sensing devices.
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While sweat analysis has attracted tremendous attention in
recent years, novel wearable and flexible sweat sensing platforms
based on different detection approaches have been developed for
in situ sweat analysis toward continuous health monitoring
(Table 1 and Fig. 2). Specifically, sweat analytes are detected
and quantified either with electrochemical and/or with optical
methods on various flexible substrates such as textiles,
81,86,90–92,98
tattoo,
75–80
and plastic.
55,82,83,85,88,94,95,99,100
The electrochemical
sweat sensors use functionalized electrodes to transduce sweat
analyte concentration into electrical signals (i.e. current or voltage)
which are transmitted to the processing component and give
quantitative results with high sensitivity. The earliest-developed
approach was based on a textile platform, which was used for
sensing sweat pH, Na
+
levels, and sweat rate.
90–92
However, for
most of the textile platforms tested in situ, the sensing results
were severely affected by suboptimal contact between the sensor
probe and human skin. In addition to textile-based platforms,
a screen-printed tattoo-based sweat sensor was developed by
Wang’s group.
75–80
The tattoo platform was implemented
through screen-printing technology on substrates such as
silicone materials (e.g. PDMS, Ecoflex, Solaris) or polymer
materials (e.g. PVA, PET, PEN). These materials have elastic
properties similar to human skin and have enhanced the
contact, adhesion and transpiration of the platform. Layers of
electrode materials (such as carbon, Ag/AgCl) and insulators are
printed on a tattoo paper sheet with designed patterns and the
electrodes are then modified with sensing membranes. The
temporary-transfer tattoo sweat biosensor is highly compact
and flexible (Fig. 2a) and enables reliable continuous and in situ
measurement of lactate, pH, Zn, ammonium and alcohol in
sweat.
75–79
An adhesive radio-frequency identification (RFID)
sensor bandage was reported which can chronologically
monitor sweat sodium using an ion selective electrode based
sodium sensor (Fig. 2b).
81
A commercial RIFD chip was adapted
Table 1 Wearable and flexible chemical sensors for continuous sweat analysis
Analyte Analytical technique Recognition element Materials and platform Ref.
Lactate Electrochemical-amperometry Lactate oxidase (LOx) Temporary transfer tattoo 75
pH Electrochemical-potentiometry Polyaniline (PANI) Temporary transfer tattoo 76
Zinc Electrochemical-stripping voltammetry Bismuth electrode Temporary transfer tattoo 77
Ammonium Electrochemical-potentiometry Nonactin ionophore based ISE Temporary transfer tattoo 78
Alcohol Electrochemical-amperometry Alcohol-oxidase Temporary transfer tattoo 79
Sodium Electrochemical-potentiometry Sodium ionophore based ISE Temporary transfer tattoo 80
Sodium Electrochemical-potentiometry Sodium ionophore based ISE Textile-based bandage patch 81
Sodium, potassium,
lactate, glucose
Electrochemical-potentiometry and
amperometry
Sodium and potassium ISEs; glucose
oxidase (GOx); LOx
PET wristband/handband 55
Ca, pH Electrochemical-potentiometry Calcium ISE; PANI PET wristband/handband 82
Zn, Cd, Pb, Cu, Hg Electrochemical-stripping voltammetry Bismuth and gold electrodes PET film 83
Humidity, glucose, pH Resistive for humidity; electrochemical
amperometry and potentiometry
PEDOT for humidity; GOx;
polyaniline
Silicone wristband 56
Humidity, glucose, pH Resistive for humidity; amperometry;
potentiometry
PEDOT for humidity; GOx; PANI Disposable silicone strip 84
pH Electrochemical-transistor ISFET for pH PET patch 85
pH, metal ions Optical-colorimetry pH sensitive colorimetric indicators;
Cu-sensitive and Fe-sensitive colori-
metric indicators
Cellulose paper and silicone
patch
86
pH Optical-colorimetry pH sensitive ionogels/dyes Adhesive plaster/wristband 87
pH, lactate, chloride,
glucose
Optical-colorimetry LDH and diaphorase in formazan
dyes; GOx with iodide; pH indicator
dye; Hg
2+
and Fe
2+
with TPTZ
PDMS patch 57
Glucose, sodium,
chloride
Electrochemical-potentiometry and
amperometry
Sodium ISEs, Ag/AgCl electrode for
chloride; glucose oxidase
PET wristband 88
Lactate, glucose Electrochemical-amperometry GOx; LOx PDMS on a skin adhesive 89
pH Optical-colorimetry pH sensitive dye Textile patch on a waistband 90
Sodium Electrochemical-potentiometry Sodium ionophore based ISE Fabric patch on a belt 91
Sodium, pH Electrochemical-impedance; optical-
colorimetry
PPy based ISE; pH sensitive dyes Fabric patch on a belt 92
EtG Electrochemical-impedance EtG antibody modified on ZnO or Au
electrode
On either glass or polyimide 93
Chloride Electrochemical-potentiometry Ag/AgCl electrodes with salt bridge PDMS on wristband or
adhesive bandage
94
Sodium, chloride Electrochemical-potentiometry Sodium ISE, Ag/AgCl electrode PET film attached to a
memory foam and strap
95
Glucose, cortisol Electrochemical-amperometry and
impedance
Glucose oxidase on ZnO film; cortisol
antibody on ZnO film
Polyamide film 96
Sodium Electrochemical-potentiometry Gold-nanodendrite ISE Silicon chip on a flexible
headband
97
Adrenaline, NaCl Electrochemical-transistor Pt for adrenaline, Ag gate for NaCl Cotton wire 98
Lactate Electrochemical-amperometry Lactate oxidase Polycarbonate membrane 99
Glucose, lactate, uric
acid, urea
Piezoelectrical LOx, GOx, uricase, and urease
immobilized on ZnO nanowires
Kapton film 100
Potassium, lactate Electrochemical-amperometry and
potentiometry
LOx; potassium ISE PET sticker on eyeglasses 101
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to allow potentiometric sensing, temperature sensing and
wireless communication with the smart phone.
Considering the complex correlation among all biophysical
and biochemical information, simultaneous, accurate and multi-
plexed detection of the signatures with the capability of perform-
ing on-site data processing and communication is in urgent
need. To this end, a fully integrated system was subsequently
developed and widely applied in sweat sensing. Gao et al. merged
plastic-based sensors that interface with the skin with integrated
circuits consolidated on a flexible printed circuit board onto a
flexible PET substrate, and achieved stable and accurate in situ
monitoring of multiple sweat analytes including glucose, lactate,
K
+
and Na
+
as well as skin temperature for signal calibration
(Fig. 2c).
55
The integrated system takes the form of a smart
wristband or headband and has stable contact with the skin;
signal conditioning, processing and bluetooth based wireless
transmission were consolidated in the flexible printed circuit
board (FPCB) on the same wristband. The same group also
developed a flexible microsensor array for heavy metal moni-
toring in sweat (Fig. 2d).
83
The ultra-low level multiplexed sweat
heavy metal analysis was realized using gold and bismuth
microelectrodes through anodic stripping voltammetry. The
effective preconcentration step and selective oxidation of metal
phase species make anodic stripping analysis the most effective
technique for monitoring the ultralow levels of heavy metals in
body fluids. The proposed wearable microsensor array was able
to simultaneously detect Zn, Cd, Pb, Cu and Hg in human sweat
during physical exercise. A graphene-based electrochemical
sensor was developed by Lee et al. and applied for sweat glucose
monitoring on a stretchable patch (Fig. 2e) and a disposable
strip (Fig. 2f).
56,84
The device shows improved sensitivity and
electrochemical activity due to the use of graphene-hybrid
materials. The serpentine bilayer of gold mesh and gold-doped
graphene formed an efficient electrochemical interface for the
stable transfer of electrical signals. The large electrochemically
active surface area of the graphene-hybrid electrode enabled the
real time detection of low levels of sweat glucose (correlation
factor between sweat and blood glucose in this work: B0.017).
The reading from the glucose sensors was calibrated with real-
time pH and temperature information obtained from pH and
temperature sensors incorporated on the same platform to
eliminate the influence of pH and temperature on the enzyme
activity and on the glucose sensor response.
In addition to electrochemical sensors, wearable sensors based
on optical methods were also developed on flexible platforms
toward sweat analysis in recent years. For optical sweat sensing,
Fig. 2 Wearable and flexible sweat biosensors. (a) Epidermal temporary-transfer tattoo sweat biosensor for in situ continuous monitoring of lactate, pH,
etc. Reproduced with permission from ref. 75. Copyright 2013 American Chemical Society. (b) Wireless sensor bandage for chronological monitoring of
sweat Na
+
concentration. A RFID chip was incorporated for wireless communication with a smart phone. Reproduced with permission from ref. 81.
Copyright 2015 IEEE. (c) Fully integrated ‘‘smart wristband’’ with a flexible sensor array and a flexible PCB for multiplexed in situ sweat analysis.
Reproduced with permission from ref. 55. Copyright 2016 Nature Publishing Group. (d) Flexible sensor array for multiplexed heavy metal monitoring in
sweat. Reproduced with permission from ref. 83. Copyright 2016 American Chemical Society. (e) Graphene-based flexible and stretchable diabetes patch
for sweat glucose monitoring. Reproduced with permission from ref. 56. Copyright 2016 Nature Publishing Group. (f) A disposable strip for sweat glucose
sensing. Reproduced with permission from ref. 84. Copyright 2017 American Association for the Advancement of Science. (g) Stretchable sensor with
serpentine structure on porous polyurethane for sweat pH sensing. Reproduced with permission from ref. 86. Copyright 2014 Wiley. (h) Microfluidic-
based device for sweat routing and colorimetric pH sensing. Reproduced with permission from ref. 87. Copyright 2012 Elsevier. (i) Stretchable
microfluidic based sweat patch for spontaneous sweat routing through serpentine channels and the reservoir, with multiplexed colorimetric sensing.
Reproduced with permission from ref. 57. Copyright 2016 American Association for the Advancement of Science.
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an optical fiber or digital camera was used to capture the colori-
metric or absorbance response to color-responsive materials
(i.e. pH dyes) of sensors, and quantitative results are generated
by image analysis on a personal computer or a mobile phone.
The optical detection method displays results to the wearer in
real time and allows for simple interpretation of the analyte
concentration. A stretchable sensor with a serpentine structure
was fabricated on porous polyurethane and was used for pH
measurements (Fig. 2g).
86
Another colorimetric sensor was
developed on a microfluidic platform using a common reservoir
to continuously draw sweat from the sensing area to four
independent reservoirs housing ionogel pH dyes (Fig. 2h).
87
In addition to pH sensing, Koh et al. developed a stretchable
microfluidic device that harvests and stores sweat from human
skin, routes sweat to different channels with colorimetric assay
reagents, and measures sweat rates as well as the levels of
multiple sweat analytes including chloride, glucose, and lactate
and pH (Fig. 2i).
57
The microfluidic network was designed to
route sweat spontaneously through the network of serpentine
channels and reservoirs. The reactions of the sensors with
sweat analytes induce changes in the color of the chromogenic
reagents. The integrated near field communication system
between a sweat monitoring device and a smartphone enables
image capture and analysis. The cellphone application can
extract RGB color information and allow the user to read the
sweat analyte concentrations.
The development of all the above-mentioned sweat sensors
has opened the door of practical applications toward real-time
monitoring of fitness and health conditions. The wearable and
flexible sweat sensors have been tested intensively in human
trials to extract more insightful information from sweat. For
example, fully integrated wearable sensor arrays developed by
Gao et al. have been used for effective and non-invasive
identification of dehydration during long-term outdoor exercise
(Fig. 3a).
55
Sweat Na
+
levels increased substantially when the
subjects had lost a large amount of water (B2.5% of the body
weight), indicating that sweat sodium can potentially serve as
an important biomarker for monitoring dehydration. The wear-
able diabetes patch developed by Lee et al. can monitor the
sweat glucose level electrochemically in real-time (sensing unit)
and actuate transcutaneous diabetic drug release (therapeutic
unit). Daily glucose monitoring in vivo was performed and
changes in the sweat glucose concentration were well correlated
with those of the blood glucose concentration (Fig. 3b).
56
Bioresorbable polymer-based micro-needles coated with the
phase-change material were coupled with multichannel thermal
actuators in the therapeutic unit such that the drug was released
in a stepwise manner when the programmed temperature
Fig. 3 Physiological and clinical investigations through human trials using wearable and flexible sweat biosensors. (a) Fully integrated headband and
wristband for dehydration status monitoring during long-term exercise. Reproduced with permission from ref. 55. Copyright 2016 Nature Publishing
Group. (b) Wearable diabetes patch for real-time monitoring of the daily sweat glucose level. Reproduced with permission from ref. 56. Copyright 2016
Nature Publishing Group. (c) Wearable optical sensor patch for real-time sweat rate monitoring of different subjects during a long-distance bicycling
race. Reproduced with permission from ref. 57. Copyright 2016 American Association for the Advancement of Science.
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exceeded the threshold as a response of glucose levels. The
stretchable optical sweat sensor developed by Koh et al. was
tested on human subjects in a competitive long-distance out-
door bicycling race (Fig. 3c).
57
The data showed that older
subjects (aged 50 to 69 years) had greater sweat rates compared
to younger subjects (aged 10 to 29 years), and male subjects
exhibited larger sweat rates than females.
The in situ sweat sensor tests described above were mostly
carried out with subjects performing vigorous exercise such as
running or cycling to generate sweat. Although sweat produc-
tion by exercising is viable for sensor testing, it might not be the
ultimate solution for continuous sweat monitoring, particularly
for sedentary individuals. In such cases, alternative methods of
sweat induction are desired and need to be introduced. One
well-developed sweat induction method is iontophoresis, which
is widely used to induce sweat excretion locally as selected,
usually on wrist. A charged substance, usually pilocarpine, is
applied to the skin at one end of the test site. With the applied
electrical potential difference between both ends of the test
site, a mild electric current carries the pilocarpine into the skin
and sweat glands are stimulated (Fig. 4a).
88,102
Based on this
sweat induction approach, Kim et al. combined pilocarpine
iontophoresis with in situ sweat alcohol sensing (Fig. 4b).
79
By applying constant current through screen-printed ‘‘tattoo’’
AgCl iontophoresis electrodes, sweat was inducted and imme-
diately used for sweat ethanol measurement by sensing electro-
des; the sensing results can be transmitted wirelessly for data
readout. The screen-printed ‘‘tattoo’’-based ethanol sensing
electrode relied on ethanol oxidase with the Prussian blue
(PB) electrode transducer. With this sweat inducing and sens-
ing device, on-body ethanol amperometric sensing yields a
significant increase in current response after alcohol consump-
tion (Fig. 4b), reflecting the good correlation between sweat and
blood ethanol levels. Similarly, a fully integrated wearable
sweat extraction and sensing platform was developed which
contains a miniaturized iontophoresis module capable of indu-
cing sweat with different excretion rate profiles and at periodic
intervals (Fig. 4c).
88
The secretion rate pattern was controlled
by controlling the compound formulation (e.g. acetylcholine,
methacholine, and pilocarpine) loaded into iontophoresis
hydrogels. The placement of the sensing electrodes between
iontophoresis electrodes enables on-site analysis of induced
sweat. The electrochemical electrodes were used to measure
glucose, Na
+
, and Cl
with high sensitivity in sweat. With this
device, diagnosis and disease investigation could be made. In
this study, real-time Cl
and Na
+
levels in sweat are much
higher in CF patients than in healthy individuals (Fig. 4c). The
device was also used to investigate the correlation between
sweat and blood glucose levels. This fully autonomous sweat
induction and analysis device leads to more potential sweat
sensing applications such as point-of-care diagnosis and health
monitoring.
The fast progress in sweat sensing platforms has expedited
research on sweat sampling, as it is critical for improving the
temporal resolution toward real-time sweat analysis and for
minimizing the evaporation and contamination of the sweat
samples. Peng et al. developed a new oil/membrane approach
with the use of carbachol for sweat stimulation that could reduce
the sample volume from the mL to nL regime and minimize
analyte contamination (Fig. 5a).
103
By using a cosmetic-grade oil
and a micro-porous membrane between skin and sensors, the
sampling intervals are on the order of minutes, and the hydro-
philic contaminants from the skin surface are blocked. Sonner
et al. optimized the sensor geometry to improve the conduction
of iontophoresis sweat flow.
