Transducer Technologies for Biosensors and Their Wearable Applications

Abstract

The development of new biosensor technologies and their active use as wearable devices have offered mobility and flexibility to conventional western medicine and personal fitness tracking. In the development of biosensors, transducers stand out as the main elements converting the signals sourced from a biological event into a detectable output. Combined with the suitable bio-receptors and the miniaturization of readout electronics, the functionality and design of the transducers play a key role in the construction of wearable devices for personal health control. Ever-growing research and industrial interest in new transducer technologies for point-of-care (POC) and wearable bio-detection have gained tremendous acceleration by the pandemic-induced digital health transformation. In this article, we provide a comprehensive review of transducers for biosensors and their wearable applications that empower users for the active tracking of biomarkers and personal health parameters.

Full text

Citation: Polat, E.O.; Cetin, M.M.;
Tabak, A.F.; Bilget Güven, E.; Uysal,
B.Ö.; Arsan, T.; Kabbani, A.; Hamed,
H.; Gül, S.B. Transducer Technologies
for Biosensors and Their Wearable
Applications. Biosensors 2022,12, 385.
https://doi.org/10.3390/
bios12060385
Received: 28 April 2022
Accepted: 27 May 2022
Published: 2 June 2022
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biosensors
Review
Transducer Technologies for Biosensors and Their Wearable
Applications
Emre Ozan Polat * , M. Mustafa Cetin, Ahmet Fatih Tabak , Ebru Bilget Güven , Bengü Özu˘gur Uysal ,
Taner Arsan , Anas Kabbani , Houmeme Hamed and Sümeyye Berfin Gül
Faculty of Engineering and Natural Sciences, Kadir Has University, Cibali, Istanbul 34083, Turkey;
mustafa.cetin@khas.edu.tr (M.M.C.); ahmetfatih.tabak@khas.edu.tr (A.F.T.); ebru.bguven@khas.edu.tr (E.B.G.);
bozugur@khas.edu.tr (B.Ö.U.); arsan@khas.edu.tr (T.A.); anas.kabbani@stu.khas.edu.tr (A.K.);
houmemehamed@stu.khas.edu.tr (H.H.); 20171709010@stu.khas.edu.tr (S.B.G.)
*Correspondence: emre.polat@khas.edu.tr
Abstract:
The development of new biosensor technologies and their active use as wearable devices
have offered mobility and flexibility to conventional western medicine and personal fitness tracking.
In the development of biosensors, transducers stand out as the main elements converting the signals
sourced from a biological event into a detectable output. Combined with the suitable bio-receptors and
the miniaturization of readout electronics, the functionality and design of the transducers play a key
role in the construction of wearable devices for personal health control. Ever-growing research and
industrial interest in new transducer technologies for point-of-care (POC) and wearable bio-detection
have gained tremendous acceleration by the pandemic-induced digital health transformation. In
this article, we provide a comprehensive review of transducers for biosensors and their wearable
applications that empower users for the active tracking of biomarkers and personal health parameters.
Keywords: transducers; biosensors; wearables; analytes; bioreceptors
1. Introduction
Living in the world of information technologies, we interact with sensors and transduc-
ers daily through smartphones, wearables, cars, cities, homes, and offices. After decades
of development and commercialization phases, sensor-and- transducer-based systems
have been greatly improved and have facilitated a more comfortable daily life. Today’s
commercially available sensors use a wide spectrum of signal conversion mechanisms for
providing instantaneous feedback on personal health, environmental conditions, events,
and changes. With the help of the transducer integrated systems, applications such as
autonomous driving, robotics, and smart homes are already making people’s lives easier,
more comfortable, and safer.
In parallel to the foreseeable progress of sensor and transducer technologies, com-
plementary connection technologies, such as the internet of things (IoT) and 5G, pave
the way toward larger-scale systems, leading to telehealth, smart cities, and autonomous
transportation. Biosensors in health and fitness wearables play a crucial role by extracting
the bio-information required as an input for the abovementioned emerging applications. To
that end, wearables aim to provide low-cost, instant feedback on biomarkers using efficient
signal conversion mechanisms. In the development of wearable biosensors, the signal
conversion mechanism simply defines the functionality and compatibility of operation on
human skin. Therefore, transducers define the wearable form factor and potential user
adoption of a particular device as the main elements of signal conversion. In this review
article, we will focus on the transducer technologies for biosensors and their wearable
applications to present a concise outlook on the prospects and challenges, as well as the
basics of biosensing for a complete understanding of the concept.
Biosensors 2022,12, 385. https://doi.org/10.3390/bios12060385 https://www.mdpi.com/journal/biosensors

Biosensors 2022,12, 385 2 of 32
Starting from the first “biosensor” demonstration by Clark and Lyons in 1962 [
1
], in
which the preparation of an enzyme electrode provided the transformation of glucose
into a detectable current output via oxygen reduction, biosensors took decisive steps in
bio-analysis, such as potentiometric urea electrodes by Guilbault and Montalvo in 1969 [
2
],
and the first commercial biosensor for glucose detection in 1975 [
3
]. Today, biosensors have
increasingly come to define an entire realm of commercial applications that encourage a
proactive approach to preserve one’s health so as to prevent illness in lieu of battling it after
the fact. The impact of biosensors has been promoted with their existence in wearables
that open a venue for research and industry to supply the demand of increasing standards
of living. To that end, the development of clinical-grade biosensors and their wearable
device applications are expected to give predictive guidance for clinical interventions. With
the achievement of clinical-grade accuracy and precision in health and fitness wearables,
personal vital sign extraction might become a prerequisite before requesting an appointment
from a health professional.
Biosensors refer to the collaboration of receptors that recognize target analytes and
transducers that translate this recognition into a detectable signal [
4
]. Biological molecules
such as enzymes, nucleic acids, antibodies, or their synthetic analogues can serve as
bio-receptors to bind the analyte of interest. To form a biosensor device that detects or
measures the biological events or changes, the targeted matching of the bio-receptor and
the analyte should be evaluated quantitatively, making the transducers indispensable
components of a biosensor [
5
,
6
]. Availability of various bio-receptors, transducers, and
possible combinations of both components constitute various ways to classify biosensors
(Figure 1).
Biosensors can monitor target analytes such as biomarkers, pathogens, or allergens.
All can be exploited as an indicator of the health status not only to diagnose but also to
monitor the patient’s prognosis [
7
–
11
]. This track record is favored, especially for elderly
individuals and individuals with alcohol or drug abuse cases [
12
,
13
]. Monitoring the
levels of exogenous substances is a leverage for the healthcare professionals to provide
full-fledged guidance.
The analytes are the biological components of a biosensor that can be extracted from
different bodily fluids such as sweat, saliva, urine, and blood (Figure 1). In addition to
bio-recognition from bodily fluids, the variety and extended functionality of transducing
mechanisms allow for the extraction of vital signs such as basal body temperature (BTT),
heart and respiratory rates, systolic and diastolic pressures, and even tremors. When
combined with the readout electronics, the output signal from the transducers can be traced,
analyzed, and recorded to evaluate the health-related quality of life (HRQOL)
[5,14–17]
. In
today’s world, the maintenance of HRQOL is raising concern in society, which increases
the demand for wearable biosensors monitoring vital processes. To that end, wearables
have brought a completely different perspective to modern medicine from fitness trackers
to medical devices in the clinical setting.
Technologies offering mobility to patients and clinicians authenticated the contribu-
tion of wearable biosensors in remote detection and the monitoring of individuals’ health
status. Perceiving the signs of potential clinical issues in advance has improved the maneu-
verability of health care professionals not only to monitor their self-quarantined patients
but also to increase their preparedness by tracking their physiological status as front-line
workers [18,19].
Determining and recording the physiological parameters and comparing them to the
critical thresholds with absolute precision has gradually become easier over the years with
the foreseeable progress of biotechnology. The bio-analytical systems have radically pro-
gressed from the times when samples were taken from the relevant person and delivered
to different laboratories to the point-of-care (POC) diagnostics, a bedside patient follow-up
unit, which is directly accessible to the person concerned [
5
,
20
]. In line with this progress,
the continuous tracking of body outputs is now provided by wearable biosensors [
5
,
21
].
The breakthrough advantages of wearable technology compared to the conventional bio-

Biosensors 2022,12, 385 3 of 32
analytical methods or POC testing devices are that their continuous monitoring does not
require an invasive way to collect samples from the person of interest and can be performed
in a user-friendly operation at a low cost [
22
–
25
]. The academic and industrial interest in
wearable technologies greatly attracted the development of new mobile devices advancing
biosensors by combining them with new materials and compact electronics. Based on our
current search on PubMed for the articles published with the keywords “wearable” and
“biosensor” in their abstracts, we can deduce that nearly half (47%) of those address the
significance of the predictive and personalized remote advantages. Likewise, wearable
biosensors specific to the diagnosis and prognosis of COVID-19 have strengthened their
use in the market for healthy individuals, as well as for patients during the pandemic [
26
].
According to Researchandmarkets’ report, the global market of wearable technologies holds
a value of USD 47.89 billion with expected growth reaching the value of USD 118.16 bil-
lion in 2028 [
27
]. Owing to advantages such as fast response, specificity, and sensitivity,
commercially developed wearable biosensors for the medical industry have gone through
a radical shift for self-testing at home and critical care at the bedside in emergencies [
28
].
Furthermore, the digital health transformation accelerated by the pandemic strongly de-
pends on wearable biosensor technologies to provide accurate and continuous sensing of
physiological information [29].
Biosensors 2022, 12, x FOR PEER REVIEW 3 of 33
Figure 1. Schematic illustration of biosensing. Analyte containing bodily fluids match with the bio-
receptor which can be an enzyme, antibody, cell, nucleic acid, or biomimetic-based. The matching
of the analyte and bio-receptor creates a change in the signal that is registered and converted to a
measurable output by the transducers. Signal conversion can be in the means of electrochemical,
optical, thermal, or gravimetric. The output signal is processed by integrated or discrete electronics
where the signal can be amplified, filtered, or sent to desired device platforms. For extended func-
tionality and mobility, modern biosensors provide wireless data communication to smart devices
with their integrated data communication module. This way, the extracted signal can be displayed
or recorded on any mobile device for personal health monitoring. (Created with BioRender.com
accessed on 27 April 2022).
Determining and recording the physiological parameters and comparing them to the
critical thresholds with absolute precision has gradually become easier over the years with
the foreseeable progress of biotechnology. The bio-analytical systems have radically pro-
gressed from the times when samples were taken from the relevant person and delivered
to different laboratories to the point-of-care (POC) diagnostics, a bedside patient follow-
up unit, which is directly accessible to the person concerned [5,20]. In line with this pro-
gress, the continuous tracking of body outputs is now provided by wearable biosensors
[5,21]. The breakthrough advantages of wearable technology compared to the conven-
tional bio-analytical methods or POC testing devices are that their continuous monitoring
does not require an invasive way to collect samples from the person of interest and can be
performed in a user-friendly operation at a low cost [22–25]. The academic and industrial
interest in wearable technologies greatly attracted the development of new mobile devices
advancing biosensors by combining them with new materials and compact electronics.
Based on our current search on PubMed for the articles published with the keywords
“wearable” and “biosensor” in their abstracts, we can deduce that nearly half (47%) of
those address the significance of the predictive and personalized remote advantages.
Figure 1.
Schematic illustration of biosensing. Analyte containing bodily fluids match with the
bio-receptor which can be an enzyme, antibody, cell, nucleic acid, or biomimetic-based. The matching
of the analyte and bio-receptor creates a change in the signal that is registered and converted to a
measurable output by the transducers. Signal conversion can be in the means of electrochemical,
optical, thermal, or gravimetric. The output signal is processed by integrated or discrete electronics
where the signal can be amplified, filtered, or sent to desired device platforms. For extended
functionality and mobility, modern biosensors provide wireless data communication to smart devices
with their integrated data communication module. This way, the extracted signal can be displayed
or recorded on any mobile device for personal health monitoring. (Created with BioRender.com
accessed on 27 April 2022).