94
Carbachol has shown prolonged
sweat stimulation in directly stimulated regions for five hours or
longer. Choi et al. developed a skin-like sweat sensor that collects
and stores sweat in a set of interconnected micro-reservoirs
(Fig. 5b).
104
The sweat sample was guided passively by the sweat
gland induced pressure to a micro-channel network that incorpo-
rates capillary bursting valves in a sequential fashion and allows
for chrono-monitoring. A soft wearable microfluidic system for
measuring secretory fluidic pressures generated by eccrine sweat
glandswasalsodemonstratedbythesamegroup.
105
The combi-
nation of iontophoresis and chrono-sampling microfluidics could
possibly bring time-dependent sweat analysis in future applica-
tions. An epidermal microfluidic system was developed by Martı
´
n
et al. for enhanced sweat sampling and metabolite detection
(Fig. 5c).
89
Theoretical modeling of the microfluidic device design
Fig. 4 Iontophoresis based sweat induction and sensing toward clinical
applications. (a) Mechanism of iontophoresis in sweat sensing applications.
Reproduced with permission from ref. 88. Copyright 2017 National Academy
of Sciences USA. (b) Tattoo-based iontophoretic-biosensing system for sweat
ethanol monitoring. A significantly increased current response of the
wearable ethanol sensor was observed after alcohol intake. Reproduced
with permission from ref. 79. Copyright 2016 American Chemical Society.
(c) A fully integrated platform for controlled iontophoresis sweat induction
and real-time sweat sensing. Real-time sensing results of sweat Na
+
and
Cl
levels distinguished cystic fibrosis patients from healthy subjects.
Reproduced with permission from ref. 88. Copyright 2017 National Academy
of Sciences USA.
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allowed optimization of the sweat sampling process. This system
enabled efficient sweat pumping to the electrochemical detection
chamber containing the enzymatic glucose and lactate sensing
electrodes.
Recent advances in sweat sensing application have proved
sweat to be an excellent candidate for non-invasive continuous
health monitoring. As more biomarkers are identified in sweat,
it can be expected that sweat sensing would be expanded to
more practical medical applications. Despite the non-invasive
nature of sweat sensing, there are still major challenges to be
addressed for continuous on-body sweat monitoring. Sweat can
be generated in different ways (e.g. heat, exercise and ionto-
phoresis) at different rates, and the levels of sweat biomarkers
(e.g. glucose) could be much lower than those in blood. Sweat
sensing at low sweat rates (i.e. in sedentary individuals) will
require miniaturized sensors with higher sensitivity. Small
volumes of sampled sweat would be susceptible to evaporation
and contamination issues, which would lower the accuracy of
the sensing results. To resolve these problems, in situ sampling of
sweat should be carefully designed and fast detection is needed.
Mixing of the newly sampled sweat and measured sweat samples
can affect the real-time sensing results and lower the chrono-
accuracy. Controlled sweat flow is necessary to improve the tem-
poral resolution toward accurate real-time measurements. For
some sweat analytes such as NaCl, their concentrations in sweat
are sweat-rate dependent, so the measured results for these ana-
lytes should be calibrated with sweat rates. To expand the broad
use of sweat-based medical or fitness monitoring, it is very critical
to identify relevant analytes present in sweat and understand the
physiological pathway of analyte secretion into sweat. Previous
studies have shown that the excretion of sweat biomarkers can
be different from each other (passive, active, or self-generating).
68
In particular, though some small molecules (e.g. Na, Cl, K, and
Ca ions) could be theorized to passively or actively partition from
blood, plasma or serum, up until now the excretion mechanisms of
many sweat biomarkers (e.g. hormones, peptides or proteins)
Fig. 5 New approaches for enhanced sweat sampling and sensing. (a) Oil-membrane approach with carbachol for sweat induction. A cosmetic-grade
oil and a micro-porous membrane were used between skin and sensors. Reproduced with permission from ref. 103. Copyright 2016 Royal Society of
Chemistry. (b) Micro-channel network based sweat sensor that collects and stores sweat in micro-reservoirs, allowing for chrono-monitoring of sweat
analytes. Reproduced with permission from ref. 104. Copyright 2017 Wiley. (c) Epidermal three-layered microfluidic device with enhanced sweat
sampling efficiency and transmission of sweat samples to sensing electrodes. Reproduced with permission from ref. 89. Copyright 2016 American
Chemical Society.
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have not been well understood as partitioning of these larger
molecules may result from complex or varied pathways.
3.2. Tear analysis
The human tear is a biofluid accumulated in the eyes and it
contains a number of salts, proteins, enzymes, and lipids. The
chemical composition of tears can reveal useful information on
ocular conditions and systemic disorders.
106,107
For example,
the increased level of proline-rich protein 4 was identified as
a biomarker for dry-eye conditions; modified expression of
cancer-related biomarkers such as complement proteins was
found in breast cancer patients; tear glucose concentration is
highly correlated with blood glucose concentration and human
tears have been used for continuous diabetes management.
108–110
Earlier wearable and flexible tear glucose sensors were
developed as flexible strips.
111–113
In the work of Kudo et al.
(Fig. 6a), a flexible and stretchable glucose sensor was prepared
on a flexible PDMS substrate and then coated by a hydrogen
peroxide permeable poly(MPC-co-DMA) membrane.
111
The
glucose sensor measures tear glucose concentration ampero-
metrically with the use of glucose oxidase (GOx). In a later work
from the same group, the tear glucose concentration was measured
with GOx immobilized on a flexible oxygen electrode, and a gas
permeable membrane was used. The device was tested in vivo
on an anesthetized rabbit and the output current was moni-
tored before and after oral glucose administration. The change
in tear glucose was detected with 10–20 min delay compared to
measured blood glucose change.
112
The strip-shaped sensor was
difficult to fix on pupil due to insufficient contact area and flimsy
movement, and Chu et al. improved this flexible strip structure by
attaching the strip sensor to a contact lens (Fig. 6b).
113
In this
case, the soft PDMS contact lens was molded; the flexible
Fig. 6 Wearable and flexible sensors for continuous tear analysis. (a) A flexible PDMS-based tear glucose sensing strip. Reproduced with permission
from ref. 111. Copyright 2006 Elsevier. (b) Soft PDMS contact lens with a glucose sensing strip attached. Reproduced with permission from ref. 113.
Copyright 2011 Elsevier. (c) Soft contact lens integrated with an amperometric glucose sensor. Reproduced with permission from ref. 114. Copyright 2011
Elsevier. (d) Glucose sensing contact lens, with an integrated antenna for wireless transmission. Reproduced with permission from ref. 115. Copyright
2012 IEEE. (e) Glucose and intraocular pressure sensing contact lens with graphene-hybrid AgNWs to improve the transparency of the lens. Increased
reflection after wearing the highly transparent contact lens possibly due to glucose binding in tear fluid. Reproduced with permission from ref. 116.
Copyright 2017 Nature Publishing Group.
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electrode was bonded to the peripheral surface and GOx was
immobilized onto the sensor. The device was tested in vivo on a
rabbit without using anesthetics. A change in the tear glucose
level was observed after oral glucose administration; continuous
tear glucose monitoring was performed and the sensor output
current remained stable (Fig. 6b). The calculated glucose concen-
tration from the measured current was consistent with human
tear glucose concentration. Yao et al. furthered this combination
of glucose sensor and contact lens by constructing a contact lens
with an amperometric glucose sensor.
114
The glucose sensor was
created on a polymer substrate and then the substrate was
shaped into a contact lens. GOx was immobilized in a titania
sol–gel layer that showed improved sensitivity as the titania
sol–gel film is very efficient for retaining the GOD activity and
preventing the enzyme detachment. A Nafion permselective film
was used to reduce potential interference from ascorbic acid,
lactate, and urea present in tears. The contact lens was tested in
glucose solutions and yielded sensitive measurements with good
interference rejection. The same group further improved the
functionality of contact lens glucose sensors and enabled con-
tinuous and real-time tear monitoring with wireless transmis-
sion (Fig. 6c).
114
A loop antenna, a wireless communication
interface chip and an electrochemical glucose sensor (Fig. 6d)
were integrated on the polymer substrate formed into a contact
lens.
115
The gold antenna loop was directly patterned on the
polymer substrate and was used for RF powering from a distance
of 15 cm. The integrated circuit (IC) was based on a 0.36 mm
2
CMOS chip with no external components. The device was tested
in buffer solutions on a PDMS eye model with an artificial tear
duct and tear drain, and was able to yield sensitive results
continuously in real time. This type of contact lens based glucose
sensor has low power consumption (3 mW) and can transmit
measurement data with a wireless readout. This platform was
later further developed by Google. More recently, Kim et al.
developed soft contact lenses for wireless detection of glucose
and intraocular pressure (Fig. 6e).
116
Using a graphene (2D
nanomaterial) hybrid with Ag nanowires (AgNWs) (1D nano-
material), the conductivity, transparency and stretchability of
the contact lens were enhanced, and the hybrid can serve as a
source/drain electrode of a field-effect transistor (FET) with
graphene as a channel. For glucose sensing, GOx was immobi-
lized on the graphene channel with a pyrene linker; oxidation of
glucose and reduction of water to hydrogen peroxide change
the concentration of charge carriers in the channels and thus
changes the drain current. SU8 was used as an additional
diffusive barrier to protect AgNWs from external damage from
molecules in tear fluid, as well as to prevent AgCl formation
that could be harmful for human eyes. The FET glucose sensor
achieved 10 times improvement in sensitivity compared to
previous evaporated metal electrodes. The bottom spiral struc-
ture is used for glucose sensing while the top spiral structure
enables wireless transmission and powering. This wireless
sensor was integrated onto the eyes of a live rabbit and showed
a higher reflection than the value before wearing due to glucose
binding in tear fluid of the rabbit (Fig. 6e). This device fully
explored the design with 2D nanomaterials and achieved
independent and simultaneous multi-analyte sensing in real
time with wireless transmission.
In addition to wearable and flexible electrochemical sensor
based tear analysis, the optical sensing strategy was also
explored toward tear glucose monitoring. Zhang et al. reported
a soft hydrogen contact lens with assembled fluorescent nano-
particles with a porous structure.
117
This contact lens was tested in
glucose solution. The fluorescence intensity and resonance energy
transfer were measured to obtain the glucose concentration. The
sensor could be used to monitor glucose levels continuously for at
least 5 days. In addition, Badugu et al. developed a contact lens
embedded with boronic acid-containing fluorophores to perform
glucose sensing.
118
Recently, a gelated colloidal crystal array (CCA)
attached contact lens was developed.
119
The CCA was embedded in
a hydrogel matrix and then attached onto a rigid gas permeable
contact lens. The CCA is able to selectively diffract visible light; the
change in glucose concentration between 0 and 50 mM will shift
the diffracted wavelengths between 567 and 468 nm, corres-
ponding to a visible color change from reddish yellow to green
and to blue. The device was tested in vitro andyieldedobservable
color changes in glucose solution and good specificity in simulated
tear fluid, but the sensitivity was severely affected in simulated
tear fluid.
Despite the excellent wearability introduced by the contact
lens platform, tear sensing still faces limitations that need to be
improved. Transparency of the contact lens is usually reduced
by the embedded sensor structure and can obstruct users’
vision in daily life. The material selection and nanostructure
design, and transparent sensor material combined with careful
structure design may be further explored to improve trans-
parency. Heating in the contact lens due to near field wireless
powering could also cause discomfort and even irritation in
users’ eyes, so further research on the selection of antenna
material or energy harvesting methods (e.g. biofuel cell) could
be useful to resolve this issue. Another limitation is the readout
distance of the near field wireless communication. The wireless
readout distance in the works of Liao et al. and Kim et al. was
15 cm and 10 mm, respectively. This means that the users need
to hold the readout device very close to eyes, which might not
be practical for prolonged continuous monitoring. Some other
issues to consider include the need for glucose monitoring
during sleep (when hypoglycemia, a dangerous situation for
diabetics, could occur), potential temporal variations in tear
glucose concentration and calibration for such variations, and
possible pupil damage related to contact lens wear such as
microbial keratitis, contact lens peripheral ulcers and inflam-
matory complications.
120
3.3. Saliva analysis
Human saliva has been considered as an attractive biofluid for
non-invasive diagnostics and monitoring.
121,122
It is a clear and
viscid biofluid secreted into the mouth by salivary glands and
the fluid contains various biomarkers, such as glucose, lactate,
phosphate, enzymes (e.g. alpha-amylase (sAA)), hormones
(e.g. cortisol, steroids), antibodies (e.g. IgA, IgG), etc.
123
Salivary
cortisol and sAA have been identified as crucial biomarkers for
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physical and psychological stress; antibodies in the saliva are
shown to be useful tools for disease diagnosis (e.g. HIV and
intestinal infections); some protein and mRNA biomarkers (e.g.
IL-8 and KRAS) in saliva were identified as cancer biomarkers
and can distinguish cancer patients from healthy individuals.
123–133
Investigation of these salivary biomarkers was facilitated by the
abundant availability of saliva and the ease of saliva sample
collection. Earlier investigation was conducted through sensing
analyte concentrations in ex vivo human saliva samples collected
by spitting or with a collection paper strip.
134,135
Sample collec-
tion is then followed by sensing out of the body, either in lab
tests or in a portable device.
123–133,136,137
As the interest in saliva sensing grew over the past years,
wearable devices for in situ saliva analysis were developed (Fig. 7).
The sample collection and sensing processes were integrated into
a single wearable and flexible platform. Mannoor et al. developed
a salivary bacteria sensing platform on tooth enamel (Fig. 7a).
138
Self-assembled antimicrobial peptides on graphene selectively
bind to bacteria in saliva, and the binding process alters the
electrical resistance that allows for bacterial quantification
(Fig. 7a). The graphene nanosensors were printed onto water-
soluble silk thin-film substrates and then integrated with
electrodes patterned with an inductive coil antenna. The
graphene/electrode/silk hybrid structure is capable of detecting
highly sensitive single-bacteria and a wireless readout could be
achieved with the active device in proximity. In addition to
the form of sensor attachment, the mouthguard was taken as
another form factor for collection–sensing integration. Kim et al.
demonstrated an integration of printable enzymatic electrodes
on a mouthguard.
139
Lactate oxidase was immobilized on the
working electrode, and the 3-electrode system was screen printed
onto the mouthguard. The lactate concentration was measured
amperometrically, and the ex vivo experiments showed promising
results for continuous on-body saliva monitoring. The same
group further explored this electrochemical mouthguard plat-
form toward salivary uric acid (UA) measurements (Fig. 7b).
140
The printed working electrode Prussian blue transducer was
chemically modified by crosslinking uricase enzyme with electro-
polymerizing o-phenylenediamine (Fig. 7b). The sensor was
fabricated on a flexible PET substrate and then integrated with
a wireless amperometic circuitry and mounted onto the mouth-
guard. This mouthguard UA sensor was tested on saliva samples
from hyperuricemia patients and healthy subjects. The detected
salivary UA levels were much higher in hyperuricemia patients
than in healthy subjects. Although real-time in-mouth monitor-
ing was not discussed, the sensor yields sensitive amperometric
measurements of UA and real-time bluetooth wireless data
transmission was achieved. Arkawa et al. also integrated bio-
sensing and real-time wireless communication on a monolithic
mouthguard (Fig. 7c).
141
Glucose oxidase was entrapped with
poly(MPC-co-EHMA) (PMEH) onto the working electrode, and
the 3-electrode system was combined with a wireless trans-
mitter onto a polyethyleneterephthalateglycol (PETG) mouthguard.
The electrode area and PMEH coating on the glucose sensor were
optimized to increase the output current response to glucose
binding. The sensor was tested with a phantom jaw in an open-
loop artificial saliva injection system and the results were
wirelessly transmitted to a personal computer (Fig. 7c). The
testing yields sensitive amperometric outputs over relevant
glucose levels in human saliva.