Biosensors 2022,12, 385 4 of 32
Playing a crucial role in signal conversion for biosensing, transducers provide output
quantity with a given relationship to input quantity [
30
]. The compatibility of biosensing
technologies to the wearable form factors strongly depends on the transducing technologies,
therefore, advancements in the transducers and the miniaturization of readout electronics
with wireless data communication technologies created stark improvements in the field of
wearable biosensing, especially in the consumer-based products that are widely available in
the market. In the next sections, we will focus on the construction of transducers by giving
insight into their signal conversion mechanisms, and provide a detailed classification of
biosensor technologies based on such mechanisms.
2. Construction and Classification of Transducers for Biosensors and Wearables
Biosensors have been evolved as the combination of bio-receptors and transducers
based on electrochemical, optical, thermal, and gravimetric methods converting the signals
from analytes containing antibodies, nucleic acids, and immunological agents, microorgan-
isms, hormones, enzymes, cells, tissues, chemical receptors, and other detectable biological
inputs. For most biosensors, device construction entails three steps; (i) implementation of a
bio-receptor that reacts with a specific analyte, (ii) integration of a transducer, and (iii) the
fixation/immobilization of a biological component to the transducer. Therefore, creating a
biosensing device strongly depends on these construction steps together with device design
strategies and integration of readout electronics for wearable devices that will provide
continuous use on the human body. It is also important to acknowledge that the isolation
from external factors such as chemical/physical conditions (temperature, contaminants,
and pH), should be further taken into consideration while constructing wearable biosensors
for specific applications [31,32].
The variety of available transducing mechanisms provides wearable biosensors worn
on the head, neck, torso, legs, feet, arms, hands, and fingers. Figure 2A shows the body
locations for which the multiple wearable form factors have been reported [
33
–
35
]. The
wearable market and reported research results include a broad spectrum of device designs
from smart helmets to skin patches (Figure 2A), for which user adoption and accuracy are
the critical factors to present a sustainable and user-friendly technology. To that end, the
construction of transducing mechanism is a key factor that can convert many biosensing
technologies to be used as wearable devices.
Transducers for biosensors depend on the type of material used, specifications of
the sensor device, and the actual signal conversion mechanism (Figure 2B). The trans-
ducer materials are generally classified as inorganic, organic, conductor, insulator, and
semiconductor, and can also be found in the form of a biological substance. While the
specifications of the transducer are mostly defined by the capabilities of the active sensing
material, the design of the device also plays an important role in the definition of the final
specifications (Figure 2B). To that end, transducer mechanism simply defines the class of
the biosensors, for instance, a biosensor is classified as an “electrochemical biosensor” if it
uses an electrochemical transducer. Following the conventional classification, from now on,
we will classify the biosensors depending on their transducers.

Biosensors 2022,12, 385 5 of 32
Biosensors 2022, 12, x FOR PEER REVIEW 5 of 33
Figure 2. Wearable biosensors and their transducer specifications. (A) Body locations for which a
variety of wearable form factors are reported [33–35]. The variety of the transducing mechanisms
and the miniaturization of electronics provide wearables that can be worn on the head, neck, torso,
legs, feet, arms, and hands. (B) Transducer specifications for construction of biosensors. Transducers
are the main components in the biosensors that register the biological signal and convert it to a
detectable output that can be in various formats. Active sensing material can be inorganic, organic,
conductor, insulator, semiconductor, or in the form of a biological substance. While the transducer
specifications are defined by the active material, the design of the transducer device can affect the
resulting wearable device specifications such as sensitivity, stability, cost, and variety of the speci-
fications that are given under the transducer specifications (middle panel of (B). Signal conversion
mechanisms are conventionally classified as the means of electrochemical, optical, thermal, and
gravimetric transducing. A more specific classification of the transducers with the signal conversion
mechanisms is given in the panel on the right-hand side of (B). (A) Created with BioRender.com
accessed on 27 April 2022, (B) Adapted from the data shown in ref. [36].
Figure 2.
Wearable biosensors and their transducer specifications. (
A
) Body locations for which a
variety of wearable form factors are reported [
33
–
35
]. The variety of the transducing mechanisms and
the miniaturization of electronics provide wearables that can be worn on the head, neck, torso, legs,
feet, arms, and hands. (
B
) Transducer specifications for construction of biosensors. Transducers are
the main components in the biosensors that register the biological signal and convert it to a detectable
output that can be in various formats. Active sensing material can be inorganic, organic, conductor,
insulator, semiconductor, or in the form of a biological substance. While the transducer specifications
are defined by the active material, the design of the transducer device can affect the resulting
wearable device specifications such as sensitivity, stability, cost, and variety of the specifications that
are given under the transducer specifications (middle panel of (
B
). Signal conversion mechanisms
are conventionally classified as the means of electrochemical, optical, thermal, and gravimetric
transducing. A more specific classification of the transducers with the signal conversion mechanisms
is given in the panel on the right-hand side of (
B
). (
A
) Created with BioRender.com accessed on
27 April 2022, (B) Adapted from the data shown in ref. [36].

Biosensors 2022,12, 385 6 of 32
2.1. Electrochemical Biosensors
Electrochemical biosensors were the first scientifically proposed and successfully
commercialized biosensors for multiple analytes [
1
,
37
–
40
]. Electrochemical biosensors
are constructed with bio-analytes and electrochemical transducers. They utilize chemical
reactions between the immobilized biomolecule and target analyte that produce/consume
ions or electrons that affect the measurable electrical properties (e.g., electric current or
potential) of the solution [
37
,
39
,
41
]. These biosensors are primarily based on Faraday’s laws
of electrolysis and Faradaic current resulting from the direct transfer of electrons in redox
reactions at the heterogeneous electrode-solution interface [
42
]. In such redox reactions,
reference electrodes, such as silver tetramethylbis (benzimidazolium) diiodide [
43
], must be
held still at a fixed potential so that they neither affect the working electrodes nor be affected
by the solution. It is, however, highly important to develop reusable and miniaturized
reference electrodes as the commercial demand has substantially risen for electrochemical
biosensing-based technologies [44].
The reported advantages of electrochemical biosensors include ease of use, better
signal-to-noise ratios with lower background noise, ability to operate on small sample
volumes, cost-effectiveness in production, and compatibility to downscaling together
with the minimal power requirements on the device operation [
40
,
45
,
46
]. On the other
hand, reported limitations are (i) the unreliable dynamic range of measurement due to
enzyme saturation kinetics, (ii) potential interference of other compounds in solutions,
(iii) oxygen requirement and concentration fluctuation in solution, particularly for glucose
measurement, (iv) the influence of pH values or ionic forces on the enzymatic activities,
and (v) biochemical and slow electron transfer processes restricting the efficiency and
speed [40].
Electrochemical biosensors can be further categorized into three main types: am-
perometric/voltammetric, potentiometric/conductometric, and impedimetric/capacitive
biosensors. All these different types of electrochemical biosensors serve in a wide spectrum
of applications ranging from detecting pollutants and pathogens to clinical diagnosis of
various diseases [
47
], such as sensors for glucose and H
2
O
2
monitoring [
48
,
49
], immunosen-
sors [
50
–
52
] for identification of several viruses, such as plum pox [
53
], fig mosaic [
54
], and
avian leukosis subgroup J [
55
]. In addition, agricultural, environmental, and industrial ap-
plications have also been reported by using electrochemically transducing biosensors [
56
].
2.1.1. Amperometric and Voltammetric Biosensors
Amperometric biosensors continuously measure the current
(i)
or the current density
(j=i
A)
(per unit area
A
), that can be generated through oxidation or reduction by a
biochemical reaction at the surface of a working electrode [
40
,
57
–
59
]. The electrode species
can be made up of graphite, noble metals, and modified forms of carbon or conducting
polymers [
60
–
62
]. The simplest form of amperometric biosensing is the direct current (DC)
measurements, which can be well-represented by their ancestor Clark oxygen electrodes [
1
].
Amperometric biosensors that provide linear concentration dependence over a defined
range [39] offer advantages of rapid robustness, portability, high sensitivity, low cost, and
low limit of detection [
46
,
63
]. However, in order to promote the electrochemical reaction of
a selected analyte at the working electrode, suitable analytes must be introduced. Some
analytes, such as protein-based ones, may not be able to serve as good redox partners in
such reactions [
64
]. It is also reported that, due to the uncompensated intrinsic resistance of
the DC measurements, the specificity and selectivity of this technique may be insufficient
for applications requiring high sensitivity [59].
Amperometric transducers have been reported to be implemented in wearable biosen-
sors successively [
65
,
66
]. By using amperometric transducing, Kim et al. have demonstrated
a wearable mouthguard with integrated wireless data communication [
67
]. Figure 3A
shows the demonstrated wearable mouthguard and the placement of the receptor and the
transducer accordingly. In this device, the authors used a wireless amperometric circuit to
detect the uric acid (UA) directly from the saliva. The UA is considered a crucial biomarker

Biosensors 2022,12, 385 7 of 32
for the detection of various diseases such as Lesch–Nyhan and Renal syndrome; therefore,
wearable amperometric transducing with the saliva as an input is a suitable and promising
method for non-invasive and low-cost monitoring of metabolites [
67
]. To that end, the
authors used screen printing to form the working electrodes on PET (polyethylene tereph-
thalate) substrate and fabricated the Prussian-blue transducer (Figure 3B) by crosslinking
the uricase enzyme and electropolymerizing o-phenylenediamine. To test the amperometric
mouthguard, the undiluted human saliva samples were first obtained from volunteers
and the concentration of the saliva samples was determined up to the artificial saliva. The
authors have recorded the changes in the current output with the application of
−
0.3 V
bias voltage (Figure 3C). The data is wirelessly communicated to the final device through
the fabricated printed circuit board (PCB) and the programmable Bluetooth low energy
(BLE) module. The inset of Figure 3C shows the linear calibration data extracted from the
amperometric changes. Figure 3D shows the stability of the demonstrated wearable sensor
through repetitive measurements. Kim et al. have demonstrated that the amperometric
wearable sensor has provided good linearity and stability that holds promise for continuous
monitoring. The inset of Figure 3D shows the percentage changes of the original current
response due to instabilities.
Biosensors 2022, 12, x FOR PEER REVIEW 7 of 33
the transducer accordingly. In this device, the authors used a wireless amperometric cir-
cuit to detect the uric acid (UA) directly from the saliva. The UA is considered a crucial
biomarker for the detection of various diseases such as Lesch–Nyhan and Renal syn-
drome; therefore, wearable amperometric transducing with the saliva as an input is a suit-
able and promising method for non-invasive and low-cost monitoring of metabolites [67].
To that end, the authors used screen printing to form the working electrodes on PET (pol-
yethylene terephthalate) substrate and fabricated the Prussian-blue transducer (Figure 3B)
by crosslinking the uricase enzyme and electropolymerizing o-phenylenediamine. To test
the amperometric mouthguard, the undiluted human saliva samples were first obtained
from volunteers and the concentration of the saliva samples was determined up to the
artificial saliva. The authors have recorded the changes in the current output with the
application of −0.3 V bias voltage (Figure 3C). The data is wirelessly communicated to the
final device through the fabricated printed circuit board (PCB) and the programmable
Bluetooth low energy (BLE) module. The inset of Figure 3C shows the linear calibration
data extracted from the amperometric changes. Figure 3D shows the stability of the
demonstrated wearable sensor through repetitive measurements. Kim et al. have demon-
strated that the amperometric wearable sensor has provided good linearity and stability
that holds promise for continuous monitoring. The inset of Figure 3D shows the percent-
age changes of the original current response due to instabilities.
Figure 3. Wearables using amperometric, potentiometric, impedimetric, and voltammetric trans-
ducer mechanisms. (A) Kim et al.’s smart mouthguard with integrated working electrodes and PCB
module with Bluetooth for wireless data communication [67]. (B) Schematic illustration of the re-
ported bio-receptor yielding amperometric changes due to binding of salivary uric acid to uricase
containing working electrode. (C) Recorded amperometric changes with the increasing concentra-
tions steps of 0.2 mM (A–F). Inset shows the resulting calibration plot with respect to uric acid con-
centration. (D) Stability of measurements with respect to time. Authors have reported the stability
of the amperometric transducers with 20 min. intervals over a total 2 h measurement. Inset shows
the deviations from the original amperometric response at t = 0. (E) Gao et al.’s wearable-integrated
sensor array allows real-time sweat analysis [68]. (F) Schematic representation of the array of sensors
including potentiometric transducing of sodium (Na
+
) and potassium (K
+
) ions, and amperometric
transducing of glucose and lactate from sweat. A resistance-based temperature sensor is also
Figure 3.
Wearables using amperometric, potentiometric, impedimetric, and voltammetric transducer
mechanisms. (
A
) Kim et al.’s smart mouthguard with integrated working electrodes and PCB module
with Bluetooth for wireless data communication [
67
]. (
B
) Schematic illustration of the reported
bio-receptor yielding amperometric changes due to binding of salivary uric acid to uricase containing
working electrode. (
C
) Recorded amperometric changes with the increasing concentrations steps
of 0.2 mM (
A
–
F
). Inset shows the resulting calibration plot with respect to uric acid concentration.
(
D
) Stability of measurements with respect to time. Authors have reported the stability of the
amperometric transducers with 20 min. intervals over a total 2 h measurement. Inset shows the
deviations from the original amperometric response at t = 0. (
E
) Gao et al.’s wearable-integrated
sensor array allows real-time sweat analysis [
68
]. (
F
) Schematic representation of the array of sensors
including potentiometric transducing of sodium (Na
+
) and potassium (K
+
) ions, and amperometric
transducing of glucose and lactate from sweat. A resistance-based temperature sensor is also included

Biosensors 2022,12, 385 8 of 32
in the array and packed together with the PCB containing a programmable microcontroller and a
Bluetooth module for wireless data communication. (
G
) Experimental demonstration of potentio-
metric transducing of Na
+
in NaCl and (
H
) K
+
in KCl solutions with increasing concentration of
ions. (
I
)
Lee et al.’s
transparent and flexible multi-sensor skin patch [
69
]. Demonstrated devices
contain gold (Au) mesh, Au film, and Au-doped graphene as an electrochemically active layer for
multiple transducing mechanisms including amperometry, (
J
) voltammetry, and (
K
,
L
) impedime-
try for the detection of glucose and pH from a limited volume of sweat. (
A
–
D
) Reproduced with
permission from [
67
], Copyright
©
Elsevier 2015, (
E
–
H
) Reproduced with permission from [
68
], Copy-
right
©
Springer Nature 2016, (
I
–
L
) Reproduced with permission from [
69
], Copyright
©
Springer
Nature 2016.
The amperometric biosensing subclass also includes voltammetry, in which the current
is measured during the controlled variations of the potential [
70
]. Promising features of
voltammetry opened significant opportunities in electrochemically transducing biosensing
systems, through voltammetric biosensors. Voltammetry is an electro-analytical method,
from which information is obtained for an analyte by varying potential, and then, the
resulting current can be measured accordingly. There are a variety of forms of voltammetry
sourced from the alternative ways of potential modulation, such as polarography [
71
,
72
],
cyclic voltammetry (CV), normal pulse, reverse pulse, linear sweep, differential staircase
and differential pulse, and square wave [
73
–
75
]. Of these, CV is one of the most widely used
techniques that provides information about redox potential and electrochemical reaction
rates of analyte solutions. The working principle of voltammetric biosensors simply utilizes
the change in current as a function of the varying potential. This principle is required
to detect analytes in solutions, in which the peak current value provides information
for identification, while peak current density is proportional to the concentration of the
corresponding species. The advantages of voltammetric biosensors are high sensitivity in
measurements and simultaneous detection of multiple analytes [
76
]. Considerable logistic
factors and features have been reported, such as the mechanical strength, cost-efficiency,
availability, stability, and easy alteration of structures. The construction of such biosensors
can be made up of various materials, such as blue-dendrimer nanocomposites [
77
], glassy
carbon, polycrystalline boron-doped diamond, carbon nanotubes [
78
], graphene, and
carbon-paste electrodes (CPEs) with multi-walled carbon nanotubes, CPEs with graphite,
or carbon microspheres [79,80].
2.1.2. Potentiometric and Conductometric Biosensors
Biosensors using potentiometric transducers measure the difference in potential (oxi-
dation or reduction potential) generated across an ion-selective membrane separating two
solutions at virtually zero-current flow for ranging analyte activities. The working princi-
ple of potentiometric biosensors relies on a current flow occurring in an electrochemical
reaction, where a ramp voltage is applied to the electrodes in the solution. Potentiomet-
ric transduction operates through a “Boltzmann distribution applied to the redox-active
molecular systems”, namely, the Nernst Potential in a simple form [59,61]:
V=V0+kBT
eln [Ox]
[Red]
where
V
denotes the electrode potential,
V0
denotes the formal potential of redox probes,
Tis absolute temperature,
kB
is the Boltzmann constant, eis the charge of an electron,
and the ratio
[Ox]
[Red]
is the concentration ratio of oxidized and reduced species. Basically,
potentiometric biosensors deal with the potential difference
V−V0
in the above equation
due to a biological event that changes the equilibrium of the redox reaction. Therefore, by
definition, it can be written as a change in the Gibbs free energy [60]:
∆G=eV−V0=kBTln [Ox]
[Red]

Biosensors 2022,12, 385 9 of 32
Potentiometric biosensors offer compatibility with downscaling and modern silicon
fabrication technologies [
81
]. Many of them are commercially available, such as pH elec-
trodes, ion-selective electrodes (ISEs), glass electrodes, and metal oxide-based sensors.
While ion-selective membranes of ions, (H
+
, F
−
, I
−
, Cl
−
), and gases (CO
2
, and NH
3
)
are currently available for biosensing, researchers focus more on creating novel mem-
brane compositions with carbon-based materials, polyvinyl chloride (PVC), and unique
ionophores [
82
] to enhance the capability and uses. Such membranes are reported to be
primarily used in ISEs and ion-sensitive field-effect transistors (ISFETs) [37,39,40].
ISFET technology is also classified under potentiometric biosensors similar to the
ISEs. ISFETs are reliable, well-applicable, analytical devices, and have many advantages in
biosensors, such as suitability for mass production, easy production of small-sized devices
with the semiconductor manufacturing process, robustness, applicability of non-conductive
materials, and fast response. ISFET is a classical metal oxide semiconductor FET with a
gate formed by a separated reference electrode and attached to the gate area via an aqueous
solution [83,84].
ISFETs have an ion-sensitive surface. When any interaction between the semiconduc-
tor and ions occurs, the surface electrical potential changes, which can subsequently be
measured. A selectively permeable polymer layer, from which ions diffuse through and
change surface potential, can also be used to cover the sensor electrode in the construction
of the ISFET (also called Enzyme-FET or ENFET, having very low detection limits but
allowing the use of small sample volumes) [
84
,
85
]. Different types of ISFET biosensors
have been reported, such as ultra-thin body dual-gate ISFET, silicon nanowire ISFET, and
biologically sensitive field-effect transistors. Common drawbacks of the ISFETs have been
reported as high dependence on pH changes and insufficient measurement capabilities for
blood and blood plasma samples with high buffering capacity [86–88].
The incorporation of the potentiometric transducer in wearable technologies has
also been frequently reported [
89
–
91
]. By using a potentiometric transducer mechanism,
Gao et al. demonstrated a fully integrated wearable sensor array for perspiration anal-
ysis [
68
]. As a potentiometric transducer, the authors have used PEDOT:PSS (poly(3,4-
ethylenedioxythiophene) polystyrene sulfonate) in the ISEs; and carbon nanotubes in the
polyvinyl butyral (PVB) reference membrane. This way, they created mechanically robust
potentiometric transducers that can be worn on the subject’s wrist or head for sweat analy-
sis (Figure 3E). The demonstrated wearable array of transducers includes potentiometric
ISEs for Na
+
and K
+
ions and amperometric transducers for glucose and lactate sensors
with a complementary resistance-based temperature sensor (Figure 3F). The integrated
sensor array in this work allows multiple transduction mechanisms so that the simulta-
neous measurement of glucose and lactate is through amperometric transduction, while
the detection of Na
+
and K
+
ions is potentiometric. The authors have demonstrated the
experimental characterization of potentiometric transduction to sense Na
+
and K
+
in NaCl
and KCl solutions, respectively, and further demonstrated the real-time sweat analysis from
a subject wearing a headband device during stationary cycling (Figure 3G,H). To provide
simultaneous real-time measurements of four different analytes from the sweat and the
complementary resistance-based temperature sensing, the authors used a programmable
microcontroller integrated with a Bluetooth module on a PCB for the data processing and
wireless data communication to a mobile phone. As shown in this study, mechanically
flexible wearables and the integration of compact electronics into them enhance potentio-
metric and amperometric transducers in wearable forms yielding real-time detection of
biomarkers in indoor and outdoor activities.
Conductometric transducers for biosensors can also be classified in the subgroup of
potentiometric transducers. Conductometry aims to measure the change in the conductivity
of ionic species with respect to the bio-recognition event. Conductometric biosensors have
been demonstrated to sense a variety of analytes such as glucose urea and arginine [
92
].
Even though they have specific advantages, such as the operation without the need for
reference electrodes, compatibility to operate at low-amplitude alternating voltages (which