As described above, wearable saliva sensors have been
developed towards continuous in-mouth monitoring. While saliva
is an attractive candidate for non-invasive health monitoring,
there are still several major roadblocks for practical implementa-
tion of saliva sensing. Considering the continuous in-mouth use
of wearable saliva sensors, the materials, devices and systems
Fig. 7 Wearable and flexible sensors for saliva analysis. (a) Bacteria sensing on tooth enamel with graphene-based nanosensors. Bacteria quantification was
performed based on electrical resistance variation. Reproduced with permission from ref. 138. Copyright 2012 Nature Publishing Group. (b) Mouthguard
device with electrochemical sensors to measure salivary uric acid (sUA) concentration. Enhanced sUA level was detected in hyperuricemia patients.
Reproduced with permission from ref. 140. Copyright 2015 Elsevier. (c) Saliva glucose sensing with the mouthguard sensing platform where GOx was
entrapped in PMEH. Reproduced with permission from ref. 141. Copyright 2016 Elsevier.
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should be fully biocompatible. In the continuous monitoring
scheme, the abundance of secreted proteins and active chemi-
cals generated from food residues can result in significant
interference with target analyte sensing, so the saliva sensor
should be of high specificity in this complex and dynamic
chemical environment. Another challenge comes from contami-
nation and sensor damage. Bacteria accumulate quickly on
surfaces inside the oral cavity due to the favorable humidity
and temperature conditions in the mouth. If the sensor is
unprotected, a bio-film will form on the sensor surface and
severely affect the sensitivity of the device. An antimicrobial or
protective coating on the sensor surface might resolve this
issue, but enhanced sensor sensitivity might be necessary after
the introduction of such protective coatings. Voluntary and invo-
luntary mouth muscle movements in daily life (e.g. talking) may
also impose mechanical stress on sensors, so the sensors should
be mounted in the mouth securely while remaining durable in lieu
of mechanical stress. Considering all these challenges, the future
effort on wearable saliva sensors should focus on careful design of
the form factors that house sensors, ensuring materials’ biocom-
patibility, sensor function and accommodating user’s comfort and
convenience at the same time.
3.4. Non-invasive blood analysis
Blood is a body fluid that transfers necessary substances such
as nutrients and oxygen to cells and transports metabolic waste
products out of cells. It sustains the physical and chemical
equilibrium of cells. A blood test has been the gold standard for
many clinical diagnostic applications.
142–144
In particular,
blood glucose has been heavily investigated and monitored
for diabetic regulation. Various devices were developed to sense
and monitor biomarkers in the blood.
145–149
Since blood is
contained in blood vessels with less accessibility, earlier devices
either performed analysis based on invasive blood draw or
required implantable sensors that are undesired for continuous
health monitoring.
Wearable devices for non-invasive blood monitoring were
recently developed, mostly for oxygenation sensing, pressure
sensing, and heart rate detection. The widely available com-
mercial product for noninvasive molecular blood sensing is the
pulse oximeter, which performs pulse oximetry on finger tips,
an optical test that measures changes in the optical properties
of hemoglobin in its oxygenated and deoxygenated states and
yields both the oxygen saturation level and blood pulse. The
pulse oximetry level has close clinical relevance in the manage-
ment of acute and chronic respiratory disease.
150–152
The earlier
versions of pulse oximeters were of portable size, and more
pulse oximeters were developed to reduce the oximeters into
wearable size. Haahr et al. developed an electronic patch that
can perform reflectance pulse oximetry.
153
The electronic patch
integrated a photodiode (sensor), a battery, and a PCB in an
adhesive patch; the chip on PCB allows for wireless commu-
nication. The microsystem could be attached to different parts
of the skin surface. Since many conventional pulse oximetry
systems are transmission-based detection systems, sensors
were applied to thinner or peripheral parts of the body which
are less opaque, but in the case of shock-induced centralization
and a resulting drop in perfusion, peripheral detection would
not yield reliable results. To resolve this limitation, Venema
et al. developed a reflection-mode oxygen saturation sensor that
could be worn in the ear canal like a headphone.
154
The sensing
was based on photoplethysmography (PPG), which allowed for
reflectance pulse oximetry at opaque body parts. Human sub-
ject testing showed that the device was very comfortable for
long-time wearing, and the tight-fit ear mold could reduce the
motion artifact, and in-ear detection capability provided stable
and more reliable results than peripheral detection. The pulse
oximetry devices described above were rigid and made of
inorganic sensors. Recent advances in flexible and conformal
electronics provide attractive routes for accurately assessing the
blood oximeter. Lochner et al. developed a pulse oximeter
sensor with organic materials on flexible substrates (Fig. 8a).
155
The green and red light-emitting diodes (LEDs) and photodiodes
were both made of organic materials, and the device was low-
cost and disposable. The all-organic optoelectronic sensor accu-
rately measured pulse oxygenation with an error of 2%. Yokota
et al. reported ultraflexible organic photonic skin with multi-
color highly efficient polymer light-emitting diodes (PLEDs) and
organic photodetectors (OPDs) for reflective pulse oximeter
monitoring (Fig. 8b).
156
The optoelectronic skins are extremely
thin (only 3 mm in thickness), lightweight and stretchable; they
can endure bending radii of 100 mm or less and repeatedly
sustain up to 60% compression. The sensor successfully measured
the blood O
2
concentration when laminated on the finger.
Recently, Kim et al. developed a miniaturized flexible and wearable
pulse oximeter (Fig. 8c).
157
Built on a silicone elastomer, the sensor
wasaroundthesizeofafingernailandwasveryflexible.Itwas
equipped with a near-field wireless chip that allows for wireless
power harvesting and data communication. The device can oper-
ate continuously for up to 3 months and can measure oxygen
saturation at almost any part of the body, including the earlobe.
Thestableandrobustbondingbetweenthedeviceandthenear-
field wireless chip allows for wireless power harvesting and data
communication. The stable and robust bonding between the
device and encapsulation layer can serve as an excellent shield
against ambient light. It should be noted that considering that
most of the skin/fingernails can introduce motion artifacts and the
opaque reflectance mode measurements are susceptible to motion
artifacts, signal-processing algorithms should be used to minimize
the motion-induced errors.
158,159
In addition to blood oxygen sensing, non-invasive monitoring
of other blood biomarkers (such as glucose and lactate) using
wearable devices was also investigated, mainly through optical
methods such as Raman and near-infrared spectroscopy. The
details could be found in this review.
160
For example, Yadav et al.
reported a sensor patch using a near-infrared (NIR) LED and
a photodiode to measure the diffuse reflectance spectra of
blood.
161
The in vivo results of this sensor patch showed a
significant difference in the blood glucose level before and after
a meal, indicating the potential of using NIR for glucose
monitoring. Recently a chemical-free blood lactate sensor was
developed with the use of microwave-range electromagnetic
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(EM) waves.
162
However, these devices usually suffer from poor
signal-to-noise ratios and are limited by calibration issues and
physiological factors. A significant improvement is needed to
replace the conventional blood glucose monitors.
3.5. Interstitial fluid (ISF) analysis
Despite the exciting progress in blood sensing, most chemical
analytes in the blood are still inaccessible non-invasively.
To bypass this sampling issue, researchers have looked at
monitoring the interstitial fluid (ISF) as a way to gain informa-
tion on chemical concentration in the blood. The composi-
tion of the ISF is very similar to blood in terms of small
molecules such as salts, proteins, glucose, and ethanol.
163
Over
the years, the ISF has been used for non-invasive diagnosis of
metabolic disorders, therapy assessment, and organ failure
assessment.
164–167
Numerous attempts were invasive with the
use of implantable devices. Yuen et al. also developed an
implantable glucose sensor with surface enhanced Raman
spectroscopy (SERS).
168
Silver films over nanosphere surfaces
functionalized with a mixed self-assembled monolayer (SAM)
were implanted in a rat and the glucose concentration was
monitored.
Non-invasive ISF monitoring devices based on reverse ion-
tophoresis (Fig. 9a) were investigated. Reverse iontophoresis,
similar to iontophoresis, applies a potential difference between
two electrodes on the skin. Ions in the body such as Na
+
act as
charge carriers and substances within the body can be extracted
through the skin. Since the skin has a net negative charge at
physiological pH, an electro-osmotic flow occurs along with
the iontophoretic current. Neutral molecules such as glucose
can be extracted to the cathode with Na
+
out of the skin.
169,170
The most well-known FDA-approved glucose sensing device,
GlucoWatch, utilized reverse iontophoresis to bring the ISF
through the skin to the externally attached sensor. The GlucoWatch
was a highly integrated wrist-watch device that features both
reverse-iontophoresis and biosensing. Amperometric electro-
chemical sensors were screen-printed; hydrogel discs with GOx
at a concentration gradient served as the reservoir for the collected
glucose. In spite of the excellent integration of its features, the
GlucoWatch requires a certain minimum duration for sufficient
Fig. 8 Wearable and flexible pulse oximeters. (a) Flexible all-organic optoelectronic pulse oximeter with a polymer LED and organic photodetectors.
Reproduced with permission from ref. 155. Copyright 2014 Nature Publishing Group. (b) Ultraflexible organic photonic skin with a multi-color polymer
LED. Reproduced with permission from ref. 156. Copyright 2016 American Association for the Advancement of Science. (c) Miniaturized flexible pulse
oximeter built on silicone elastomer and equipped with a near-field wireless chip. Reproduced with permission from ref. 157. Copyright 2017 Wiley.
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glucose extraction and it still suffered from inaccuracy,
171,172
interference from sweating,
160
and the product was removed
from the market due to filed reports on skin irritation and
blisters during continuous usage.
160
The reverse iontophoresis
technique was also used in other ISF sensors. Bandodkar et al.
introduced an all-printed flexible temporary tattoo-based
glucose sensor (Fig. 9b).
173
Both sensing electrodes and
reverse-iontophoresis electrodes (and agarose hydrogel coating)
were printed onto a tattoo paper, and the use of a GOx-modified
Prussian blue transducer allowed for the detection at a lower
potential than the GlucoWatch. The current density required for
ISF extraction was also significantly lower than the GlucoWatch.
The device was tested on healthy individuals for pre- and post-
meal glucose monitoring, and the results showed the rise in
the glucose level after a meal (Fig. 9c). Recently, Chen et al.
developed a flexible and non-invasive ISF glucose monitoring
system with electrochemical twin channels (ETC) and reverse
iontophoresis.
174
With the use of ETC, high-density hyaluronic
acid (HA) is transdermally repelled into the ISF under the
anode; the extra HA (positively charged) raises the ISF osmotic
pressure and thus promotes intravascular blood glucose trans-
port into the ISF (Fig. 9d). The increased glucose concentration
in the ISF can then increase the flux of reverse iontophoresis at
a low-current level. As a result, more ‘‘real’’ blood glucose is
measured and the blood–ISF glucose correlation is improved.
The biosensor consists of layers of poly(methyl methacrylate)
(PMMA), polyimide (PI), a sand-dune nanostructured gold thin
film, a transducer layer (PB), and a transfer/glucose oxidase
(GOx) immobilization layer (Fig. 9e). The device was tested on
human subjects with hourly measurement during a 1 day period,
Fig. 9 Wearable sensors for non-invasive monitoring of the interstitial fluid (ISF). (a) ISF sensing mechanism with the use of reverse iontophoresis.
Reproduced with permission from ref. 173. Copyright 2015 American Chemical Society. (b) All-printed tattoo-based ISF glucose sensor. Reproduced with
permission from ref. 173. Copyright 2015 American Chemical Society. (c) Tattoo-based sensor measured pre- and post-meal ISF glucose level on a
human subject. Reproduced with permission from ref. 173. Copyright 2015 American Chemical Society. (d) Application of electrochemical twin channels
(ETC) for increasing intravascular glucose transport into ISF and for increasing the flux of reverse iontophoresis. Reproduced with permission from
ref. 174. Copyright 2017 American Association for the Advancement of Science. (e) Biosensor layout for the ETC-based ISF glucose monitoring system.
Reproduced with permission from ref. 174. Copyright 2017 American Association for the Advancement of Science. (f) Human subject testing of the
ETC-based ISF glucose monitoring system over a day, with concurrent finger-pricked blood glucose measurements. Reproduced with permission from
ref. 174. Copyright 2017 American Association for the Advancement of Science.
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and the results showed a good match and a high correlation
between the ISF and blood glucose level, with a 1 hour time lag
in the ISF glucose concentration (Fig. 9f).
In addition to reverse iontophoresis, ISF extraction could be
achieved with sonophoresis, which employs 20 kHz ultrasound
to increase the permittivity of the skin to the interstitial fluid.
The glucose in the ISF then flows to the skin for electrochemical
sensor detection.
175
But this technique creates micropores in
the skin and was considered minimally invasive. Based on this
technique, Pu et al. developed a polyimide-based flexible electro-
chemical sensor integrated with an ISF-collecting microfluidic
chip for continuous glucose monitoring.
176
The authors thus
worked to increase the sensitivity at lower glucose concentration
by modifying graphene onto the working electrode with inkjet
printing and introducing gold nanoparticles onto the graphene
layer to improve uniform and enhanced electrochemical activity
and electron transfer rate.
177
In vitro testing showed promising
results for continuous glucose monitoring but on-body testing
was not yet accomplished. In addition to glucose monitoring,
ethanol was also continuously monitored using a wireless wear-
able device via poration of the stratum corneum of the skin.
178
Micropores (open for 3 days) with the diameter of human hair
were created by a handheld porator, and an ISF harvesting
system was positioned on the pore location to collect the fluid
with an electromechanical pump in the harvesting unit. The
electrochemical sensor inside the harvesting unit then measured
the ethanol concentration. Using a similar extraction method,
Venugopal et al. reported a wearable sensor for the continuous
assessment of cortisol using electrochemical impedance spectro-
scopy in the ISF.
179
As described above, there is great potential for ISF-based
biomarker monitoring, but several limitations should be noted
and addressed. The delay of glucose and ethanol response either
due to diffusion from the blood to the ISF or due to ISF extraction
and collection may impose delayed detection of dangerous con-
ditions (i.e. hyperglycemia or hypoglycemia). Skin irritation due
to continuous ISF extraction is also an issue to be addressed.
Sporadic sampling with on-off switch or the use of smaller reverse
iontophoresis current with conformal design for ISF extraction
may be applied to reduce such adverse effects. In addition,
interference from external glucose sources (i.e. sweat glucose
and glucose residue on skin) could impose inaccurate readings;
inconsistent ISF extraction efficiency over the collection area or
over time may also introduce inconsistency in the measured
results. Finally, a fully integrated sensing platform is yet to be
constructed, which requires sensor powering, signal processing,
and wireless communication.
3.6. Wound monitoring
Would healing undergoes three stages, including vasculariza-
tion, granulation, and re-epithelialization stages. In the vascu-
larization stage, growth factors and other proteins are released
and form a temporary matrix on the wound; in the granulation
stage, various biological components are secreted; in the
re-epithelialization stage, cell apoptosis and destruction of the
provisory matrix tissue occur.
180
The chemical composition of
wound exudates changes significantly over the three stages
and is crucial for proper healing, and relevant parameters such
as pH, uric acid and C-reactive protein (CRP) concentration of
wound exudates could indicate the stage of healing and the
presence of infection.
181–186
Monitoring the wound healing
process (wound management) can reduce hospitalization time,
prevent amputations, and aid in therapy studies. Of all wound-
related morbidities, chronic ulcer has affected 25% of diabetics
and is the leading cause of non-traumatic amputation in
developing countries;
187
it also occurs in around 1–2% of the
American population.
188
Chronic ulcers can be especially difficult
to treat and highly susceptible to infection. The use of wound
management would benefit these patients tremendously. Quali-
tative wound assessment was done by visual inspection, which
could be subjective and sometimes yields inconsistency due to
variation in lighting and angle. Quantitative assessment of
wound parameters was investigated over the years and wearable
wound management devices were constructed to monitor para-
meters such as pH, CRP concentration, hydration, skin tempera-
ture, etc.
189
Earlier wearable device based wound monitoring was
integrated as a system with a wound dressing patch.
190
Voirin
and coworkers utilized the optical detection method to measure
pH and CRP concentration. PHEMA/DMAEM hydrogels change
volume in response to pH variation, and OptoDex coating on the
substrate was used as a receptor for CRP. A sensitive layer was
deposited on a waveguide substrate; changes in hydrogel volume
and specific adsorption of CRP by the OptoDex coating lead to a
varied refractive index of the interface. This change in refractive
index was then detected using a spectrometer. The device was
used to monitor changes in pH and CRP concentration in serum
and observable changes were recorded.