Biosensors 2022,12, 385 10 of 32
prevents Faraday processes on electrodes), insensitivity to light, and easy integration
and miniaturization, the wearable biosensing applications of conductometric transducing
methods have been suppressed by the limitations such as higher signal-to-noise ratio (>2%)
causing lower sensitivity, low specificity resulting in the incapability of distinguishing
between simultaneous reactions, and occurrence of polarization in the electrodes of the
double-layer capacitance during the reaction. In addition, the response value of such
biosensors highly depends on medium conditions, such as pH and ionic strength, and
buffer capacity [
93
]. Thus, the selectivity of the conductometric method is presumed to
be low and, consequently, its potential use for wearable different applications encounters
technical difficulties.
2.1.3. Impedimetric and Capacitive Biosensors
Impedimetric biosensors depend on the impedance measurements resulting from the
redox reactions in the analyte/electrode interface due to a biological recognition. This
technique includes the application of a small perturbation bias to sense the change in the
oscillations of the current response. Impedimetric measurements require a definition of
a mathematical transfer function for the electrochemical impedance, which is a complex
function of frequency denoted by
[Z∗(ω)]
. In the simplest form, the complex impedance
function is sourced from the ratio of the perturbed voltage with a frequency of ω,
V(t) = V0+Vpeiωt
to the output current response with a phase difference of ∅,
I(t) = I0+Ipeiωt−∅
yielding,
[Z∗(ω)]=V(t)
I(t)=Z+iZ0
where
Z
and
Z0
are the real and complex parts of the impedance, respectively. The im-
pedimetric biosensors have been reported for the detection of bacteria and whole cells [
94
]
Moreover, impedimetric transducing yields the electrochemical impedance spectroscopy
(EIS) technique, which is a highly used biosensing technology based on scanning over the
perturbation frequency to measure the resulting impedance changes in the interface of
bio-receptor and analyte [
95
]. Impedimetric transducing has also been incorporated into
wearables. Lee et al. have demonstrated a wearable device in the form of a semi-transparent
and flexible skin patch using impedimetric transducers for the active monitoring of glucose
in diabetic patients [
69
]. The demonstrated wearable device consists of multiple sensors for
humidity, glucose, pH, and tremor detection (Figure 3I). To provide a mechanically stable
and optically transparent electrochemical interface, the authors used a serpentine mesh
of gold (Au) and Au-doped graphene, yielding a stable transfer of signals. The device
uses a sweat uptake layer, sensing components, and therapeutic components such as mi-
croneedles, a heater, and a temperature sensor to release the drug above the threshold skin
temperature. The authors demonstrated multiple transducing of their device, including
voltammetric, and impedimetric transducing by using Au film, Au mesh, and Au-coated
graphene (Figure 3J–L). All three electrode structures are tested in phosphate-buffered
saline with Fe(CN)
63−/4−
to provide voltammetric (CV) (Figure 3J), and impedimetric
(Figure 3K,L) responses. The authors further demonstrated the stability of the transducing
mechanisms under the application of mechanical stress and applied the wearable patch to
a diabetic mouse to prove the successful drug release correlated with the sensor operation.
The device is finally reported to be attached to healthy individuals from whom the glucose
and pH measurements are taken. The device shows a statistically high level of correlation
factor (p< 0.001, R
2
= 0.89) to the commercially available glucose assay kit for sweat and
reliability to the blood glucose meter measurement.
Capacitive transducing is another promising method for bio-detection and has been
widely investigated for potential integration into wearables for human health monitoring.

Biosensors 2022,12, 385 11 of 32
In a biosensor that uses capacitive transducing (capacitive biosensor), an analyte binds
by interacting with bio-receptors grafted or immobilized on the electrode surface [
96
].
The capacitive transducers are demonstrated to present compatibility in the detection of
analytes, e.g., hormones or DNA fragments, by registering the signal changes due to the
bio-events modulating the dielectric properties or thickness of the immobilized sensing
layer [
97
]. In this regard, the development of affinity-based capacitive biosensors was first
demonstrated in the 1980s depending on changes in dielectric properties, dimension, shape,
and charge distribution, when an antibody/antigen complex formed on the surface of an
electrode [98].
Analog to the parallel plate capacitors, the active surface of capacitive biosensors serves
as one of the plates of the parallel-plate capacitor which detects changes in capacitance that
can be expressed as:
C=ε·A/d
where
ε
is the dielectric medium permittivity, Ais the area of plates, and dis the distance
between plates. If the capacitance between the plates is needed to vary, one can simply
change the d,
ε
, or A. Preferably, the capacitance can be measured using a bridge circuit,
where the output of transducer impedance is given by:
XC= 1/2πf·C
where Cis the capacitance, and fis the excitation frequency.
In general, affinity-based capacitive biosensors can detect both conductive and non-
conductive target samples, which can operate independently of the readout distance due to
a change in the electric field around it [
99
]. The second plate in the capacitor analogy is the
analyte to be detected for the conductive target samples. On the other hand, to measure non-
conductive target samples, a metallic plate serves as the second plate and the target sample
is the insulator between the parallel plates. Since a capacitive sensor measures the change
in dielectric properties and thickness of the dielectric layer, the precision of the capacitive
biosensor varies according to parameters such as the concentration of charged ions at the
electrode–analyte interface, the distance of the electrode plates, and the content of the
analyte. Capacitive transducers provide label-free detection as they utilize a method based
on measuring changes upon binding of the analyte to a ligand/receptor immobilized on
the electrode surface [
100
]. Measurements can be directly performed in real-time, without
dependence on expensive labels. In this respect, biosensors using capacitive transducers
have novel advantages over labeled biosensors. However, apart from the analyte-receptor
interaction at the electrode interface, any binding resulting from incomplete separation of
the sample can also give an output signal which is difficult to distinguish from the original
signal [
101
]. To minimize measurement errors, sample preparation and sensor design
studies should be carried out meticulously.
The capacitive transducers provide many advantages in biosensors, such as high
sensitivity, low operation power, low loading effect because of high input impedance, and
good frequency response. Capacitive transducers in biosensing have been demonstrated
to be very effective in a variety of applications, including cancer tracking [
102
], bacteria
growth monitoring [
103
], chemical solvent detection, DNA hybridization [
104
], and virus
detection [
105
]. With capacitive transducers, biosensors can detect the presence of a
wide spectrum of substances, regardless of the variety of the target molecules (analytes)
in contact.
Capacitive transducers have been widely investigated for potential integration into
smart wearables for human health monitoring. Just as in the affinity-based capacitive
biosensors, capacitive transducing also forms the basis of flexible pressure and strain sensor
systems yielding electronic skin technology (smart skin), that aims to represent the next
generation of biosensing applications such as monitoring the specific motions of robotic
arms and prosthesis, and human–machine interface [
106
]. Smart skins have been inspired
by the ability of human skin to convert external pressure and strain signals into electrical

Biosensors 2022,12, 385 12 of 32
signals. According to the pressure sensing properties of biological skin, pressure sensing can
mainly be used for low-pressure range (0–10 kPa) pulse and micro-touch applications, while
such capacitive pressure sensors can be used for surface pressure distribution applications
in the high-pressure range (10–100 kPa). As the applied pressure rises to a few kPa, the
sensitivity decreases significantly, which, in turn, limits its practical applications. Due to
the distance between the arterial trees and the skin surface, the amplitude of the pulse-
sourced mechanical waves is lost in the propagation path through the soft tissue, yielding
a weak mechanical on our skin surface. Therefore, the pressure range and sensitivity of the
sensor are crucial for the implementation of smart skin applications [
107
]. In recent years,
significant progress has been made in improving the sensitivity of capacitive pressure
sensors to detect ultra-low-pressure changes [
108
]. In addition, it is pivotal to examine as
many reversible pressure changes as possible and to determine whether the capacitance
change is stable or not [109].
On the way to develop electronic skins, elastomers (e.g., polydimethylsiloxane, silk
fibroin, polyurethane, polyethylene terephthalate, etc.) are used as capacitive transducer
materials acting as flexible matrices to deform the conductive filler network (e.g., carbon
materials, metal nanoparticles, liquid metals, and conductive polymers) that cause changes
in capacitance, conductivity, or resistance [
110
]. They convert external pressure and strain
into electronic output to measure bio-parameters or the interaction of a body part to an
external stimulus such as pressure. The performance of such sensors depends on the
properties of the matrix, the mesh of various materials, and the interactions between them.
Compared to conventional sensors based on rigid semiconductors, metals, and ceramics,
elastomers are advantageous since they exhibit the highest level of strain behavior for
wearable applications. Polydimethylsiloxane (PDMS) microstructures have been used to
increase the sensitivity of pressure sensors; however, the sensitivety value was found to be
less than 1 kPa
−1
even in the low-pressure range [
111
]. On the other hand, it was stated
by Mannsfeld et al. [
112
] that the fabrication of microstructured-PDMS films requires a
challenging 4-step process. As an alternative solution to the low-pressure sensitivity draw-
back, it was proposed to integrate organic thin-film transistors for diversification of the
sensing mechanism. With the inclusion of micro-structured PDMS as a dielectric layer in a
flexible organic thin-film transistor (OTFT)-based pressure sensor, a high sensitivity that is
enough (8.2 kPa
−1
) to allow the use of flexible pressure sensors in mobile health monitoring
was achieved for cardiovascular medicine [
113
]. However, this high sensitivity was only
possible in the very narrow pressure range due to the use of PDMS for the dielectric layer.
In another report, the sensitivity of the pressure sensor formed by the combination of PDMS
elastomer, polymethyl methylacrylamide (PMMA), and silver nanowire was calculated
as 3.8 kPa
−1
[
114
]. Alternately, the pressure sensitivity of a capacitive sensor prepared
with the carbon/silicon structure was found to be in the very high range (
0–700 kPa
) but
it was found to exhibit very low sensitivity (0.025 kPa
−1
) [
115
]. Results of the studies
using capacitive transducing for pressure-sensing capabilities [
108
,
112
–
122
] are interpreted
quantitatively and presented in Table 1for easy comparison. According to the reported
results, high filler content is often required to form a conductive network in the elastomer
matrix to achieve measurable electrical signal amplitudes [
106
]. In contrast, such conduc-
tive composite materials show low breakdown stress limiting the design parameters in
wearable applications.
2.2. Optical Biosensors
The most predominant type of biosensors is optically transducing biosensors (optical
biosensors) owing to a variety of the optical methodologies to transduce the biologically
generated signals. Optical biosensors are widely used and inestimable tools for medi-
cal research, including clinical diagnosis of genetic diseases, drug design, neuroscience,
healthcare monitoring, protein detection, and identification [
123
]. Optical detection is
based on the interaction of the optical field with bio-receptors where the analytes bind
and trigger a biochemical reaction [
124
]. Optical biosensors can be categorized as “label-

Biosensors 2022,12, 385 13 of 32
free”, and “label-based” depending on the targeted binding or simultaneous detection of a
range of biomarkers. Commonly reported optical biosensing techniques are based on spec-
trometry, fiber-optics (FOBs) [
125
], interferometry [
126
], and surface plasmon resonance
(SPR) [
127
,
128
]. Moreover, similar to the conventional optical phenomena, absorption,
fluorescence [
129
], refraction, optical diffraction [
85
], phosphorescence, and Raman scat-
tering are highly incorporated in optical biosensing. With the commonly used optical
methods, one can measure the spatial and temporal properties of the light signal resulting
from a biological event, such as the modulation of amplitude, decay time, polarization,
phase, and/or energy that provide diverse knowledge about the properties of an analyte or
event [130].
Table 1.
Reported sensitivity ranges and pressure sensitivity values for the flexible capacitive sensors.
Material Sensitivity Range (kPa) Pressure Sensitivity (kPa−1) Reference
Ecoflex 0–5 0.601 [108]
PDMS square pyramid microstructure 0–2 0.55 [112]
PDMS (microstructure)/PiI2T-Si 0–8 8.2 [113]
PMMA/PDMS/PVP/Silver 45–500 3.8 [114]
Carbon/Silicon 0–700 0.025 [115]
PDMS porous structure 0–0.33 0.26 [116]
ACC/PAA/Alginate 0–1 0.17 [117]
MAA/DMAPS 0–5 9 [118]
PEDOT:PSS/PDMS/silica 0–10 1 [119]
AgNW-PMMA 0–1 2.76 [120]
Graphene Micropyramid 0–4 7.68 [121]
Au/PET/PDMS micropillar 0–16 0.42 [122]
Optical biosensors empower the user over case-specific application areas with the
advantages of real-time detection, continuous interaction at a low-cost [
127
,
128
], compati-
bility with small volume samples, remote sensing in out-of-reach areas [
123
–
125
], and fast
detection [
131
–
134
]. For instance, FOBs can be adapted for the detection of a variety of
bio-events from the growth of Escherichia coli to the ovary cells of hamsters. In addition,
the SPR technique makes the identification of molecular binding possible by offering high
sensitivity and label-free detection [
128
]. SPR imaging is also used in clinical studies for
screening biomarkers and therapeutic targets [
127
]. On the other hand, evanescent-wave
biosensors reflect great detection sensitivities in a short period of time and are capable
of disease diagnosis with high accuracy due to their ability to assess kinetics and affinity
of interactions by optically monitoring biomolecular interactions concurrently [
135
,
136
].
Immunochromatographic test scripts and lateral flow immunoassays have been present for
a long time in the field of health monitoring, and such scripts and immunoassays have a
wide range of use from home pregnancy test kits to the recently developed SARS-CoV-2
diagnostic kits [136].
In the wearable field, the integration of optical biosensors into compact electron-
ics yielded wearable optoelectronics for the continuous and noninvasive extraction of
vital signs. Advancements in optoelectronic integration provided wearable photoplethys-
mography (PPG), in which a light source sends light to the skin, and modulations of
reflected/transmitted light are measured to record heart rate and related cardiac param-
eters [
34
]. The wearable devices that contain PPG have shown an abrupt increase re-
cently [
137
], and today, almost all fitness trackers commonly make use of this technology
together with electrocardiography (ECG) sensors to provide the most accurate results by
comparison and compensation of both measurement technologies. In PPG, the vessels
absorb the incident light at a known wavelength (commonly at visible and near-IR spec-
tra) and the reflected light reaches the light sensor (photodetector, PD) where the pulse
modulated light intensity is registered as heart rate (HR). Unlike ECG, PPG depends on
the local profusion to extract the HR yielding freedom of location for the measurement
sites on the body. However, the thickness and structure of the skin are important factors

Biosensors 2022,12, 385 14 of 32
in vital sign extraction. Skin conditions may change depending on personal parameters
such as age, sex, or due to a medical condition [
138
]. In principle, human skin provides a
unique interface for wearable devices to extract physiological parameters, however, in a
large number of wearable transducing mechanisms (e.g., electrochemical biosensors) the in-
formation received with the help of biosensors is removed from the surface of the skin [
138
].
For instance, conventional ECG electrodes that are attached to the skin surface sense the
electrical signal sourced from the heart, therefore, the HR measurement can deviate up to a
skin condition factor such as body fat [
139
]. On the other hand, PPG uses a light source of
a certain wavelength (400–1000 nm) that can provide a maximum penetration reaching the
dermis and hypodermis, hence, it can bypass the anatomical factors that interfere with the
measurements [
34
]. In parallel, the effect of skin tone has also been discussed frequently
in the context of optical heart rate wearables [
140
]. Based on the classical skin phototype
classification of Fitzpatrick [
141
], previous experimental studies revealed that the PPG HR
measurements vary up to 15% due to the melanin density difference between dark and
light skins [
140
]. Therefore, the effect of skin tone should be considered a major factor in
the construction of optical transducers for biosensors and wearables.
The PPG signal contains vital information that can be extracted as respiratory rate
(RR) [
142
], blood oxygen saturation (SpO
2
) [
143
] (Figure 4A), blood pressure (BP) [
144
], and
cardiac output [
145
]. Moreover, it was demonstrated to be possible to extract clinical physi-
ological parameters directly from the measured PPG signal such as vascular assessment
and autonomic function [146].
The clinical and technological research on PPG has reached its top speed with the
integration into wearables and consumer electronics. To that end, breakthrough skin
conformable pulse oximeters using material technologies have been reported. Yokota et al.
have demonstrated an ultra-flexible pulse oximeter (Figure 4B) wrapped on a finger [
147
].
Unlike the rigid conventional oximeters, the authors have used flexible polymer light-
emitting diodes (PLEDs) and organic photodetectors (OPDs) that can be laminated on the
skin. By integrating green and red PLEDs together with an OPD, the authors have created
a fully flexible pulse oximeter. The authors performed a PPG with an ultra-flexible pulse
oximeter at two different wavelengths to find the blood oxygen saturation by differentiating
the absorbance percentage of oxygenated and deoxygenated hemoglobin. The reported
device can register a SpO
2
change of 9% (between the oxygenation states of 99% and 90%)
(Figure 4C,D). The authors have reported the stable device operation at ambient conditions
for the long term and the ultra-flexible device structure can withstand a fairly good amount
of applied mechanical stress (300 cycles of stretching for OPDs and 1000 for OPDs bending
down to 100 µm).
Similarly, Kim et al. have demonstrated a wireless epidermal oximeter that can be
laminated on human skin [
148
]. The oximeter includes a red (625 nm, InGaAIP) and
commercially available infrared (IR) LED, a photodiode, and the complementary circuitry
that are strain engineered to the stretchable configuration (Figure 4E,F). To achieve the
required device stability in the optical detection of SpO
2
, Kim et al. have electrically and
physically insulated the metal traces and located them on the neutral plane of the elas-
tomeric substrates [
148
]. The authors have implemented the near field communication
(NFC) technology to wirelessly measure blood pulse oxygenation and used an ultrathin
medical adhesive to bind the devices to the skin. To block the environmental light, the
demonstrated epidermal oximeter uses a black textile and an astable oscillator controls
the current in the LEDs [
148
]. This way, the authors detected the variations in the concen-
tration of oxyhemoglobin (
∆
O
2
Hb) and deoxyhemoglobin (
∆
HHb) by the photodetectors’
responses during red and IR illuminations. The demonstrated wireless oximeter devices
were simultaneously operated with a commercial NIRS (Near IR spectroscopy) oximeter
and variations in the
∆
O
2
Hb and
∆
HHb were recorded from the adjacent regions of the
forearm (Figure 4G,H). A good agreement in the oxygenation results, mechanically flexible
form factor, and wireless data and power communication promise for future use of the skin
conformable devices as a gold standard in the wearable field.