Over the recent years, flexible sensors for continuous wound
monitoring were demonstrated. Mehmood et al. developed a
flexible telemetric sensing system to monitor pressure, tempe-
rature and moisture of wound.
191
Interfaced with off-the-shelf
sensors, a flexible circuit was fabricated and wireless RF data
transmission was achieved. Hattori et al. developed a sensor
platform that can softly and reversibly laminate at wounds and
micro-metal resistors provide precise measurement of tempera-
ture and thermal conductivity of the skin near the wounds.
192
The measurements were processed on a computer with 3D FEM
to yield time-dependent mapping of temperature and thermal
conductivity data. Punjiya et al. constructed a smart bandage
platform in chronic wound healing (Fig. 10a).
193,194
The smart
bandage measures wound pH with pH sensitive polyaniline
(PANI) coated threads, which increases the open circuit
potential when pH is lowered. Custom CMOS readout electro-
nics were used for wireless readout, and 2D pH mapping was
achieved. Guinovart et al. introduced a resistive pH wound
monitoring device that uses modified screen-printed Ag/AgCl
electrodes on a commercial bandage (Fig. 10b).
195
A PVB coated
membrane was used on the reference electrode (which displayed
a more stable potential compared to the bare solid Ag/AgCl
reference electrode), and the working electrode consisted of
PANI, which has potentiometric change due to pH change based
on the transition between emeraldine salt and emeraldine base.
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Kassal et al. constructed a uric acid sensing bandage with a
wireless connection feature.
196
Prussian blue (PB) modified
carbon electrodes were screen printed onto a commercial
bandage, and uricase was then immobilized on the working
electrode (Fig. 10c). Uric acid (UA) oxidation produces hydrogen
peroxide, which was reduced catalytically by the PB-carbon
electrode; in this way, UA detection was achieved at a very low
negative working potential. The bandage was connected to a
potentiostat that measures and stores biosensor current output
and transfers data wirelessly. Liu et al. fabricated a uric acid
sensor with conductive thread and coated the polyester thread
with Ag/AgCl or carbon ink.
197,198
Sensors were then embroi-
dered and uricase was immobilized onto the working electrode.
The embroidered structure has stable and consistent output in
simulated wound fluid between flat and bending conditions.
Mostafalu et al. demonstrated a smart bandage encased in a
stretchable and flexible 3D-printed platform for wound oxygen
monitoring (Fig. 10d).
199
Silver and electroplated zinc electrodes
on parylene-C were used for oxygen sensing, and a thin layer of
PDMS was used as an oxygen-selective membrane for the sensor.
Farooqui et al. utilized an inkjet-printed smart bandage to
continuously monitor irregular bleeding, pH and external pres-
sure and wirelessly transmit output data (Fig. 10e).
200
Capacitive
sensing was used for bleeding and pressure sensing: irregular
bleeding causes a change in the dielectric constant between two
electrodes on either side of the bandage strip; pressure leads to a
change in the distance between the two electrodes. Resistive
sensing was used in pH measurements: the conductivity of the
carbon-based electrode varies when the electrode is exposed to
solutions of different pH, and the change in resistance is
Fig. 10 Wearable sensors for wound monitoring. (a) Smart bandage for chemical sensing of wound pH using pH-sensitive PANI coated threads. CMOS
wireless readout and 2D mapping of pH levels were incorporated. Reproduced with permission from ref. 193 and 194. Copyright 2017 IEEE. (b) Wound pH
wound monitoring with the PANI working electrode and the PVB-coated reference electrode. Reproduced with permission from ref. 195. Copyright 2014
Wiley. (c) Wound uric acid sensing bandage with wireless readout. Uricase was immobilized on carbon working electrodes modified with Prussian blue.
Reproduced with permission from ref. 196. Copyright 2015 Elsevier. (d) Wound oxygen monitoring with a smart bandage encased in a 3D-printed flexible
platform. Zn–Ag electrodes were used for oxygen sensing. Reproduced with permission from ref. 199. Copyright 2015 IEEE. (e) Inkjet-printed smart
bandage for continuous monitoring of irregular bleeding, pH and external pressure using capacitive and resistive sensing. Reproduced with permission
from ref. 200. Copyright 2016 Nature Publishing Group.
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measured for pH sensing. A double-sided detachable PCB was
inkjet printed on Kapton adhesive tape, and an on-chip micro-
controller was used to store measured data. The device was
tested on human skin and it exhibited great mechanical resi-
lience and displayed a text of ‘‘change bandage’’ when bleeding
happened on the site. Melai et al. used the graphene oxide
coated working electrode on a screen-printed substrate to moni-
tor pH in wounds.
201
A simultaneous pH and glucose sensing
wound patch was recently developed.
202
The parameters were
measured via measuring fluorescence signals from a pH indi-
cator sensitive dye-carboxynaphtho-fluorescein probe, and a
metabolite-sensing enzymatic system (based on glucose oxidase
and horseradish peroxidase) that responds to glucose concen-
tration. The system was able to detect low glucose concentration
in artificial wound exudate. Since the results could be seen by
visual inspection of the fluorescence intensity, it provides direct
and qualitative wound status information to users, and can help
to distinguish between an autonomously healing and a chronic
wound at an early stage.
Although multiple wearable devices have been developed to
provide useful information on wound healing, multi-parameter
sensing was yet to be achieved to give a more comprehensive
understanding of the wound healing status. The devices
described above either performed in vitro testing on wound
exudate or on human skin for a relatively short time, and thus
did not examine if wearing such modified bandages for long
time may have adverse effects on the wound healing process or
lead to discomfort for users. Since the wound healing process
spans over at least a few days, continuous monitoring could be
achieved by sporadic sampling, but the replacement of sensor-
bandages would require high consistency in sensing perfor-
mance between individual sensor-bandage. If the same piece of
bandage could stay on wound for extended time, storage of
measured data or real-time wireless data transmission to an
external device is needed for the temporal progress of the
wound healing status.
3.7. Breath analysis
Breath is the gas–vapor mixture that humans exhale through
either nasal or oral cavity. The potential for breath measurement
is enormous, since it is completely noninvasive and inherently
safe.
203
Breath collection can be done on any individual who can
breathe, even including neonates and Alzheimer patients. The
composition of breath is complex: it includes a mixture of
nitrogen, oxygen, CO
2
, and water vapor. The remainder includes
as many as 500 different compounds, either endogenous (pro-
duced by physiological processes) or exogenous (ingested food,
air, etc.). Among these endogenous compounds are acetone,
carbon monoxide, ammonia, and nitric oxide.
204
The breath
carbon monoxide test has been used for neonatal jaundice
diagnosis; breath ammonia can be used for assessment of
asthma and hemodialysis; the breath nitric oxide test was used
to monitor the asthma therapy process; breath acetone could be
used for diabetes monitoring in place of finger-pricking, etc.
204–208
In addition to breath vapor composition, breath rate and breath
condensate are also used for analysis. The breath rate profile is
a useful analysis for individuals with sleep apnea, asthma, and
chronic obstructive pulmonary disease (COPD).
204,209,210
Breath
condensate, a liquid mixture of chilled breath, contains fluids
from airway-lining and pulmonary tissues. Species such as
cytokines, reactive oxygen chemicals, and leukotrienes were found
in breath condensate and could be used for disease diagnosis and
monitoring.
204,209,210
Breath sensing technology has been developed since the
1970s, and most sensing processes require bulky equipment for
breath collection. A portable breath sensing device was developed
in the late 2000s: using the electrostatic self-assembly of super-
hydrophilic SiO
2
nanoparticles, Corres et al. built a humidity
sensor and integrated it into an optical fiber.
211
The sensor
performance has good reproducibility and low hysteresis; when
the sensor was coupled with methylene blue dye, fast response
was achieved and characteristic hyperventilation respiration
patterns were detected. A portable breath acetone sensor was
developed.
212
The Si-doped epsilon-WO
3
nanoparticle was used as
the sensing material due to its high sensitivity and selectivity to
acetone at high relative humidity, which enabled low-concentration
detection of acetone. Si-Doped epsilon-WO
3
nanostructured
films were flame-deposited and annealed in situ. During human
testing, the sensor measured acetone concentration conti-
nuously when the test person was at rest and during physical
activity. The results were sensitive and consistent with those
detected by the standard proton transfer reaction mass spectro-
metry (PTR-MS) technique. Borini et al. also developed a flexible
breath humidity sensor with graphene oxide (Fig. 11a) at a
wearable size, and measured the sensor performance using a
controlled humidity generator.
213
The impedance of the graphene oxide sensor varies accord-
ing to the relative humidity and temperature. The 2D graphene
oxide film at 15 nm thickness reached ultrafast response to a
modulated humid flow (B30 ms response and recovery time)
and maintained full-scale output at the same time. The device
is highly flexible (Fig. 11a) and transparent, and its wearable
size and low cost of production made it a good candidate for
continuous breath analysis. The sensor response can also
distinguish different patterns of human behaviors (speaking
vs. breathing, whistle tunes by different users), which may be
useful for user recognition in breath monitoring. Using the
humidity change as an indicator of exhalation, Caccami et al.
measured the graphene oxide resistance change to monitor
respiratory activity in real time (Fig. 11b).
214
A flexible antenna
was prototyped on Kapton to enable RFID and wireless data
transmission. The antenna and the graphene oxide sensor were
integrated onto a facemask, and this wearable device was tested
on a subject. The results showed its ability to monitor indi-
vidual exhalation–inhalation peaks and distinguish normal
breath from apnea conditions. Wang et al. developed a flexible
chemical sensor for breath acetone measurement (Fig. 11c).
215
Using a 3D biomimetic design resembling a hierarchical butterfly
wing, a porous chitosan–graphene oxide biocomposite was
synthesized. The structure allows for excellent mechanical
strength and flexibility in addition to biocompatibility and
electronic activity. This biocomposite layer was sandwiched
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between copper electrodes and paper substrate and the impe-
dance changed according to acetone concentration. The sensor
was implemented as a flexible wristband and tested withpulsated
ejection of simulated diabetic breath flow over the sensor (2 ppm
acetone vapor with 85% relative humidity). The sensing results
were calibrated with relative humidity to prevent overestimation
of acetone concentration. The device exemplifies great sensitivity
and specificity in real-time acetone monitoring. Ammu et al. used
inkjet-printed single-walled carbon nanotube (CNT) films on PET
and cellulosics (Fig. 11d) substrates to detect nitrogen dioxide
and chlorine.
216
The penetration of the gas molecules into
the CNT films increases the distance between the conducting
pathways of CNTs, causing the resistance changes of the sensor.
The inkjet-printed CNT films showed a decrease in resistance for
NO
2
and Cl
2
as compared to an increase in resistance for other
common gas vapors (Fig. 11d). Without the use of the vapor
concentration process, NO
2
vapor was detected at as low as
125 ppb and Cl
2
vapor was detected at as low as 500 ppb in
ambient air. The inkjet printing process helped to control the
film thickness and maintain consistency among sensors. The
group experimented with CNT films printed on a PET substrate
and CNT films printed on cellulosics (paper and cloth) to test
sensor recovery performance: the cellulosic substrates aided in
spontaneous gas recovery in an ambient environment at room
temperature, and the PET substrates required photoirradiation
for sensor recovery.
In addition to single-analyte sensing, the aggregation of
chemical sensors into arrays was complemented with multi-
variate data analysis to yield quantification and identification
of compounds with combinatorial selectivity. Recognition of a
specific compound or a specific mix of compounds is done by
all sensors, which is parallel to the behavior of the human
olfaction sensor. Due to this similarity, these gas sensor arrays
were called ‘‘electronic noses’’ (‘‘e-nose’’). The concept was
Fig. 11 Wearable and flexible sensors for gas/breath analysis. (a) Graphene-oxide-based flexible breath humidity sensor capable of measuring breath
humidity and distinguishing human behaviors such as whistling. Reproduced with permission from ref. 213. Copyright 2013 American Chemical Society.
(b) Graphene-oxide-based sensor for respiratory activity (breath humidity) monitoring. A flexible antenna was integrated for RF identification and wireless
data transmission. The device took the form of a face mask. Reproduced with permission from ref. 214. Copyright 2017 IEEE. (c) Flexible acetone
chemical sensor based on a porous chitosan–graphene oxide biocomposite. Impedance variation was observed under pulsated ejection of simulated
diabetic breath. Reproduced with permission from ref. 215. Copyright 2017 Royal Society of Chemistry. (d) Flexible inkjet-printed single-walled carbon
nanotube (CNT) based chemical sensor for Cl
2
and NO
2
sensing. Decrease in sensor resistance was observed for Cl
2
and NO
2
vapor, compared to other
common gas vapors. Reproduced with permission from ref. 216. Copyright 2012 American Chemical Society. (e) Flexible exhaled breath sensor-array
based on modified gold-nanoparticles for ovarian carcinoma (OC) diagnosis. Discriminant factor analysis (DFA) separated breath samples collected from
subjects with OC as well as from controls. Reproduced with permission from ref. 221. Copyright 2015 American Chemical Society. (f) Flexible printed
CNTs/polymer sensor-array for armpit odor monitoring (e.g. ammonia, acetic acid).
223
Reproduced with permission from ref. 223. Copyright 2014 MDPI.
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demonstrated in the early 1980s but used for gas analysis. Since
the early 2000s, electronic noses have been used in breath
analysis, especially for classification of individuals into pathology
of interest. Many studies have used electronic noses for diagnosis
of diseases such as lung cancer, pneumonia, asthma, COPD,
tuberculosis, kidney functions, neurological diseases, etc.
217–219
Based on this electronic nose concept, Seesaard et al. developed a
fabric-based wearable chemical amine sensor prototype with
carboxylic-functionalized single-walled carbon nanotube/polymer
(SWNT–COOH) nanocomposite.
220
4 types of polymers were used
to sense different analytes such as ethanol, ammonium hydro-
xide, pyridine, etc. The composite solution was drop-coated onto
interdigitated electrodes prepared by embroidery of conductive
thread into fabrics. The fabricated sensor array was used for odor
analysis of two individuals on their urine, armpit and exhaled
breath samples ex vivo. Principal component analysis (PCA)
discrimination showed that the fabricated electronic nose was
capable of distinguishing the odor profile of two different
individuals. Haick and coworkers introduced a flexible sensor
array based on molecularly modified gold nanoparticles for
diagnosis of ovarian carcinoma (OC) from exhaled breath
(Fig. 11e).
221
The flexible sensor could selectively detect ppb
level of volatile organic compounds (VOCs) that are related with
ovarian cancers in exhaled breath and could distinguish them
from environmental VOCs. Through the use of discriminant
factor analysis (DFA) based on the leave-one-out method, in
actual breath samples from subjects with OC as well as from
controls, the sensitivity, selectivity and accuracy were found to be
81.3%, 82.9% and 83.6%, respectively. The same group further
developed a self-healable and flexible sensing platform based on
functionalized gold nanoparticle films that can be used for
sensing pressure variation as well as 11 kinds of VOCs.
222
Lorwongtragool et al. demonstrated a novel wearable electronic
nose based on flexible printed CNTs/polymer sensor array for
monitoring of armpit odor (Fig. 11f).
223
This e-nose based sensor
was fabricated on a polyethylene naphthalate (PEN) substrate
through a low cost and scalable inject-printing technique. Both
composite-like layer and composite film of MWCNTs/polymer
were used as the sensing layers. The sensor was designed as a
compact armband and showed feasibility to detect VOCs (e.g.
ammonia, acetic acid, acetone and ethanol) released from the
armpit regions of the human body.
Despite recent advances on breath sensing, challenges still
exist in many aspects. Inherent breath confounders (i.e. con-
tamination from ambient air, interference from humidity,
ingested materials, etc.) are difficult to eliminate from breath
collection and analysis. Sensing of biomarkers in gas is signifi-
cantly influenced by humidity. Exposure to water vapor will
induce decreased resistance of n-type oxide semiconductors
and thus affect the gas-triggered response. To resolve humidity
Fig. 12 Fusion of wearable and flexible physical and chemical sensors. (a) Electrocardiogram sensing integrated with lactate sensing. Exercising subject
testing yields reasonable correlation between lactate levels and heart rate profile. Reproduced with permission from ref. 224. Copyright 2016 Nature
Publishing Group. (b) Temperature sensing integrated with multi-analyte sensing. Temperature compensation improved the consistency of lactate and
glucose measurement at different temperatures. Reproduced with permission from ref. 55. Copyright 2016 Nature Publishing Group.