Biosensors 2022,12, 385 15 of 32
Another commonly used wearable optical transducing method is colorimetry. Col-
orimetric transducers have the key enabling properties to develop wearables, such as
easy detection and user-friendly feedback by the simple color-changing feedback mech-
anism. Colorimetric wearables have been commonly demonstrated to be laminated on
the skin surface and give the user necessary optical feedback by the color change of the
active transducers [
149
,
150
]. Choe et al. have demonstrated a wearable colorimetric patch
based on a thermoresponsive plasmonic microgel embedded in a stretchable hydrogel
film [
151
]. The authors have demonstrated plasmonic microgel film that undergoes large
and reversible color shifts with respect to temperature changes without the change in the
overall sensor volume (Figure 4I). To control the colorimetric response, the authors fabri-
cated raspberry-shaped plasmonic microgels by decorating them with gold nanoparticles
(AuNPs) on thermoresponsive poly(N-isopropylacrylamide) (PNIPAM) hydrogels. To
provide the repetitive colorimetric response under the successive heating/cooling cycles
of the colloidal solution, the authors have incorporated the plasmonic microgels into the
flexible polyacrylamide (PAAm) hydrogel film and realized reliable thermoresponsive color
shifts. The thermoresponsive plasmonic structure is encapsulated in two PDMS (poly-
dimethylsiloxane) films with thicknesses of 150
µ
m to form the mechanically flexible device
structure to operate on human skin (Figure 4J). This way the wearable colorimetric patches
undergo an efficient peak shift of 176 nm (545 nm to 721 nm), leading to a high-contrast
colorimetric response, and devices are reported to exhibit a stable operation after 10 cycles
of heating and cooling (Figure 4K,L).
Biosensors 2022, 12, x FOR PEER REVIEW 15 of 33
medical adhesive to bind the devices to the skin. To block the environmental light, the
demonstrated epidermal oximeter uses a black textile and an astable oscillator controls
the current in the LEDs [148]. This way, the authors detected the variations in the concen-
tration of oxyhemoglobin (ΔO
2
Hb) and deoxyhemoglobin (ΔHHb) by the photodetectors’
responses during red and IR illuminations. The demonstrated wireless oximeter devices
were simultaneously operated with a commercial NIRS (Near IR spectroscopy) oximeter
and variations in the ΔO
2
Hb and ΔHHb were recorded from the adjacent regions of the
forearm (Figure 4G,H). A good agreement in the oxygenation results, mechanically flexi-
ble form factor, and wireless data and power communication promise for future use of
the skin conformable devices as a gold standard in the wearable field.
Another commonly used wearable optical transducing method is colorimetry. Col-
orimetric transducers have the key enabling properties to develop wearables, such as easy
detection and user-friendly feedback by the simple color-changing feedback mechanism.
Colorimetric wearables have been commonly demonstrated to be laminated on the skin
surface and give the user necessary optical feedback by the color change of the active
transducers [149,150]. Choe et al. have demonstrated a wearable colorimetric patch based
on a thermoresponsive plasmonic microgel embedded in a stretchable hydrogel film [151].
The authors have demonstrated plasmonic microgel film that undergoes large and re-
versible color shifts with respect to temperature changes without the change in the overall
sensor volume (Figure 4I). To control the colorimetric response, the authors fabricated
raspberry-shaped plasmonic microgels by decorating them with gold nanoparticles
(AuNPs) on thermoresponsive poly(N-isopropylacrylamide) (PNIPAM) hydrogels. To
provide the repetitive colorimetric response under the successive heating/cooling cycles
of the colloidal solution, the authors have incorporated the plasmonic microgels into the
flexible polyacrylamide (PAAm) hydrogel film and realized reliable thermoresponsive
color shifts. The thermoresponsive plasmonic structure is encapsulated in two PDMS (pol-
ydimethylsiloxane) films with thicknesses of 150 µm to form the mechanically flexible
device structure to operate on human skin (Figure 4J). This way the wearable colorimetric
patches undergo an efficient peak shift of 176 nm (545 nm to 721 nm), leading to a high-
contrast colorimetric response, and devices are reported to exhibit a stable operation after
10 cycles of heating and cooling (Figure 4K,L).
Figure 4.
Wearables using optical transducers. (
A
) Schematic illustration of the reflective pulse oxime-
try from finger [
147
]. OPD records the reflected light intensity that is illuminated synchronously by
the green and red PLEDs to extract the PPG and the resulting SpO
2
. (
B
) Photograph of
Yokota et al.’s
ultra-flexible polymer oximeter wrapped on the finger [
147
]. (
C
) Demonstrated SpO
2
extraction at
green and red wavelengths by the change of the PPG signal intensity [
147
]. (
D
) A 9% change in the
SpO
2
value (from 99% to 90%) yields a detectable intensity change in the resulting PPG signal at
both wavelengths [
147
]. (
E
) Epidermal wireless pulse oximeter demonstrated by Kim et al. [
148
].
The device is skin conformable by the medical adhesive and the strain engineered device structure.
(
F
) Components of the wireless epidermal oximeter [
148
]. The device includes NFC technology to
wirelessly detect PPG and SpO
2
. All device components are encapsulated and located on the neutral
plane of the elastomeric substrate, yielding the mechanically robust device operation on the human
skin. (G,H) Demonstrated simultaneous measurements from the commercially available NIRS bulk

Biosensors 2022,12, 385 16 of 32
oximeter and the epidermal wireless oximeter [
148
]. Although the two systems use different con-
figurations to optically extract the hemoglobin concentration, the resulting curves exhibit similar
trends and values. (
I
) Choe et al.’s colorimetric thermoresponsive wearable [
151
]. The incorporated
plasmonic microgel structure undergoes an efficient spectral shift during the heating and cooling pro-
cesses, yielding colorimetric feedback to the user about the skin temperature. (
J
) The demonstrated
plasmonic microgels are embedded in PDMS films to provide a device structure with enhanced
mechanical stability [
151
]. (
K
,
L
) Recorded spectral shifts during 10 cycles of heating/cooling [
151
].
The wearable colorimetric patches exhibit a reversible color change with a 176 nm peak shift from
545 nm to 721 nm yielding high-contrast colorimetric feedback to the user. (
A
–
D
) Reproduced with
permission from [
147
] under the terms of the Creative Commons Attribution-NonCommercial license,
Copyright
©
2016, American Association for the Advancement of Science. (
E
–
H
) Reproduced with
permission from [
148
] under the terms of the Creative Commons Attribution-Noncommercial license,
Copyright
©
2016, American Association for the Advancement of Science. (
I
–
L
) Reproduced with
permission from [
151
] under the terms of Creative Commons Attribution 4.0 International License,
Copyright © 2018 Springer Nature.
2.3. Thermal/Calorimetric/Thermometric Biosensors
Biosensors using thermal/calorimetric/thermometric transducers measure energy
changes (heat, q) of a system and its surroundings. Such biosensors are constructed by
immobilization of bio-elements onto temperature sensors, in which the bio-recognition
sourced energy change of a system and its surroundings (i.e., heat exchange) is deter-
mined [
152
]. The working principle of such biosensors involves: (i) the entrance of analyte
solution through the substrate inlet and the measurement of temperature by a thermal trans-
ducer (e.g., a thermistor [
153
,
154
] microelectromechanical system, thermocouple, resonator,
or thermophile [
155
]), (ii) the flow of solution through a packed column consisting of immo-
bilized enzymes where enzyme-catalyzed reactions occur, resulting in the loss or generation
of heat, and (iii) the remeasurement of temperature by a separate thermal transducer while
solution proceeds towards outlet [
156
]. This way, the thermal transducers quantify the
change of heat that occurs within endothermic and exothermic enzyme-catalyzed chemical
reactions in biological systems to interpret findings with respect to the analyte concen-
tration, molar enthalpy, and product formation or the total number of molecules in such
reactions. To detect the temperature change from first principles, we consider the total heat
in the system as:
q=−np(∆H)
and,
q=−Cp(∆T),
yielding a temperature change:
∆T=−(∆H)np/Cp
where qis the total heat, n
p
is the number of moles of the product,
∆
Tis the change in
temperature recorded by the enzyme thermistor (ET), and
∆
His the molar enthalpy change.
Cpis the heat capacity of the system including the solvent [157].
Thermal transducing is also commonly used in resistive devices to reflect the tempera-
ture change as a change of resistance. The change in temperature with respect to resistance
can be described by the Steinhart–Hart equation:
1/T=A+B(ln R) + C(ln R)3

Biosensors 2022,12, 385 17 of 32
where, A,B, and Care the experimentally derived coefficients. From this, the relation
between the resistance at temperature T, (
RT
) and the resistance at T= 0, (
RT0
) can be
written as:
RT=RT0eβ(1
T−1
T0)
for narrow temperature ranges where
β
is a material constant (ranging between 4000 and
5000 K) [158].
Starting with the conventional ET devices containing immobilized enzyme columns,
this class of biosensor has been progressed into the micro and multi-sensing (hybrid) ver-
sions of such devices together with the commercialization [
159
–
161
]. Thermal transducers
are highly suitable for a wide range of applications in the detection of bioprocesses with
a groundbreaking development of thermometric enzyme-linked immunosorbent assay
(TELISA) combining the fundamentals of ELISA [
154
,
162
] for the determination of hor-
mones, antibodies, and other biomolecules generated during the fermentation process,
environmental monitoring [163], clinical diagnosis [164], and food analysis [165].
Based on the abovementioned literature, the current progress promises the devel-
opments aiming to design microdevices with multi-analytical capabilities in addition to
portability and digital data reporting. For instance, periodic and conventional measurement
of body temperature with a thermometer can help regarding early diagnosis of malfunc-
tioning metabolic processes of the body and symptoms of illnesses such as fever, infection,
depression, or even insomnia problems. Real-time and continuous monitoring of such mea-
surements is accurately possible with wearable temperature sensing systems by gathering
diagnostic information and understanding signals of underlying diseases. Lightweight and
flexible temperature sensing devices can be attached to the human skin without the aware-
ness of the user and provide continuous monitoring. In the market of wearables, Yono’s
earbud [
166
], Ava Science’s wristband [
167
], Empatica’s watch [
168
], Oura’s ring [
169
],
and VitalConnect’s chest patch [
166
] are great examples of functioning as a thermometer,
temperature sensors, a skin temperature measurand, and a thermo-resistor, respectively.
2.4. Gravimetric/Piezoelectric/Mass-Sensitive Biosensors
Piezoelectric biosensors are formed by the coupling of a bio-component with piezoelec-
tric transducers that are usually based on a quartz crystal coated with gold electrodes [
170
].
Piezoelectricity is a reversible process in which an electrical charge accumulation is induced
due to the applied mechanical stress causing deformation or vibration of the material [
171
].
Piezoelectric biosensors work upon the detection of frequency changes occurring on the
transducer surface. When the target analyte attaches to the material, the resulting mass
shift on the crystal component causes resonance frequency alterations. To that end, phys-
ical, chemical, or biological microcantilevers detecting changes in cantilever bending or
vibrational frequency can be categorized as gravimetric/mass-sensitive biosensors.
In the construction of piezoelectric biosensors, the end application plays a highly
important role, since the piezoelectric materials and designs differ for specific purposes
such as label-free detecting [
172
] bacteria or virus [
173
], or cancer biomarkers [
174
]. There-
fore, the selection of the mass-sensitive transducer elements includes a variety of factors
ranging from the type and the thickness of electrodes (e.g., gold, chromium, platinum,
titanium, etc.) to the detected bio-agents (e.g., warfare agent, virus, bacteria, etc.). Owing
to various construction options and possible features, there is a high demand for piezo-
electric transducers and their biosensing applications. While the precision and sensitivity
are the advantageous specifications, the common downside of mass-sensitive transducers
is the temperature-dependent sensitivity yielding the loss of such features at extremely
low or high temperatures. In that sense, any change in the temperature causes thermal
inconsistency of various features (e.g., dielectric, piezoelectric, or electromechanical) and
weakening of acoustic waves and dielectric losses. All these factors should be considered
while constructing a piezoelectric biosensor to maintain the effectiveness of the sensor at
various temperatures [175].