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interference, dehydration via increased sensor temperature might
be considered; humidity absorbing materials such as NiO may be
used to separate water molecules from the target analyte sensor,
and regeneration of such absorbers is needed for continuous
monitoring.
224
The devices developed were mostly tested in a
controlled environment with low air disturbance, which is less
realistic compared to a real-life situation. Future wearable devices
should be able to accommodate different ambient conditions
including temperature, humidity, and airflow. Lastly, full device
integration for in situ real-time wireless monitoring was not yet
accomplished in many wearable breath sensors. Continuous mon-
itoring for diseases such as asthma will require a fully integrated
flexible device with wireless readout and steady yet wearable power
supply. In addition to the challenges described above, the concept
of electronic nose could be expanded to biofluid analysis, combi-
natorial analysis of gas and biofluids, or among different biofluids.
This concept may benefit disease classification in a broader
spectrum than simply gas-based classification.
3.8. Fusion of physical and chemical sensors
Aside from wearable sensors based solely on chemical sensing
principles, various wearable sensors fused both physical and
chemical sensing onto a single platform. Such physical parameters
include body temperature, pressure, strain, and heart rate. The
implementation of these physical parameters, in addition to
chemical parameters, can be used for more comprehensive evalua-
tion of physical conditions such as wound healing progress,
191
diabetic ulceration,
187
etc. Imani et al. integrated electrocardiogram
with lactate sensing on a thin, flexible polyester patch (Fig. 12a).
224
The patch consists of a screen-printed three-electrode ampero-
metric lactate sensor and two electro-cardiogram electrodes, which
enabled simultaneous measurement of sweat lactate and an
electrocardiogram in real time. The device was placed on the
chests of exercising human subjects and recorded sweat lactate
concentration and heart rate in real time.
The measured physical parameters can also be used to calibrate
chemical sensing results.
55,56
For instance, enzyme activity usually
varies with temperature, and enzyme-based chemical sensors can
be strongly affected by temperature variation in real-life applica-
tions. The use of temperature measurement to calibrate chemical
sensing results aids in the consistency in sensor performance over
time. The flexible sensor that Gao et al. developed incorporated
physical based temperature monitoring with electrochemical
detection of ions and metabolites (Fig. 12b).
55
As the temperature
increases from 22 1Cto401C, an increase in GOx and LOx activity
wasobservedwhichledtooverestimationofactualglucoseand
lactate concentration. With the real time temperature sensing
and compensation, accurate and consistent readings of glucose
and lactate concentration were achieved.
4. Conclusions and outlook
In this review, we have summarized and highlighted recent
advances in wearable flexible chemical sensors toward conti-
nuous health monitoring. These include wearable chemical
sensors for non-invasive biomarker analysis in tears, saliva,
sweat, ISF and blood, as well as sensors that can detect the gas
molecules from exhaled breath. Unlike most previously reported
wearable sensors that mainly tracked physical activities and vital
signs, the new generation of wearable chemical sensors enable
real-time and fast detection of accessible biomarkers from the
human body, and allow for collection of large-scale information
about the individual’s dynamic health status at the molecular
level. A number of key features have been identified as the
advantages of the development of the wearable chemical sensors:
flexibility, stretchability, biocompatibility, low-cost production
and real-time continuous monitoring. These new non-invasive
wearable and flexible chemical sensors have promising prospects
in a variety of healthcare fields. A key example is that they provide
low-cost solutions for early diagnosis, real-time in-home monitor-
ing and management of chronic diseases. The affordable wireless
health-monitoring devices will also lead to major improvements
in patient monitoring, particularly for the developing countries or
rural areas where the medical resources are limited, unavailable,
expensive, and ineffective.
Despite significant progress made in the past few years
in the field of wearable and flexible chemical sensors toward
continuous health monitoring, there are many key challenges
to address and technological gaps to bridge before realizing the
full potential of wearable chemical sensors. One major problem
is that, the amount of available wearable chemical sensors that
can accurately, reliably and continuously analyze the broad
spectrum of the body biomarkers is still very limited. Recently
developed wearable chemical sensors mainly focus on monitoring
major metabolites and electrolytesinbodyfluids.Mostofthese
sensors still require further optimization in sensor stability and
assessment in human trials before exploitation and routine use
can happen. In addition, there are few reports on wearable devices
that can monitor peptides, hormones, proteins, and DNAs/RNAs
in body fluids or in exhaled breath. In fact, sweat, saliva, tears and
exhaledbreathcontainawealthofsuchanalytesintrace
amounts, many of which are reported to have good correlations
with blood analytes or close relations with various health condi-
tions and diseases. Continuous monitoring of these biomarkers
through wearable devices would provide insightful information
for screening and early diagnosis of a broad range of major health
conditions. However, one of the challenges lies in the detection of
the extremely low concentrations of these biomarkers in body
fluids. Developing highly sensitive and selective sensors with
proper pre-concentration techniques is one method of tackling
this problem. Moreover, current protein or DNA sensors usually
require multi-step preparation protocols with long waiting time
and additional washing steps, which are not desirable for con-
tinuous wearable monitoring. Next-generation wearable chemical
sensors must explore novel materials and detection techniques to
target these challenging biomarkers to further expand the reach of
wearable sensing technology for non-invasive, personalized health
monitoring. Considering the complex physiological process of the
human body, multiplexed chemical sensing or fusion of sensors
for physical and chemical analysis could be extremely important
to obtain accurate and insightful physiological information.
Review Article Chem Soc Rev
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Although wearable sensing has achieved great advances in
the past few years, several key milestones are yet to be reached.
One of the milestones is powering the wearable devices. The
tremendous progress in wearable biosensing and the increasing
need for multi-tasking (sensing, processing, and communication)
on wearable platforms have urged the development of efficient
and sustainable power sources. For non-invasive skin-worn wear-
able devices, the power sources must also be flexible and stretch-
able to comply with skin contour and cope with mechanical
stress. To this end, flexible solar cells, piezoelectric devices,
thermoelectric devices, and wearable batteries were constructed
in the past as abiotic energy sources.
225–227
In addition, recent
studies have cast light on wearable electrochemical biofuel cells
energy harvesting from biotic sources and showed the great
potential of biofuel cells for wearable applications. Though the
biocompatibility and auto-power features of biotic energy sources
have made biofuel cells a great candidate for wearable applica-
tions, the long-term stability of the enzymes and redox mediator
(if present) need to be improved for proper function of the biofuel
cell. Stabilization of the enzymes should be considered and could
possibly be achieved by engineering modified enzymes,
228
by
applying enzyme stabilizers
229
or using better enzyme immobili-
zation methods.
230
Although great efforts have been invested in
non-invasive flexible fuel cells and biofuelcells, the current power
is generally insufficient to power many current wearable electro-
nic devices. In addition to increasing the power output of such
flexible fuel cell, an alternative is to create miniaturized sensing
devices with lower power consumption; using multi-source
energy harvest may be also promising.
Another milestone for many wearable sensing devices lies in
data processing and seamless system-level integration. In order
to acquire interpretable results from wearable sensing devices,
the electrical signals from sensors need to be properly processed
and then transmitted for analysis and display. The electrical
signals need to be converted into recordable values and sampled
by the processor. Raw data sampled can be susceptible to
inherent or ambient noises, and signal processing reduces these
noise effects to extract useful signals from sensors. Transmission
of the processed data to an external platform can allow for display
and analysis of the processed data. In many cases, mathematical
manipulation (i.e. difference between electrical potentials, rate-
time profile) that requires specific analytics or even algorithm is
needed to gain interpretable results corresponding to the para-
meter of interest. The main function of processing is noise
reduction in the recorded data, but researchers have recently
incorporated mathematical calculation and digital processing
into the on-board platform.
231
For applications that require large
data storage and complex computation, processed data need to
be transmitted to a computing device (i.e. mobile phone, com-
puter). Although wired connection remains an option, it imposes
inconvenience on users and thus not desirable in the final device
layout. At present, bluetooth low energy (BLE) and near-field
communication (NFC) have been widely used in wearable sensor
platforms and allow for real-time data streaming and analysis.
However, BLE can be a huge drain for power supply, and NFC
requires close proximity with receiver electronics. In addition to
respective constraints, both technologies are not suitable for
high-density data transmission, where several users have multi-
ple sensors interacting with receivers at a high sampling rate. A
transmission that fulfills ideal connectivity is yet to be developed.
The analytics after signal transmission should also be carefully
designed to convey relevant, useful and interpretable information
to users. This will require development of efficient algorithms
and user-friendly display platforms. The remote storage and
processing of these personal data introduce concerns about data
security and user privacy, and intensive research efforts have
been made on cryptologic algorithms.
54,231
The wearable sensing technology will not be realized without
efficient commercialization. To date, most current commercial
effort on wearable sensing devices has been devoted to the
adaptation of current sensing methods through system minia-
turization and conformal design. There is still a lack of wearable
sensing products for molecular monitoring in the commercial
market, due to issues such as sensor stability, biofouling and
sensor degradation over time. Therefore, future commercial
products may require strategies for easy replacement of sensing
components (e.g. disposable and low cost chemical sensors) with
reusable electronics components. Although there are still no
commercially available wearable chemical sensors for conti-
nuous monitoring of human sweat, tears, interstitial fluids or
saliva, a number of start-up companies such as Eccrine Systems
(sweat monitoring), Kenzen (sweat monitoring) and MouthSense
(saliva monitoring) are receiving increased attention and invest-
ment. As the interest grows, more endeavors from both research
and clinical fields will help identify strategies to resolve the
current limitations and issues related to chemical sensors. It is
expected that the commercial wearable chemical monitor will be
available to the customers within the next decade.
Although current research efforts on wearable chemical sens-
ing technology primarily focus on developing reliable wearable
and flexible chemical sensing systems, as the technologies
move forward, understanding the physiological relevance of bio-
markers in body fluids and exhaled breath to determine their
utility for non-invasive health monitoring will be the main
bottleneck to be addressed toward a broader use of wearable
chemical sensors. To this end, accumulation of large sets of data
across longitudinal and cohort studies can help to investigate
the potential correlations and generate possible predictive algo-
rithms. These studies present a key challenge to sweat sensor
technology development in the immediate future and would
require the close collaboration among engineers, scientists as well
as clinicians. An attractive long-term vision of wearable devices is
to heterogeneously integrate a wide range of sensor networks
(e.g. biomolecules, vital signs) on small wearable patches that
can continuously and non-invasively monitor the user’s health.
Big data analytics and machine learning tools can be used
towards parsing the vast time series of multiplexed sensor data
from these population studies to identify subtle patterns and
correlations. The resulting system, coupled with the big data
mined from this technology, will supplant traditional reactive,
episodic healthcare with predictive and proactive diagnostics.
Such technology can pave the way for future development in
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Chem.Soc.Rev. This journal is ©The Royal Society of Chemistry 201 8
wearable sensors for personalized, real time health analysis.
In certain cases, upon further investigation and development,
such non-invasive wearable electrochemical sensors can poten-
tially be used as a counterpart to blood testing not only for
patients in medical applications but also for the general public in
daily life. Through cohort studies using the technology develop-
ment of wearable and flexible electronics for continuous mole-
cular analysis, we expect that the large sets of data and health
information collected using wearable sensors from individuals’
daily activities will ultimately generate predictive algorithms and
revolutionize the traditional healthcare setting.
Conflicts of interest
There are no conflicts to declare.
Acknowledgements
This work is supported by the California Institute of Technology.
References
1 M. A. Hamburg and F. S. Collins, N. Engl. J. Med., 2010,
363, 301–304.
2 J. Larry Jameson and D. L. Longo, Obstet. Gynecol. Surv.,
2015, 70, 612–614.
3 J. Gubbi, R. Buyya, S. Marusic and M. Palaniswami, Future
Gener. Comput. Syst., 2013, 29, 1645–1660.
4 M. Hassanalieragh, A. Page, T. Soyata, G. Sharma, M. Aktas,
G. Mateos, B. Kantarci and S. Andreescu, Proc. IEEE Int.
Conf. Serv. Comput., 2015, 285–292.
5 S. Patel, H. Park, P. Bonato, L. Chan and M. A. Rodgers,
J. NeuroEng. Rehabil., 2012, 9, 1–17.
6 S. C. Mukhopadhyay, IEEE Sens., 2015, 15, 1321–1330.
7 M. Stoppa and A. Chiolerio, Sensors, 2014, 14, 11957–11992.
8 D.-H. Kim, R. Ghaffari, N. Lu and J. A. Rogers, Annu. Rev.
Biomed. Eng., 2012, 14, 113–128.
9 D.-H. Kim, N. Lu, R. Ma, Y. Kim, R. Kim, S. Wang, J. Wu,
S. M. Won, H. Tao, A. Islam, K. J. Yu, T. Kim, R. Chowdhury,
M.Ying,L.Xu,M.Li,H.-J.Chung,H.Keum,M.McCormick,
P.Liu,Y.Zhang,F.G.Omenetto,Y.Huang,T.Colemanand
J. A. Rogers, Science, 2011, 333, 838–843.
10 M. L. Hammock, A. Chortos, B. C. K. Tee, J. B. H. Tok and
Z. Bao, Adv. Mater., 2013, 25, 5997–6038.
11 J.-Y. Sun, X. Zhao, W. R. K. Illeperuma, O. Chaudhuri,
K. H. Oh, D. J. Mooney, J. J. Vlassak and Z. Suo, Nature,
2012, 489, 133–136.
12 J. A. Rogers, T. Someya and Y. Huang, Science, 2010, 327,
1603–1607.
13 M. Kaltenbrunner, T. Sekitani, J. Reeder, T. Yokota,
K. Kuribara, T. Tokuhara, M. Drack, R. Schwo
¨diauer,
I. Graz, S. Bauer-Gogonea, S. Bauer and T. Someya, Nature,
2013, 499, 458–463.
14 A. M. Hussain and M. M. Hussain, Adv. Mater., 2016, 28,
4219–4249.
15 S. Choi, H. Lee, R. Ghaffari, T. Hyeon and D.-H. Kim, Adv.
Mater., 2016, 28, 4203–4218.
16 K. Chen, W. Gao, S. Emaminejad, D. Kiriya, H. Ota,
H. Y. Nyein, K. Takei and A. Javey, Adv. Mater., 2016, 28,
4397–4414.
17 M. Amjadi, A. Pichitpajongkit, S. Lee, S. Ryu and I. Park,
ACS Nano, 2014, 8, 5154–5163.
18 C. Pang, C. Lee and K. Y. Suh, J. Appl. Polym. Sci., 2013, 130,
1429–1441.
19 D.J.Lipomi,M.Vosgueritchian,B.C.-K.Tee,S.L.Hellstrom,
J. A. Lee, C. H. Fox and Z. Bao, Nat. Nanotechnol., 2011, 6,
788–792.
20 Y. Gao, H. Ota, E. W. Schaler, K. Chen, A. Zhao, W. Gao,
H. M. Fahad, Y. Leng, A. Zheng, F. Xiong, C. Zhang,
L.-C. Tai, P. Zhao, R. S. Fearing and A. Javey, Adv. Mater.,
2017, 29, 1701985.
21 B. Nie, R. Li, J. Cao, J. D. Brandt and T. Pan, Adv. Mater.,
2015, 27, 6055–6062.
22 L. Pan, A. Chortos, G. Yu, Y. Wang, S. Isaacson, R. Allen,
Y. Shi, R. Dauskardt and Z. Bao, Nat. Commun., 2014,
5, 3002.
23 J. Lee, H. Kwon, J. Seo, S. Shin, J. H. Koo, C. Pang, S. Son,
J. H. Kim, Y. H. Jang, D. E. Kim and T. Lee, Adv. Mater.,
2015, 27, 2433–2439.
24 C. Wang, D. Hwang, Z. Yu, K. Takei, J. Park, T. Chen, B. Ma
and A. Javey, Nat. Mater., 2013, 12, 899–904.
25 K. Takei, T. Takahashi, J. C. Ho, H. Ko, A. G. Gillies,
P. W. Leu, R. S. Fearing and A. Javey, Nat. Mater., 2010,
9, 821.
26 Z. Wen, M. H. Yeh, H. Guo, J. Wang, Y. Zi, W. Xu, J. Deng,
L. Zhu, X. Wang, C. Hu, L. Zhu, X. Sun and Z. L. Wang,
Sci. Adv., 2016, 2, e1600097.
27 H. Wang, P. Hu, J. Yang, G. Gong, L. Guo and X. Chen, Adv.
Mater., 2015, 27, 2348–2354.
28 S. P. Lacour, G. Courtine and J. Guck, Nat. Rev. Mater.,
2016, 1, 16063.
29 T. Q. Trung and N.-E. Lee, Adv. Mater., 2016, 28, 4338–4372.
30 Y. Khan, A. E. Ostfeld, C. M. Lochner, A. Pierre and
A. C. Arias, Adv. Mater., 2016, 28, 4373–4395.
31 A. Chortos, J. Liu and Z. Bao, Nat. Mater., 2016, 15,
937–950.
32 G. Schwartz, B. C. Tee, J. Mei, A. L. Appleton, D. H. Kim,
H. Wang and Z. Bao, Nat. Commun., 2013, 4, 1859.
33 S. Y. Kim, S. Park, H. W. Park, D. H. Park, Y. Jeong and
D.-H. Kim, Adv. Mater., 2015, 27, 4178–4185.
34 J. Y. Sun, C. Keplinger, G. M. Whitesides and Z. Suo, Adv.
Mater., 2014, 26, 7608–7614.
35 C. Dagdeviren, Y. Su, P. Joe, R. Yona, Y. Liu, Y. S. Kim,
Y. Huang, A. R. Damadoran, J. Xia, L. W. Martin, Y. Huang
and J. A. Rogers, Nat. Commun., 2014, 5, 4496.
36 X. W. Wang, Y. Gu, Z. P. Xiong, C. Zheng and T. Zhang,
Adv. Mater., 2014, 26, 1336–1342.
37 B. Xu, A. Akhtar, Y. Liu, H. Chen, W. H. Yeo, S. I. Park,
B. Boyce, H. Kim, J. Yu, H. Y. Lai, S. Jung, Y. Zhou, J. Kim,
S. Cho, Y. Huang, T. Bretl and J. A. Rogers, Adv. Mater.,
2016, 28, 4462–4471.
Review Article Chem Soc Rev
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
This journal is ©The Royal Society of Chemistry 201 8 Chem.Soc.Rev.
38 B. Zhu, Z. Niu, H. Wang, W. R. Leow, H. Wang, Y. Li,
L. Zheng, J. Wei, F. Huo and X. Chen, Small, 2014, 10,
3625–3631.
39 S. Xu, Y. Zhang, L. Jia, K. E. Mathewson, K.-I. Jang, J. Kim,
H. Fu, X. Huang, P. Chava, R. Wang, S. Bhole, L. Wang,
Y. J. Na, Y. Guan, M. Flavin, Z. Han, Y. Huang and
J. A. Rogers, Science, 2014, 344, 70–74.
40 R. C. Webb, Y. Ma, S. Krishnan, Y. Li, S. Yoon, X. Guo,
X. Feng, Y. Shi, M. Seidel, N. H. Cho, J. Kurniawan, J. Ahad,
N. Sheth, J. Kim, J. G. Taylor, J. T. Darlington, K. Chang,
W. Huang, J. Ayers and A. Gruebele, Sci. Adv., 2015,
1, e1500701.
41 D.-H. Kim, N. Lu, R. Ghaffari, Y.-S. Kim, S. P. Lee, L. Xu,
J. Wu, R.-H. Kim, J. Song, Z. Liu, J. Viventi, B. de Graff,
B. Elolampi, M. Mansour, M. J. Slepian, S. Hwang,
J. D. Moss, S. M. Won, Y. Huang, B. Litt and J. A. Rogers,
Nat. Mater., 2011, 10, 316–323.
42 Y. Liu, J. J. Norton, R. Qazi, Z. Zou, K. R. Ammann, H. Liu,
L. Yan, P. L. Tran, K.-I. Jang, J. W. Lee, D. Zhang,
K. A. Kilian, S. H. Jung, T. Bretl, J. Xiao, M. J. Slepian,
Y. Huang, J.-W. Jeong and J. A. Rogers, Sci. Adv., 2016,
2, e1601185.
43 Y. Liu, M. Pharr and G. A. Salvatore, ACS Nano, 2017, 11,
9614–9635.
44 Y. S. Rim, S.-H. Bae, H. Chen, N. D. Marco and Y. Yang,
Adv. Mater., 2016, 28, 4415–4440.
45 T. Q. Trung and N.-E. Lee, Adv. Mater., 2016, 28, 4338–4372.
46 M. Gao, L. Li and Y. Song, J. Mater. Chem. C, 2017, 5,
2971–2993.
47 J. S. Heo, J. Eom, Y.-H. Kim and S. K. Park, Small, 2017,
14, 1703034.
48 A. Nag, S. C. Mukhopadhyay and J. Kosel, IEEE Sens. J.,
2017, 17, 3949–3960.
49 A.J.Atkinson,W.A.Colburn,V.G.DeGruttola,D.L.DeMets,
G.J.Downing,D.F.Hoth,J.A.Oates,C.C.Peck,R.T.
SchooleyandB.A.Spilker,Clin.Pharmacol.Ther., 2001, 69,
89–95.
50 K. Strimbu and J. A. Tavel, Curr. Opin. HIV AIDS, 2010, 5,
463–466.
51 A. Hulanicki, S. Glab and F. Ingman, Pure Appl. Chem.,
1991, 63, 1247–1250.
52 A. J. Bandodkar and J. Wang, Trends Biotechnol., 2014, 32,
363–371.
53 G. Matzeu, L. Florea and D. Diamond, Sens. Actuators, B,
2015, 211, 403–418.
54 A. J. Bandodkar, I. Jeerapan and J. Wang, ACS Sens., 2016,
1, 464–482.
55 W. Gao, S. Emaminejad, H. Y. Nyein, S. Challa, K. Chen,
A. Peck, H. M. Fahad, H. Ota, H. Shiraki, D. Kiriya,
D. H. Lien, G. A. Brooks, R. W. Davis and A. Javey, Nature,
2016, 529, 509–514.
56 H. Lee, T. K. Choi, Y. B. Lee, H. R. Cho, R. Ghaffari, L. Wang,
H. J. Choi, T. D. Chung, N. Lu, T. Hyeon, S. H. Choi and
D.-H. Kim, Nat. Nanotechnol.,2016,11,566572.
57 A. Koh, D. Kang, Y. Xue, S. Lee, R. M. Pielak, J. Kim,
T. Hwang, S. Min, A. Banks, P. Bastien, M. C. Manco,
L. Wang, K. R. Ammann, K.-I. Jang, P. Won, S. Han,
R. Ghaffari, U. Paik, M. J. Slepian, G. Balooch, Y. Huang
and J. A. Rogers, Sci. Transl. Med., 2016, 8, 366ra165.
58 J. Wang, Analytical Electrochemisitry: Controlled-Potential
Techniques, John Wiley/Sons, Hoboken, 2006.
59 E. Bakker, P. Bu
¨hlmann and E. Pretsch, Chem. Rev., 1997,
97, 3083–3132.
60 C. McDonagh, C. S. Burke and B. D. MacCraith, Chem. Rev.,
2008, 108, 400–422.
61 S. Coyle, V. F. Curto, F. Benito-Lopez, L. Florea and
D. Diamond, Wearable Sensors, 2014, pp. 65–83.
62 S. Coyle, F. Benito-Lopez, R. Byrne and D. Diamond,
Wearable and Autonomous Biomedical Devices and Systems
for Smart Environment, Lecture Notes in Electrical Engineering,
2010, pp. 177–193.
63 D. Diamond, S. Coyle, S. Scarmagnani and J. Hayes, Chem.
Rev., 2008, 108, 652–679.
64 M. Bariya, H. Y. Y. Nyein and A. Javey, Nature Electron.,
2018, 1, 160–171.
65 K. Mitsubayashi, M. Suzuki, E. Tamiya and I. Karube, Anal.
Chim. Acta, 1994, 289, 27–34.
66 K. Wilke, A. Martin, L. Terstegen and S. S. Biel, Int.
J. Cosmet. Sci., 2007, 29, 169–179.
67 K. Sato, W. Kang, K. Saga and K. Sato, J. Am. Acad.
Dermatol., 1989, 20, 537–563.
68 Z. Sonner, E. Wilder, J. Heikenfeld, G. Kasting, F. Beyette,
D.Swaile,F.Sherman,J.Joyce,J.Hagen,N.Kelley-Loughnane
and R. Naik, Biomicrofluidics, 2015, 9, 031301.
69 M. J. Buono, Exp. Physiol., 1999, 84, 401–404.
70 T. Kamei, T. Tsuda, Y. Mibu, S. Kitagawa, H. Wada,
K. Naitoh and K. Nakashima, Anal. Chim. Acta, 1998, 365,
259–266.
71 R. Keller and J. Sands, FASEB J., 2014, 28, LB719.
72 Y. Y. Al-Tamer, E. A. Hadi and I. E. I. Al-Badrani, Urol. Res.,
1997, 25, 337–340.
73 K. Sato, Rev. Physiol., Biochem. Pharmacol.,1977,79,51131.
74 B. P. O’Sullivan and S. D. Freedman, Lancet, 2009, 373,
1891–1904.
75 W. Jia, A. J. Bandodkar, G. Valde
´s-Ramı
´
rez, J. R. Windmiller,
Z.Yang,J.Ramı
´
rez, G. Chan and J. Wang, Anal. Chem., 2013,
85, 6553–6560.
76 A. J. Bandodkar, V. W. S. Hung, W. Jia, G. Valde
´s-Ramı
´rez,
J. R. Windmiller, A. G. Martinez, J. Ramı
´rez, G. Chan,
K. Kerman and J. Wang, Analyst, 2013, 138, 123–128.
77 J. Kim, W. R. de Araujo, I. A. Samek, A. J. Bandodkar,
W. Jia, B. Brunetti, T. R. L. C. Paixa
˜o and J. Wang, Electro-
chem. Commun., 2015, 51, 41–45.
78 T. Guinovart, A. J. Bandodkar, J. R. Windmiller, F. J. Andrade
and J. Wang, Analyst, 2013, 138, 7031.
79 J. Kim, I. Jeerapan, S. Imani, T. N. Cho, A. Bandodkar,
S. Cinti, P. P. Mercier and J. Wang, ACS Sens., 2016, 1,
1011–1019.
80 A. J. Bandodkar, D. Molinnus, O. Mirza, T. Guinovart,
J. R. Windmiller, G. Valde
´s-Ramı
´
rez, F. J. Andrade, M. J.
Scho
¨ning and J. Wang, Biosens. Bioelectron., 2014, 54,
603–609.
Chem Soc Rev Review Article
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
Chem.Soc.Rev. This journal is ©The Royal Society of Chemistry 201 8
81 D. P. Rose, M. E. Ratterman, D. K. Griffin, L. Hou, N. Kelley-
Loughnane, R. R. Naik, J. A. Hagen, I. Papautsky and J. C.
Heikenfeld, IEEE Trans. Biomed. Eng.,2015,62, 1457–1465.
82 H. Y. Y. Nyein, W. Gao, Z. Shahpar, S. Emaminejad,
S. Challa, K. Chen, H. M. Fahad, L.-C. Tai, H. Ota, R. W.
Davis and A. Javey, ACS Nano, 2016, 10, 7216–7224.
83 W. Gao, H. Y. Y. Nyein, Z. Shahpar, H. M. Fahad, K. Chen,
S. Emaminejad, Y. Gao, L.-C. Tai, H. Ota, E. Wu, J. Bullock,
Y. Zeng, D.-H. Lien and A. Javey, ACS Sens., 2016, 1,
866–874.
84 H.Lee,T.K.Choi,Y.B.Lee,H.R.Cho,R.Ghaari,L.Wang,
H. J. Choi, T. D. Chung, N. Lu, T. Hyeon, S. H. Choi and
D.-H. Kim, Nat. Nanotechnol., 2016, 11, 566–572.
85 S. Nakata, T. Arie, S. Akita and K. Takei, ACS Sens., 2017, 2,
443–448.
86 X. Huang, Y. Liu, K. Chen, W.-J. Shin, C.-J. Lu, G.-W. Kong,
D. Patnaik, S.-H. Lee, J. F. Cortes and J. A. Rogers, Small,
2014, 10, 3083–3090.
87 V. F. Curto, C. Fay, S. Coyle, R. Byrne, C. O’Toole, C. Barry,
S. Hughes, N. Moyna, D. Diamond and F. Benito-Lopez,
Sens. Actuators, B, 2012, 171–172, 1327–1334.
88 S. Emaminejad, W. Gao, E. Wu, Z. A. Davies, H. Y. Y. Nyein,
S. Challa, S. P. Ryan, H. M. Fahad, K. Chen, Z. Shahpar,
S. Talebi, C. Milla, A. Javey and R. W. Davis, Proc. Natl.
Acad. Sci. U. S. A., 2017, 114, 4625–4630.
89 A. Martı
´
n, J. Kim, J. F. Kurniawan, J. R. Sempionatto,
J. R. Moreto, G. Tang, A. S. Campbell, A. Shin, M. Y. Lee,
X. Liu and J. Wang, ACS Sens., 2017, 2, 1860–1868.
90 D. Morris, S. Coyle, Y. Wu, K. T. Lau, G. Wallace and
D. Diamond, Sens. Actuators, B, 2009, 139, 231–236.
91 B. Schazmann, D. Morris, C. Slater, S. Beirne, C. Fay,
R. Reuveny, N. Moyna and D. Diamond, Anal. Methods,
2010, 2, 342–348.
92 S. Coyle, K.-T. Lau, N. Moyna, D. O’Gorman, D. Diamond,
F. Di Francesco, D. Costanzo, P. Salvo, M. G. Trivella,
D. E. De Rossi, N. Taccini, R. Paradiso, J.-A. Porchet,
A. Ridolfi, J. Luprano, C. Chuzel, T. Lanier, F. Revol-Cavalier,
S. Schoumacker, V. Mourier, I. Chartier, R. Convert,
H. De-Moncuit and C. Bini, IEEE Trans. Inf. Technol.
Biomed., 2010, 14, 364–370.
93 A. P. Selvam, S. Muthukumar, V. Kamakoti and S. Prasad,
Sci. Rep., 2016, 6, 23111.
94 D.-H. Choi, J. S. Kim, G. R. Cutting and P. C. Searson, Anal.
Chem., 2016, 88, 12241–12247.
95 Z. Sonner, E. Wilder, T. Gaillard, G. Kasting and
J. Heikenfeld, Lab Chip, 2017, 17, 2550–2560.
96 R. D. Munje, S. Muthukumar and S. Prasad, Sens. Actuators, B,
2017, 238, 482–490.
97 S. Wang, Y. Wu, Y. Gu, T. Li, H. Luo, L.-H. Li, Y. Bai, L. Li,
L. Liu, Y. Cao, H. Ding and T. Zhang, Anal. Chem., 2017, 89,
10224–10231.
98 N. Coppede
`,G.Tarabella,M.Villani,D.Calestani,S.Iannotta
and A. Zappettini, J. Mater. Chem. B, 2014, 2, 5620–5626.
99 E. L. Tur-Garcı
´a, F. Davis, S. D. Collyer, J. L. Holmes,
H. Barr and S. P. J. Higson, Sens. Actuators, B, 2017, 242,
502–510.
100 W. Han, H. He, L. Zhang, C. Dong, H. Zeng, Y. Dai, L. Xing,
Y. Zhang and X. Xue, ACS Appl. Mater. Interfaces, 2017, 9,
29526–29537.
101 J. R. Sempionatto, T. Nakagawa, A. Pavinatto, S. T. Mensah,
S. Imani, P. Mercier and J. Wang, Lab Chip, 2017, 17,
1834–1842.
102 J. B. Sloan and K. Soltani, J. Am. Acad. Dermatol., 1986, 15,
671–684.