Biosensors 2022,12, 385 18 of 32
With the demonstration of in situ interfacial mass detection by piezoelectric transduc-
ers [
176
], piezoelectric/gravimetric/mass-sensitive biosensors have been widely used in
the medical field as immunosensors for the detection of bacteria and viruses. Moreover,
piezoelectric genosensors for the detection of DNA or RNA fragments regarding their
distinct sequence of bases have been reported [
177
]. Important applications of this class
of transducers include cell and tissue characterization, healthcare monitoring, pressure
sensing, and detection of endotoxins [
178
,
179
] cholesterol [
180
,
181
], pesticides [
182
], and
breast cancer [
183
,
184
]. Moreover, piezoelectric biosensors show potential for applications
in food quality detection, and environmental and clinical analysis [185].
Piezoelectric transducers have found a wide spectrum of wearable device applica-
tions. Han et al. have demonstrated a self-powered electronic skin based on a piezo-
biosensing [
23
]. The authors have implemented enzyme/ZnO nanoarrays as the basis of
the piezo-biosensing unit and demonstrated an electronic skin that can detect lactate, glu-
cose, uric acid, and urea in the perspiration [
185
]. Very recently, Su et al. have demonstrated
muscle fibers inspired piezoelectric textiles for wearable physiological monitoring [
186
]. To
mimic the muscle fibers, the authors have used the surface modification of polydopamine
(PDA) and monitored the real-time heart rate and dynamic output profile for the voice
recognition with the skin attachable wearable. Alternately, piezoelectric transducers are
frequently used as energy harvesters in stretchable and self-powered wearables and im-
plantable devices [
187
]. Dagdeviren et al. have reported biocompatible piezoelectric
energy harvesters based on the piezoelectricity of ZnO material [
188
]. Similarly, Zhu et al.
have reported piezoelectric nanogenerators that are mechanically flexible for wearable
applications [
189
]. With its key properties, nanomaterial-based piezoelectricity presents a
compatible platform as transducers for wearable biosensing and energy harvesting. This
multi-discipline approach may lead to a new developing orientation of health and fitness
wearables by promoting such flexible self-powered multifunctional nano-systems.
3. Supplementary Technologies for Wearable Biosensing
In this section, we provide a general outlook on the supplementary technologies em-
powering the wearable use of transducers for biosensing. We summarize and highlight
the enabling supplementary technologies for wearable transducers to represent a user-
friendly and low-cost platform by providing key factors, such as ease of access, continuous
and non-invasive analysis, mechanical robustness, enhanced user adoption, and usability.
Furthermore, we extend our perspective to the energy sources, the data communication
technologies, the location and position services, and the biocompatibility within the supple-
mentary technologies providing a framework for the development of wearable biosensors.
3.1. Microfluidics and Biomedical Microelectromechanical Systems (Bio-MEMS)
Microfluidics provides an efficient platform for transducers by forming the core of the
sensing elements and providing advanced liquid holding and storage capabilities [
190
].
Microfluidics-based point of care medical sensors are of ever-growing interest due to
their versatility as wearable lab-on-the-body systems [
191
–
193
]. This section is dedicated
to a brief outlook on the importance and varieties of the conformal biomedical micro-
electromechanical (bio-MEMS) sensors as a supplementary technology to the wearable
transducer-based sweat analysis. Sweat is a compatible substance with wearable medical
sensor systems owing to its transparent and hypotonic properties [
191
,
194
]. Furthermore,
natural sweating induces high enough pressure (~70 kPa) to sustain pressure different
for the microfluidic channels [
191
,
194
,
195
]. However, sweat-based sensing is prone to
false-concentration-based analyte readings due to vaporization [
191
,
194
,
196
,
197
] and inho-
mogeneous local sweat gland density [
195
]. It has been demonstrated that the measurement
error could be as high as 114% [
192
]. Furthermore, the sensory system should include
microfluidics, electronics, and a power supply to collate and relay data on the collected
and analyzed sweat in real-time, which results in a rather bulky system. Moreover, the
said system should be conformal if desired to be employed in direct contact with the skin

Biosensors 2022,12, 385 19 of 32
on different parts of the body. Despite these shortcomings and constraints, analysis of the
constituents of sweat is a crucial part of medical diagnosis with an ever-growing focus
owing to the variety of critical biomarkers that can be detected in tiny volumes.
Biomarkers present in the sweat that can be detected by wearable microfluidics-based
medical sensors are (i) glucose, (ii) lactate, (iii) pH, (iv) chloride, (v) creatinine, (vi) tyrosine,
(vii) uric acid, (viii) potassium, (ix) sodium, (x) ascorbic acid, (xi) cortisol, (xii) dopamine,
and (xiii) adrenaline. The concentration of glucose in sweat is 1% of that of the plasma and
its accurate measurement is very important for continuous monitoring of diabetes [
193
,
198
].
On the other hand, lactate is an indicator of physical effort signaling the transition from
aerobic to anaerobic metabolism [
194
,
199
]. Determining the level of pH helps in ascertaining
the neuromuscular condition of the subject [
194
], while chloride levels are used to help with
the diagnosis of cystic fibrosis [
194
,
195
]. The level of creatinine is an important indicator in
confirming the hydration status of the metabolism and renal health [
195
] and determining
tyrosine level from sweat is generally used for the diagnosis of liver diseases and various
psychiatric disorders [
197
]. Thus, the analytes found in sweat serve as crucial markers
for the performance of vital organs and critical systems in the human body; therefore,
using sweat as a working fluid in a wearable Bio-MEMS sensor is crucial. The reported
wearable microfluidic devices that use sweat to analyze the abovementioned analytes are
summarized in Table 2.
Table 2. Biomarkers and driving mechanisms in wearable microfluidic Bio-MEMS *.
Analyte Ranges Limit of Detection Sensitivity Literature
Glucose 0–400 µM 1.5–7 µM1.08–3.5 mA mM−1cm−2[191,192,195,198,200]
Lactate 0–100 mM 0.2–2 mM 36.2 µAµM−1cm−2[191,192,194,195,199]
pH 4–8.5 – 71.4 mV pH−1[191,194,196,201]
Chloride 0–625 mM 5–39 mM – [191,195,196]
Creatinine 0–1000 µM 15.6 µM – [195]
Tyrosine 0–160 µM 3.6 µM0.61 µAµM−1cm−2[197]
Uric Acid 0–140 µM 0.74 µM3.50 µAµM−1cm−2[196]
Potassium 0.1–100 mM – – [194,195]
Sodium 0.2–200 mM – 56 mV dec−1[194,196,201]
Ascorbic Acid 0.02–10 mM 0.013–10 µM0.78 ×105C mol−1[193,200]
Cortisol/Cortisol-BSA 5–100 ng/mL – – [200]
1–8 mg/mL – –
Dopamine 1–100 µM 0.05–1 µM1.1 ×105C mol−1[193]
Adrenaline 10–500 µM 2–10 µM0.8 ×105C mol−1[193]
Microfluidic Drive Reported Pressure Values Literature
Pressure of Sweat Glands 70–72 kPa [191–202]
Capillary Force (pressure difference) 100–400 Pa [191,194,195,198]
Active Valves (thermo-responsive hydrogels) 15–300 mmHg [192]
Passive Valves (bursting valves) Laplace-Young Equation [200]
* Data collated solely from the work reported exclusively in the selected references.
In the wearable microfluidic wearable Bio-MEMS, the sweat is collected and regulated
via the natural pressure difference of sweat glands [
191
,
192
,
194
–
202
]. The capillary force
is usually strong enough to drive the fluid through microchannels leading to the cham-
ber where various sensor types are employed [
191
,
195
,
198
]. The flow regulation can be
managed via active thermo-responsive hydrogel valves controlled by microheaters [
192
],
or passive capillary pressure-bursting valves [
200
]. Capillary force arises due to the hy-
drophilicity of the wetted surface and allows liquids to travel long distances without large
pressure difference when the characteristic length-scale of the ducts is in microns [
195
,
198
].
The microchannel itself will exert a shear resistance on the flow, which will be balanced
with the capillary force, and a static pressure build-up will occur when there is no flow.
This pressure might be released in terms of sudden hydrodynamic pressure by changing
the shape, and thus the contact angle, along the way. A bursting valve operates on the very
same principle [
200
]. On the other hand, an active valve, such as the thermo-responsive
system [
192
], does not require plugging the duct entirely but rather partially obstructs
the duct, inducing substantial head loss on the flow thus depriving it of the mechanical
energy necessary to overcome the shear resistance. The reported works using these main

Biosensors 2022,12, 385 20 of 32
driving mechanisms for the control of sweat flow in wearable Bio-MEMS are summarized
in Table 2.
3.1.1. Energy Sources and Detection Mechanisms
Although passive effects such as the capillary force or secondary ducts do not require
a power source, active valves, as well as sensory electronics along with powering the
detection mechanism and achieving real-time or ad-hoc relay of data, require a power
supply. Power supply in wearable Bio-MEMS sensors is based on either energy scavenging
or electrochemical conversion aside from the fact that microfluidics relies on natural pres-
sure difference of sweat [
194
,
195
] except for thermo-responsive valves that are polymers
actuated via a change in temperature [
200
]. The power source exploited in such systems
should be of low voltage and current for the safety and practicality of the system. Also,
the source should be either easily replicable or rechargeable, avoiding the requirement for
complicated reassembly procedures. To that end, the most practical common solution is to
employ batteries to generate electric currents through electrochemical conversion through
battery packs. On the other hand, the electronics can be fitted with energy scavenging
MEMS hardware using electromagnetic fields of a certain frequency, such as the one emitted
by a smartphone [
200
], or the body movement of the subject wearing the system trans-
forming mechanical vibration to electricity by means of induction and triboelectric effect
during physical exercise [
201
]. Furthermore, the external energy might only be required
for near field communication (NFC) to pair the wearable Bio-MEMS sensor and the device
employed to collect the data [
195
]. Furthermore, smartphones can incorporate the wearable
system in a vast IoT framework [
191
]. The reported microfluidic wearables using various
energy sources and the utilized detection mechanisms are summarized in Table 3.
Table 3. The reported energy sources and utilized detection mechanisms in microfluidic wearables.
Energy Source Literature
Electrochemical Conversion (Rechargeable Battery Pack) [192,194,196,197]
Energy Scavenging (Radio Frequency) [200]
Energy Scavenging (Mechanical Motion) [201]
Natural Pressure Difference (~70 kPa) [191,192,194–202]
Detection Mechanism Literature
Colorimetric (Fluorescent and visible light) [191,195,200]
Strain (Swelling) [202]
Galvanic (Capacitance) [196]
Electrochemical (Mediator molecules) [192,193,196,197,199]
3.1.2. Data Transmission
Another important aspect for wearables to represent a mobile technology is their data
transmission. To that end, a wearable system would usually exploit wireless means of com-
munication. The wireless connection usually suffers from a limited onboard power source
with wearable systems; therefore, long rage real-time radio frequency (RF) communication
is not feasible, whereas short-range data transfer options are widely preferred. There are
currently three main methods in the data communication of wearables, namely, wireless, i.e.,
Bluetooth [
192
,
194
,
196
,
197
,
201
], NFC [
195
,
200
], and visual data capture [
191
,
195
,
200
]. Blue-
tooth communication can be real-time while NFC and visual data recording are designed
to be intermittent. While visual data recording does not require either special electronics
or power sources dedicated to the task, wireless and NFC need custom circuitry, an RF
antenna, and a device to be paired and collate the data of interest [
192
,
194
–
197
,
200
,
201
].
In addition, active flow control can be achieved by wireless communication via a mobile
device [
192
]. The presence of integrated electronics increases the device in size to a relative
bulk, however, still practical to apply on the skin.
The transmitted sensory data are of different origins. There are simplistic systems
using NFC to pair the sensor with a device such as a smartphone to start an application