103 R. Peng, Z. Sonner, A. Hauke, E. Wilder, J. Kasting,
T.Gaillard,D.Swaille,F.Sherman,X.Mao,J.Hagen,
R. Murdock and J. Heikenfeld, Lab Chip, 2016, 16, 4415–4423.
104 J. Choi, D. Kang, S. Han, S. B. Kim and J. A. Rogers, Adv.
Healthcare Mater., 2017, 6, 1601355.
105 J. Choi, Y. Xue, W. Xia, T. R. Ray, J. T. Reeder, A. J.
Bandodkar, D. Kang, S. Xu, Y. Huang and J. A. Rogers,
Lab Chip, 2017, 17, 2572–2580.
106 I. Jalbert, Exp. Eye Res., 2013, 117, 138–146.
107 F. J. Holly and M. A. Lemp, Surv. Ophthalmol., 1977, 22,
69–87.
108 N. von Thun und Hohenstein-Blaul, S. Funke and F. H.
Grus, Exp. Eye Res., 2013, 117, 126–137.
109 D. K. Sen and G. S. Sarin, Br. J. Ophthalmol., 1980, 64,
693–695.
110 J. T. Baca, D. N. Finegold and S. A. Asher, Ocul. Surf., 2007,
5, 280–293.
111 H. Kudo, T. Sawada, E. Kazawa, H. Yoshida, Y. Iwasaki and
K. Mitsubayashi, Biosens. Bioelectron., 2006, 22, 558–562.
112 S. Iguchi, H. Kudo, T. Saito, M. Ogawa, H. Saito, K. Otsuka,
A. Funakubo and K. Mitsubayashi, Biomed. Microdevices,
2007, 9, 603–609.
113 M. X. Chu, K. Miyajima, D. Takahashi, T. Arakawa, K. Sano,
S. Sawada, H. Kudo, Y. Iwasaki, K. Akiyoshi, M. Mochizuki
and K. Mitsubayashi, Talanta, 2011, 83, 960–965.
114 H. Yao, A. J. Shum, M. Cowan, I. La
¨hdesma
¨ki and B. A.
Parviz, Bios. Bioelectron., 2011, 26, 3290–3296.
115 Y. T. Liao, H. Yao, A. Lingley, B. Parviz and B. P. Otis, IEEE
J. Solid-State Circuits, 2012, 47, 335–344.
116 J. Kim, M. Kim, M.-S. Lee, K. Kim, S. Ji, Y.-T. Kim, J. Park,
K. Na, K.-H. Bae, H. K. Kim, F. Bien, C. Y. Lee and
J.-U. Park, Nat. Commun., 2017, 8, 14997.
117 J. Zhang, W. Hodge, C. Hutnick and X. Wang, J. Diabetes
Sci. Technol., 2011, 5, 166–172.
118 R. Badugu, J. R. Lakowicz and C. D. Geddes, Analyst, 2004,
129, 516–521.
119 J.-L. Ruan, C. Chen, J.-H. Shen, X.-L. Zhao, S.-H. Qian and
Z.-G. Zhu, Polymers, 2017, 9, 125.
120 K. Dumbleton, Contact Lens Anterior Eye, 2002, 25,
137–146.
121 M. A. Javaid, A. S. Ahmed, R. Durand and S. D. Tran, J. Oral
Biol. Craniofacial Res., 2016, 6, 67–76.
122 Y. Goswami, R. Mishra, A. Agrawal and L. Agrawal, J. Dent.
Med. Sci., 2015, 14, 80–87.
123 R. S. P. Malon, S. Sadir, M. Balakrishnan and E. P. Co
´rcoles,
BioMed Res. Int., 2014, 2014,120.
124 U. M. Nater and N. Rohleder, Psychoneuroendocrinology,
2009, 34, 486–496.
Review Article Chem Soc Rev
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
This journal is ©The Royal Society of Chemistry 201 8 Chem.Soc.Rev.
125 A. van Stegeren, N. Rohleder, W. Everaerd and O. T. Wolf,
Psychoneuroendocrinology, 2006, 31, 137–141.
126 M. V. Thoma, C. Kirschbaum, J. M. Wolf and N. Rohleder,
Biol. Psychol., 2012, 91, 342–348.
127 T. F. Oberlander, J. Weinberg, M. Papsdorf, R. Grunau,
S. Misri and A. M. Devlin, Epigenetics, 2008, 3, 97–106.
128 H. M. Burke, M. C. Davis, C. Otte and D. C. Mohr, Psycho-
neuroendocrinology, 2005, 30, 846–856.
129 M. Castro, P. C. L. Elias, C. E. Martinelli Jr., S. R. R.
Antonini, L. Santiago and A. C. Moreira, Braz. J. Med. Biol.
Res., 2000, 33, 1171–1175.
130 I.-H. Kim, C. Kim, K. Seong, M.-H. Hur, H. M. Lim and
M. S. Lee, J. Evidence-Based Complementary Altern. Med.,
2012, 2012, 1–9.
131 K. Fransen, T. Vermoesen, G. Beelaert, J. Menten, V. Hutse,
K. Wouters, T. Platteau and E. Florence, J. Virol. Methods,
2013, 194, 46–51.
132 E. H. El and W. A. Arafa, J. Egypt. Soc. Parasitol., 2004, 34,
1095–1104.
133 L. R. Bigler, C. F. Streckfus, L. Copeland, R. Burns, X. Dai,
M. Kuhn, P. Martin and S. A. Bigler, J. Oral Pathol. Med.,
2002, 31, 421–431.
134 E.-M. Poll, I. Kreitschmann-Andermahr, Y. Langejuergen,
S. Stanzel, J. M. Gilsbach, A. Gressner and E. Yagmur, Clin.
Chim. Acta, 2007, 382, 15–19.
135 M. Navazesh, Ann. N. Y. Acad. Sci., 1993, 694, 72–77.
136 R. C. Stevens, S. D. Soelberg, S. Near and C. E. Furlong,
Anal. Chem., 2008, 80, 6747–6751.
137 W. Yantasee, C. Timchalk, K. K. Weitz, D. A. Moore and
Y. Lin, Talanta, 2005, 67, 617–624.
138 M. S. Mannoor, H. Tao, J. D. Clayton, A. Sengupta,
D. L. Kaplan, R. R. Naik, N. Verma, F. G. Omenetto and
M. C. McAlpine, Nat. Commun., 2012, 3, ncomms1767.
139 J. Kim, G. Valde
´s-Ramı
´rez, A. J. Bandodkar, W. Jia,
A. G. Martinez, J. Ramı
´rez, P. Mercier and J. Wang, Analyst,
2014, 139, 1632–1636.
140 J. Kim, S. Imani, W. R. de Araujo, J. Warchall, G. Valde
´s-
Ramı
´rez, T. R. L. C. Paixa
˜o, P. P. Mercier and J. Wang,
Biosens. Bioelectron., 2015, 74, 1061–1068.
141 T. Arakawa, Y. Kuroki, H. Nitta, P. Chouhan, K. Toma,
S.Sawada,S.Takeuchi,T.Sekita,K.Akiyoshi,S.Minakuchi
and K. Mitsubayashi, Biosens. Bioelectron., 2016, 84, 106–111.
142 R. P. Peters, M. A. van Agtmael, S. A. Danner,
P. H. Savelkoul and C. M. Vandenbroucke-Grauls, Lancet
Infect. Dis., 2004, 4, 751–760.
143 M. Hallek, B. D. Cheson, D. Catovsky, F. Caligaris-Cappio,
G. Dighiero, H. Do
¨hner, P. Hillmen, M. J. Keating,
E. Montserrat, K. R. Rai and T. J. Kipps, Blood, 2008, 111,
5446–5456.
144 A. D. Association, Diabetes Care, 2014, 37, S81–S90.
145 S. K. Garg, S. Schwartz and S. V. Edelman, Diabetes Care,
2004, 27, 734–738.
146 M. S. Boyne, D. M. Silver, J. Kaplan and C. D. Saudek,
Diabetes, 2003, 52, 2790–2794.
147 B. Seok Lee, J.-N. Lee, J.-M. Park, J.-G. Lee, S. Kim, Y.-K.
Cho and C. Ko, Lab Chip, 2009, 9, 1548–1555.
148 B. Seok Lee, Y. Ui Lee, H.-S. Kim, T.-H. Kim, J. Park,
J.-G. Lee, J. Kim, H. Kim, W. Gyo Lee and Y.-K. Cho,
Lab Chip, 2011, 11, 70–78.
149 T. Zheng, N. Pierre-Pierre, X. Yan, Q. Huo, A. J. O.
Almodovar, F. Valerio, I. Rivera-Ramirez, E. Griffith, D. D.
Decker, S. Chen and N. Zhu, ACS Appl. Mater. Interfaces,
2015, 7, 6819–6827.
150 Y. Mendelson, Clin. Chem., 1992, 38, 1601–1607.
151 N. D. Giardino, L. Chan and S. Borson, Appl. Psychophysiol.
Biofeedback, 2004, 29, 121–133.
152 M. Nitzan, A. Romem and R. Koppel, Med. Devices, 2014, 7,
231–239.
153 R. G. Haahr, S. B. Duun, M. H. Toft, B. Belhage, J. Larsen,
K. Birkelund and E. V. Thomsen, IEEE Trans. Biomed.
Circuits Syst., 2012, 6, 45–53.
154 B. Venema, N. Blanik, V. Blazek, H. Gehring, A. Opp and
S. Leonhardt, IEEE Trans. Biomed. Eng., 2012, 59, 2003–2010.
155 C. M. Lochner, Y. Khan, A. Pierre and A. C. Arias, Nat.
Commun., 2014, 5, 5745.
156 T. Yokota, P. Zalar, M. Kaltenbrunner, H. Jinno,
N. Matsuhisa, H. Kitanosako, Y. Tachibana, W. Yukita,
M. Koizumi and T. Someya, Sci. Adv., 2016, 2, e1501856.
157 J. Kim, P. Gutruf, A. M. Chiarelli, S. Y. Heo, K. Cho, Z. Xie,
A. Banks, S. Han, K.-I. Jang, J. W. Lee, K.-T. Lee, X. Feng,
Y. Huang, M. Fabiani, G. Gratton, U. Paik and J. A. Rogers,
Adv. Funct. Mater., 2017, 27, 1604373.
158 J. M. Goldman, M. T. Petterson, R. J. Kopotic and
S. J. Barker, J. Clin. Monit. Comput., 2000, 16, 475–483.
159 R. Yousefi, M. Nourani, S. Ostadabbas and I. Panahi, IEEE
J. Biomed. Health Inform., 2014, 18, 670–681.
160 S. K. Vashist, Anal. Chim. Acta, 2012, 750, 16–27.
161 J. Yadav, A. Rani, V. Singh and B. M. Murari, Proc. Int. Conf.
Signal Process. Integr. Netw., 2014, 591–594.
162 A. Mason, O. Korostynska, J. Louis, L. E. Cordova-Lopez,
B. Abdullah, J. Greene, R. Connell and J. Hopkins, IEEE
Trans. Biomed. Eng., 2017, 1.
163 N. Fogh-Andersen, B. M. Altura, B. T. Altura and
O. Siggaard-Andersen, Clin. Chem., 1995, 41, 1522–1525.
164 A. Sieg, F. Jeanneret, M. Fathi, D. Hochstrasser, S. Rudaz,
J. L. Veuthey, R. H. Guy and M. B. Delgado-Charro, Eur.
J. Pharm. Biopharm., 2009, 72, 226–231.
165 I. T. Degim, S. Ilbasmis, R. Dundaroz and Y. Oguz, Pediatr.
Nephrol., 2003, 18, 1032–1037.
166 B. Leboulanger, R. H. Guy and M. B. Delgado-Charro, Eur.
J. Pharm. Sci., 2004, 22, 427–433.
167 A. J. Bandodkar and J. Wang, Trends Biotechnol., 2014, 32,
363–371.
168 J. M. Yuen, N. C. Shah, J. T. Walsh, M. R. Glucksberg and
R. P. Van Duyne, Anal. Chem., 2010, 82, 8382–8385.
169 G. Rao, P. Glikfeld and R. H. Guy, Pharm. Res., 1993, 10,
1751–1755.
170 B. Leboulanger, R. H. Guy and M. B. Delgado-Charro,
Physiol. Meas., 2004, 25, R35.
171 L. J. De Groot, G. Chrousos, K. Dungan, K. R. Feingold,
A. Grossman, J. M. Hershman, C. Koch, M. Korbonits,
R. McLachlan, M. New, J. Purnell, R. Rebar, F. Singer and
Chem Soc Rev Review Article
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
Chem.Soc.Rev. This journal is ©The Royal Society of Chemistry 201 8
A. Vinik, Endotext, MDText.com, Inc., South Dartmouth
(MA), 2000.
172 S. V. Edelman, Diabetes Technol. Ther., 2001, 3, 283–284.
173 A. J. Bandodkar, W. Jia, C. Yardımcı, X. Wang, J. Ramirez
and J. Wang, Anal. Chem., 2015, 87, 394–398.
174 Y. Chen, S. Lu, S. Zhang, Y. Li, Z. Qu, Y. Chen, B. Lu,
X. Wang and X. Feng, Sci. Adv., 2017, 3, e1701629.
175 H. Yu, D. Li, R. C. Roberts, K. Xu and N. C. Tien,
J. Microelectromech. Syst., 2012, 21, 917–925.
176 Z.Pu,R.Wang,K.Xu,D.LiandH.Yu,IEEE Sens., 2015, 1–4.
177 Z. Pu, C. Zou, R. Wang, X. Lai, H. Yu, K. Xu and D. Li,
Biomicrofluidics, 2016, 10, 011910.
178 M. Venugopal, K. E. Feuvrel, D. Mongin, S. Bambot,
M. Faupel, A. Panangadan, A. Talukder and R. Pidva, IEEE
Sens. J., 2008, 8, 71–80.
179 M. Venugopal, S. K. Arya, G. Chornokur and S. Bhansali,
Sens. Actuators, A, 2011, 172, 154–160.
180 H. P. Lorenz and M. T. Longaker, Surgery, Springer,
New York, NY, 2008, 191–208.
181 K. F. Cutting, Br. J. Community Nurs., 2003, 8, S4–S9.
182 R. O. Niskanen, O. Korkala and H. Pammo, J. Bone Joint
Surg. Br., 1996, 78, 431–433.
183 P. Kujath and A. Michelsen, Dtsch. Arztebl. Int., 2008, 105,
239–248.
184 M. L. Fernandez, Z. Upton, H. Edwards, K. Finlayson and
G. K. Shooter, Int. Wound J., 2012, 9, 139–149.
185 D. W. Edlow and W. H. Sheldon, Proc. Soc. Exp. Biol. Med.,
1971, 137, 1328–1332.
186 B. Greener, A. a. Hughes, N. p. Bannister and J. Douglass,
J. Wound Care, 2005, 14, 59–61.
187 M. Turns, Br. J. Nurs., 2011, S19–S25.
188 F. Gottrup, Am. J. Surg., 2004, 187, S38–S43.
189 T. R. Dargaville, B. L. Farrugia, J. A. Broadbent, S. Pace,
Z. Upton and N. H. Voelcker, Biosens. Bioelectron., 2013, 41,
30–42.
190 S. Pasche, S. Angeloni, R. Ischer, M. Liley, J. Luprano and
G. Voirin, Adv. Sci. Technol., 2008, 57, 80–87.
191 N. Mehmood, A. Hariz, S. Templeton and N. H. Voelcker,
BioMed. Eng. OnLine, 2015, 14, 17.
192 Y. Hattori, L. Falgout, W. Lee, S.-Y. Jung, E. Poon, J. W. Lee,
I. Na, A. Geisler, D. Sadhwani, Y. Zhang, Y. Su, X. Wang,
Z. Liu, J. Xia, H. Cheng, R. C. Webb, A. P. Bonifas, P. Won,
J.-W. Jeong, K.-I. Jang, Y. M. Song, B. Nardone,
M. Nodzenski, J. A. Fan, Y. Huang, D. P. West, A. S.
Paller, M. Alam, W.-H. Yeo and J. A. Rogers, Adv. Healthcare
Mater., 2014, 3, 1597–1607.
193 M. Punjiya, H. Rezaei, M. A. Zeeshan and S. Sonkusale,
Proc. IEEE Int. Conf. Solid-State Sens. Actuators Microsyst.,
2017, 1700–1702.
194 M. Punjiya, P. Mostafalu and S. Sonkusale, Proc. IEEE Int.
Midwest Symp. Circuits Syst., 2017, 495–498.
195 T. Guinovart, G. Valde
´s-Ramı
´rez, J. R. Windmiller, F. J.
Andrade and J. Wang, Electroanalysis, 2014, 26, 1345–1353.
196 P.Kassal,J.Kim,R.Kumar,W.R.deAraujo,I.M.Steinberg,
M. D. Steinberg and J. Wang, Electrochem. Commun., 2015,
56, 6–10.
197 X. Liu and P. B. Lillehoj, Biosens. Bioelectron., 2017, 98,
189–194.
198 X. Liu and P. B. Lillehoj, Proc. Micro Electro Mech. Syst.,
2017, 377–380.
199 P. Mostafalu, W. Lenk, M. R. Dokmeci, B. Ziaie,
A. Khademhosseini and S. R. Sonkusale, IEEE Trans.
Biomed. Circuits Syst., 2015, 9, 670–677.
200 M. F. Farooqui and A. Shamim, Sci. Rep., 2016, 6, 28949.
201 B. Melai, P. Salvo, N. Calisi, L. Moni, A. Bonini, C. Paoletti,
T. Lomonaco, V. Mollica, R. Fuoco and F. Di Francesco,
Proc. IEEE Eng. Med. Biol. Soc., IEEE Int. Conf., 38th, 2016,
pp. 1898–1901.
202 D. A. Jankowska, M. B. Bannwarth, C. Schulenburg,
G. Faccio, K. Maniura-Weber, R. M. Rossi, L. Scherer,
M. Richter and L. F. Boesel, Biosens. Bioelectron., 2017,
87, 312–319.
203 M. K. Nakhleh, H. Amal, R. Jeries, Y. Y. Broza, M. Aboud,
A. Gharra, H. Ivgi, S. Khatib, S. Badarneh, L. Har-Shai,
L. Glass-Marmor, I. Lejbkowicz, A. Miller, S. Badarny,
R. Winer, J. Finberg, S. Cohen-Kaminsky, F. Perros,
D. Montani, B. Girerd, G. Garcia, G. Simonneau,
F. Nakhoul, S. Baram, R. Salim, M. Hakim, M. Gruber,
O. Ronen, T. Marshak, I. Doweck, O. Nativ, Z. Bahouth,
D. Shi, W. Zhang, Q. Hua, Y. Pan, L. Tao, H. Liu, A. Karban,
E. Koifman, T. Rainis, R. Skapars, A. Sivins, G. Ancans,
I. Liepniece-Karele, I. Kikuste, I. Lasina, I. Tolmanis,
D. Johnson, S. Z. Millstone, J. Fulton, J. W. Wells,
L. H. Wilf, M. Humbert, M. Leja, N. Peled and H. Haick,
ACS Nano, 2017, 11, 112–125.
204 T. H. Risby and S. F. Solga, Appl. Phys. B: Lasers Opt., 2006,
85, 421–426.
205 M. J. Maisels and E. Kring, Pediatrics, 2006, 118, 276–279.
206 T. Hibbard and A. J. Killard, Crit. Rev. Anal. Chem., 2011,
41, 21–35.
207 M. G. Persson, L. E. Gustafsson, O. Zetterstro
¨m,
V. Agrenius and E. Ihre, Lancet, 1994, 343, 146–147.
208 C. N. Tassopoulos, D. Barnett and T. Russell Fraser, Lancet,
1969, 293, 1282–1286.
209 J. Hodgkinson and R. P. Tatam, Meas. Sci. Technol., 2013,
24, 012004.
210 K.-H. Kim, S. A. Jahan and E. Kabir, TrAC, Trends Anal.
Chem., 2012, 33, 1–8.
211 J. M. Corres, I. R. Matias, M. Hernaez, J. Bravo and
F. J. Arregui, IEEE Sens. J., 2008, 8, 281–285.
212 M. Righettoni, A. Tricoli, S. Gass, A. Schmid, A. Amann and
S. E. Pratsinis, Anal. Chim. Acta, 2012, 738, 69–75.
213 S. Borini, R. White, D. Wei, M. Astley, S. Haque, E. Spigone,
N. Harris, J. Kivioja and T. Ryha
¨nen, ACS Nano, 2013, 7,
11166–11173.
214 M. C. Caccami, M. Y. S. Mulla, C. D. Natale and G. Marrocco,
Proc. 11th Eur. Conf. Antennas Propag., 2017, pp. 3394–3396.
215 L. Wang, J. A. Jackman, J. Hyeon Park, E.-L. Tan and
N.-J. Cho, J. Mater. Chem. B, 2017, 5, 4019–4024.
216 S. Ammu, V. Dua, S. R. Agnihotra, S. P. Surwade,
A. Phulgirkar, S. Patel and S. K. Manohar, J. Am. Chem.
Soc., 2012, 134, 4553–4556.
Review Article Chem Soc Rev
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
This journal is ©The Royal Society of Chemistry 201 8 Chem.Soc.Rev.
217 C. Di Natale, R. Paolesse, E. Martinelli and R. Capuano,
Anal. Chim. Acta, 2014, 824, 1–17.
218 K. Arshak, E. Moore, G. M. Lyons, J. Harris and S. Clifford,
Sens. Rev., 2004, 24, 181–198.
219 J. W. Gardner, H. W. Shin and E. L. Hines, Sens. Actuators,
B, 2000, 70, 19–24.
220 T. Seesaard, P. Lorwongtragool and T. Kerdcharoen,
Sensors, 2015, 15, 1885–1902.
221 N. Kahn, O. Lavie, M. Paz, Y. Segev and H. Haick, Nano
Lett., 2015, 15, 7023–7028.
222 H. Jin, T.-P. Huynh and H. Haick, Nano Lett., 2016, 16,
4194–4202.
223 P. Lorwongtragool, E. Sowade, N. Watthanawisuth,
R. Baumann and T. Kerdcharoen, Sensors, 2014, 14,
19700–19712.
224 S. Imani, A. J. Bandodkar, A. M. V. Mohan, R. Kumar,
S. Yu, J. Wang and P. P. Mercier, Nat. Commun., 2016,
7, 11650.
225 M. Kaltenbrunner, M. S. White, E. D. Głowacki, T. Sekitani,
T. Someya, N. S. Sariciftci and S. Bauer, Nat. Commun.,
2012, 3, 770.
226 Y.-H. Lee, J.-S. Kim, J. Noh, I. Lee, H. J. Kim, S. Choi, J. Seo,
S. Jeon, T.-S. Kim, J.-Y. Lee and J. W. Choi, Nano Lett., 2013,
13, 5753–5761.
227 X. Pu, L. Li, H. Song, C. Du, Z. Zhao, C. Jiang, G. Cao,
W. Hu and Z. L. Wang, Adv. Mater., 2015, 27, 2472–2478.
228 V. G. H. Eijsink, A. Bjørk, S. Gåseidnes, R. Sirevåg,
B. Synstad, B. van den Burg and G. Vriend, J. Biotechnol.,
2004, 113, 105–120.
229 A. J. Bandodkar, W. Jia, J. Ramı
´rez and J. Wang, Adv.
Healthcare Mater., 2015, 4, 1215–1224.
230 C.Mateo,J.M.Palomo,G.Fernandez-Lorente,J.M.Guisan
and R. Fernandez-Lafuente, Enzyme Microb. Technol., 2007,
40, 1451–1463.
231 A. Pantelopoulos and N. G. Bourbakis, IEEE Trans. Syst.,
Man, Cybern., Syst., Part C: Appl. Rev., 2010, 40, 1–12.
Chem Soc Rev Review Article
Published on 03 April 2018. Downloaded by California Institute of Technology on 03/04/2018 16:48:34.
View Article Online
  • ... Flexible electronics are expected to bring out a revolution in diverse fields of technology, such as electronic skin [1,2], robotics [3,4] or health-monitoring devices [5][6][7], among others. Most of the recent advances in this context have been possible due to the emergence of new conductive and flexible materials, many of which have reported outstanding results in terms of electrical conduction and integration, such as the polycrystalline silicon (poly-Si) [8,9] or several semiconducting metal oxides (e.g., SnO2, TiO2, ZnO or ITO) [10][11][12]. ...
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  • ... Those [4][5][6]. Previous reviews have described current developments in wearable sensor technologies covering aspects on fabrication [7], materials and systems [8], electrophysiological wearables [9], flexible wearables and electronics [10] sensors and electrochemical sensors [11], sweat monitoring [12], wireless and communications [13] as well as energy harvesting developments [14]. Conversely, this review explores advances specific to electrochemical sensors integrated with microneedles which are designed to interface with dermal interstitial fluid (ISF). ...
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  • ... 3 Wearable chemical sensors have recently received tremendous attention towards the monitoring of the wearer's health and fitness 1,2 . Newly developed wearable bioelectronic devices have thus been shown useful for tracking continuously and non-invasively (bio)chemical markers, such as metabolites and electrolytes, in different body fluids 3 . ...
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  • ... Wearable biosensors that interface with the human skin have received much attention because of the popularization of portable electronic consumers, such as flexible medical sensors or various smart watches/bracelets [1][2][3][4][5][6]. Recent advances in materials science, mechanics, and electronics establish the foundations for stretchable and flexible sensors that can conform to the complex, textured surface of the skin. ...
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    Wearable battery-free perspiration analyzing sites based on sweat flowing on ZnO nanoarrays was fabricated.Coupling of hydrovoltaic effect and enzymatic reaction were analyzed.The wearable wireless physiological status monitoring system has potential application in constructing sports big data. Wearable battery-free perspiration analyzing sites based on sweat flowing on ZnO nanoarrays was fabricated. Coupling of hydrovoltaic effect and enzymatic reaction were analyzed. The wearable wireless physiological status monitoring system has potential application in constructing sports big data. We fabricated wearable perspiration analyzing sites for actively monitoring physiological status during exercises without any batteries or other power supply. The device mainly consists of ZnO nanowire (NW) arrays and flexible polydimethylsiloxane substrate. Sweat on the skin can flow into the flow channels of the device through capillary action and flow along the channels to ZnO NWs. The sweat flowing on the NWs (with lactate oxidase modification) can output a DC electrical signal, and the outputting voltage is dependent on the lactate concentration in the sweat as the biosensing signal. ZnO NWs generate electric double layer (EDL) in sweat, which causes a potential difference between the upper and lower ends (hydrovoltaic effect). The product of the enzymatic reaction can adjust the EDL and influence the output. This device can be integrated with wireless transmitter and may have potential application in constructing sports big data. This work promotes the development of next generation of biosensors and expands the scope of self-powered physiological monitoring system.
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    Sweat potentially contains a wealth of physiologically relevant information, but has traditionally been an underutilized resource for non-invasive health monitoring. Recent advances in wearable sweat sensors have overcome many of the historic drawbacks of sweat sensing and such sensors now offer methods of gleaning molecular-level insight into the dynamics of our bodies. Here we review key developments in sweat sensing technology. We highlight the potential value of sweat-based wearable sensors, examine state-of-the-art devices and the requirements of the underlying components, and consider ways to tackle data integrity issues within these systems. We also discuss challenges and opportunities for wearable sweat sensors in the development of personalized healthcare. This Review Article examines the development of wearable sweat sensors, considering the challenges and opportunities for such technology in the context of personalized healthcare.
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    Wearable electronics are emerging as a platform for next-generation, human-friendly, electronic devices. A new class of devices with various functionality and amenability for the human body is essential. These new conceptual devices are likely to be a set of various functional devices such as displays, sensors, batteries, etc., which have quite different working conditions, on or in the human body. In these aspects, electronic textiles seem to be a highly suitable possibility, due to the unique characteristics of textiles such as being light weight and flexible and their inherent warmth and the property to conform. Therefore, e-textiles have evolved into fiber-based electronic apparel or body attachable types in order to foster significant industrialization of the key components with adaptable formats. Although the advances are noteworthy, their electrical performance and device features are still unsatisfactory for consumer level e-textile systems. To solve these issues, innovative structural and material designs, and novel processing technologies have been introduced into e-textile systems. Recently reported and significantly developed functional materials and devices are summarized, including their enhanced optoelectrical and mechanical properties. Furthermore, the remaining challenges are discussed, and effective strategies to facilitate the full realization of e-textile systems are suggested.
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    Despite tremendous recent efforts, non-invasive sweat monitoring is still far from delivering its early analytical promise. Here, we describe a flexible epidermal microfluidic detection platform fabricated through hybridization of lithographic and screen-printed technologies for efficient sweat sampling and continuous real-time electrochemical monitoring of sweat glucose and lactate levels. This soft, skin-mounted device judiciously merges lab-on-a-chip and electrochemical detection technologies, integrated with a miniaturized flexible electronic board for real-time wireless data transmission to a smart device. Modelling of the device design and sweat flow conditions allowed optimization of the sampling process and microchannel layout for achieving rapid filling of the detection reservoir (within 8 min, from starting exercise). The wearable micro-device enabled efficient natural sweat pumping to the electrochemical detection reservoir containing the enzyme-modified electrode transducers. The fabricated device can easily be mounted on the epidermis without hindrance to the wearer and displays resiliency against continuous mechanical deformation expected from such epidermal wear. Amperometric biosensing of lactate and glucose from the rapidly generated sweat, using the corresponding immobilized oxidase enzymes, was wirelessly monitored during cycling activity of different healthy subjects. This ability to monitor sweat glucose levels introduces new possibilities for effective diabetes management, while similar lactate monitoring paves the way for new wearable fitness applications. The new epidermal microfluidic electrochemical detection strategy represents an attractive alternative to recently reported colorimetric sweat-monitoring methods, and hence holds considerable promise for practical fitness or health monitoring applications.
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    Skin is the largest organ of the human body, and it offers a diagnostic interface rich with vital biological signals from the inner organs, blood vessels, muscles, and dermis/epidermis. Soft, flexible, and stretchable electronic devices provide a novel platform to interface with soft tissues for robotic feedback and control, regenerative medicine, and continuous health monitoring. Here, we introduce the term “lab-on-skin” to describe a set of electronic devices that have physical properties, such as thickness, thermal mass, elastic modulus, and water-vapor permeability, which resemble those of the skin. These devices can conformally laminate on the epidermis to mitigate motion artifacts and mismatches in mechanical properties created by conventional, rigid electronics while simultaneously providing accurate, non-invasive, long-term, and continuous health monitoring. Recent advances in the design and fabrication of soft sensors with more advanced capabilities and enhanced reliability suggest an impending translation of these devices from the research lab to clinical environments. Regarding these advances, the first part of this manuscript reviews materials, design strategies, and powering systems used in soft electronics. Next, the paper provides an overview of applications of these devices in cardiology, dermatology, electrophysiology, and sweat diagnostics, with an emphasis on how these systems may replace conventional clinical tools. The review concludes with an outlook on current challenges and opportunities for future research directions in wearable health monitoring.
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    As chemical sensors are in great need for portable and wearable analytical applications, effort is highly desired to develop all-solid-state ion-selective electrode (ISE) and reference electrode (RE) platform that can match simplicity and stability. Here we proposed a wearable sensor platform with a new type of all-solid-state ISE based on a gold nanodendrites (AuNDs) array electrode as the solid contact and a polyvinyl acetate/inorganic salt (PVA/KCl) membrane coated all-solid-state RE. A simple and controllable method was developed to fabricate the AuNDs on a microwell array patterned chip by one-step electrodeposition without additional process. For the first time, the AuNDs electrodes with different real surface area and double layer capacitance were developed as solid contact of the Na+-ISE to investigate the relationship between the performance of the ISE and the value of surface area. The as-prepared AuNDs-ISE with larger surface area (~7.23 cm2) exhibited an enhanced potential stability in comparison with the smaller surface area of AuNDs-ISE (~1.85 cm2) and the bare Au ISE. Important as the ISE, the PVA/KCl membrane coated Ag/AgCl RE exhibited a very stable potential even after 3 months’ storage. Finally, a wearable “Sweatband” sensor platform was developed for efficient sweat collection and real-time analysis of sweat sodium during an indoor exercise. This all-solid-state ISE and RE integrated sensor platform provided a very simple and reliable way to construct diverse portable and wearable devices for healthcare, sports, clinical diagnosis and environmental analysis applications.