Biosensors 2022,12, 385 21 of 32
for image processing out of colorimetric data [
195
]. Colorimetric data is obtained by the
reaction of analytes with certain chemicals and enzymes that emit certain frequencies if
excited by fluorescent or visible light [
191
,
195
,
200
] that will reveal the spatial concentration.
These systems do not necessarily require real-time data transfer and additional onboard
power dedicated to the colorimetric analysis [
195
]. In addition to colorimetric sensors,
there are systems with hybrid sensing compartments or completely different transducing
approaches, e.g., galvanic, strain, piezoresistive effect, and electrochemical. The reported
devices with various transducing mechanisms are summarized in Table 3.
Electrochemical biosensors are exclusively microfluidic and rely on custom electrodes
coated with specific mediator molecules, such as carbon nanomaterials, to choose and
react to a targeted analyte [
196
,
197
,
199
]. The induced current by the transducer is directly
related to the flux of the molecules of interest in the vicinity of the electrode [
192
]. Thus, the
transportation and diffusion of analytes in the microfluidic network and reservoirs are of
utmost importance [
197
]. Also, strain sensors could exploit swelling of bulk material, e.g.,
hydrogels, as sweat gets absorbed while stretching a conductive fabric changing its re-
sistance [
202
]. Similarly, the sweat rate can be detected based on the change in galvanic
properties of specifically designed microchannel geometries.
3.1.3. Biocompatibility
In microfluidic wearables, the sensors and antennas, along with a printed electrical
interface to the necessary PCB, are expected to be implemented on biocompatible conformal
structures to seamlessly fit on any surface. Furthermore, the sensor is supposed to be at-
tached to the skin via medical-grade adhesives. Sweat itself is supposed to be collected and
stored by the carefully tailored network of microfluidic channels and chambers embedded
in this elastic structure. Such a network can be implemented in PDMS [
196
,
198
,
199
,
201
],
silicone rubber [194], and medical adhesives [197].
Polyimide (PI) [
197
,
198
], PDMS [
193
], and polyethylene terephthalate (PET) [
197
]
are demonstrated as a cover to the microchannels, as well as for the printing of the sen-
sory electrodes. Moreover, different porous materials could be used to collect sweat,
such as hydrogel [
202
] and Ecoflex [
198
], in certain volumes. Likewise, paper and cot-
ton can be used for sweat collection, fluid flow compartmentalization, and sensor place-
ment [
191
,
193
]. In addition, the thermo-responsive valves can be manufactured out of
poly(N-isopropylacrylamide) [192].
To reach the desired wearable microfluidic device components, microchannels and
reservoirs can be implemented via different fabrication techniques such as photolithogra-
phy, screen printing, laser engraving, laying cotton fibers, and embossment. Table 4presents
the list of the reported wearable fabrication techniques and biocompatible microfluidic
device materials. The said techniques are utilized to manufacture tailored microchannels
and electrodes meeting different production constraints. For instance, laser engraving is uti-
lized as a faster and cheaper solution as it reduces the need for infrastructure and expertise
required for photolithography [
195
–
197
,
199
,
201
,
203
]. Also, making use of easy-to-obtain
materials such as cotton or paper when feasible greatly simplifies the manufacturing pro-
cedure and reduces the overall cost [
191
,
193
]. Likewise, the techniques of screen-printing
and embossment are employed as a cost-effective solution to microchannel manufacturing
while preserving the dexterity to obtain a channel network with bifurcations [
194
–
197
].
Nevertheless, manufacturing electrodes for electrochemical or galvanic sensing or building
the associated interfacial electronics entails the use of photolithography [195,196,199,203].
Overall, the end product is biocompatible and flexible so that the wearable biosensor
will perform without exhibiting any chemical reaction to the skin or the sweat. To that end,
it is important to acknowledge that the flow rates in these microfluidic networks are usually
on the order of 0.1–1
µ
L/min and they are designed to accommodate fluidic conditions of
natural sweating.

Biosensors 2022,12, 385 22 of 32
Table 4.
Biocompatible materials and fabrication techniques for microfluidic-based wearable biosensors.
Materials Literature
PDMS [196,198,199,201]
Silicone Rubber [194]
Medical Adhesive [197]
Polyimide [197,198]
Polyethylene Terephthalate [197]
Hydrogel [202]
Ecoflex [198]
Paper [191]
Cotton [193]
Poly(N-isopropylacrylamide) [192]
Fabrication Technique Literature
Photolithography [195,196,199,203]
Screen printing [194]
Laser Engraving [197,201]
Laying Cotton Fibers [191]
Embossment [195]
3.2. Location and Position Services
Currently, available wearable devices are mostly equipped with location/position
services for various purposes such as finding a route, counting the steps, or calculating
exercise output to give extended feedback to the user. Wearables are more efficient for
location services than smartphones since they are bonded to the user in various form
factors. This section provides a general outlook on current advancements in the location
and position systems as a supplementary technology for wearables.
The global positioning system (GPS) used in outdoor positioning is a satellite-based
navigation system developed by the United States Department of Defense [204]. Outdoor
position information can be calculated when four or more GPS satellites are in the line of
sight [
204
]. The GPS enables critical outdoor positioning, predominantly for civil, com-
mercial, and military applications. To that extent, Assisted-GPS (A-GPS) technology is a
positioning system used in mobile devices to find the user’s location within the A-GPS
address server via the base station [
205
]. The accuracy provided by GPS in outdoorposition-
ing is between 3 m and 15 m, while usually around 10 m for indoor positioning
[204,205]
.
On the other hand, the accuracy of the A-GPS in outdoor positioning is 15 m while lim-
ited to 50 m for indoors [
205
]. Although these accuracy values allow sufficient position
determination outdoors, since people spend more than 80% of their time indoors, suitable
indoor geolocation systems are in high demand. Unfortunately, the use of GPS satellites in
indoor localization is not sufficient enough with the weakening and loss of GPS signals
due to the atmospheric delays, signal reflections (multipath), steel structures, roofs, and
building walls [
204
,
205
]. Therefore, in the last two decades, a serious amount of work has
been carried out on developing new technologies that enable a reliable indoor positioning
with high accuracy and low average error. When the GPS signals cannot be reached, the
positioning problem is solved by using different technologies such as infrared, ultrasonic,
cellular, radio frequency recognition (RFID), wireless network (Wi-Fi), Bluetooth beacon, or
ultra-wideband (UWB) sensors [
206
–
209
]. In some studies, even visible light [
210
,
211
] and
technologies that use the earth’s magnetic field have been used [
212
,
213
]. However, new
algorithms and methods are needed to improve the results.
Among the new indoor positioning methodologies, ultra-wideband (UWB) sensor
technology steps forward with its ideal indoor distance estimation, indoor geolocation,
indoor tracking, and navigation [
214
]. The UWB sensor technology is important in pro-
viding centimeter-scale positioning accuracy indoors. UWB is a radio technology used in
short-range high-bandwidth communication. UWB has higher bandwidth of 500 MHz;
therefore, signals usually reach the receiver in more than one way [
215
]. However, high

Biosensors 2022,12, 385 23 of 32
bandwidth allows a range of frequencies to be used at different times; thus, UWB can be
used to solve multipath problems and interference effects. UWB transmitters consume rela-
tively low power compared to other indoor geolocation technologies, making them a more
efficient option by providing a longer battery life for wearables. The power consumption
of wearable UWB transmitters is generally less than 1 mW, while the power consumption
of UWB receivers is around 400 mW [216].
The UWB frequency range for communication applications is between 3.1 and 10.6 GHz [
217
].
This frequency range makes UWB signals less affected by disturbances and prevents them
from being affected by Wi-Fi and Bluetooth signals, mainly operating in the 2.4 GHz
frequency band. In indoor positioning, where the number of people and objects is high,
the field of view (LOS-line of sight) may be relatively blocked, which may cause delay and
deviation in the received signal. It is important to note that the error rate increases if the
user to be positioned is out of sight, but it is difficult to conclude that the absorbing effects
of the human body will increase these errors [
218
,
219
]. Time of flight (ToF) and time of
arrival (ToA) methods are used to determine indoor location with UWB sensors.
UWB sensors with the time difference of arrival method (TDoA) draw attention in
medical applications [
220
,
221
]. Another technique, the angle of arrival (AoA), allows the
signals to be compared with the signal strength of the angles received from at least two
sources. In this way, the object’s location can be found from the angle of the intersection of
the signals [
222
]. With the ToF and TDoA methods, indoor positioning can be performed
with an average error of around 20 cm [
223
]. Since UWB technology requires a special
transmitter and receiver infrastructure, it has not yet entered the informatics market, except
for a few industrial applications [
224
]. However, the current advancements and planning
for the integration of UWB sensors in mobile phones show that UWB sensors will be used
in a wide variety of areas in the very near future, including wearables [225,226].
4. Conclusions, Discussion, and Outlook
In this review, we have provided a detailed outlook on the transducers for biosensors
and their wearable application, including their construction, classification, and wearable
uses. We have also summarized the recent progress in wearable biosensors and the role of
transducers in the development of enabling applications ranging from personal health and
fitness tracking to clinical and environmental uses.
Transducers and their integration into biosensors and wearables have paved the way
toward a radical shift in the main trends of consumer electronics and the application of
western medicine. To that end, the development of wearable biosensor-based telehealth
systems may provide real-time control of personal health parameters without physically
visiting a healthcare unit.
Design and implementation of new material-based technologies allow transducers
that are not possible by the conventional methodologies. Together with the recent progress
in supplementary technologies, such as microfluidics, wireless data communication, and
location/position services, wearable biosensors present viable solutions in a diverse ap-
plication spectrum from healthcare-focused devices to enzyme electrodes, providing the
transformation of glucose into a detectable current output via oxygen reduction, bn which
they prepare, organic [
137
], and piezoelectric polymers [
186
] as active transducers and
the formation of electronic components on unconventional substrates such as paper [227],
textile [
228
] and silk [
229
], together with strain engineering of device components [
148
],
hold utmost importance for the realization of the wearable technologies with enhanced
feel and form factors. To that end, we have provided a material- and design-based per-
spective on the transducers to be used in biosensors and their wearable applications. On
the other hand, Si microprocessor technologies are not expected to be outperformed by a
nanomaterial soon; however, current proof of concept demonstrations on two-dimensional
material technologies provided computational opportunities with high performance and
mechanical flexibility [230].

Biosensors 2022,12, 385 24 of 32
In line with the current progress of biosensors, future wearable applications are
expected to monitor multiple vital biomarkers continuously and non-invasively. Clinical-
grade biomarker extraction is not mandatory in the context of consumer wearables; how-
ever, the implementation of new materials and designs hold great promise for telehealth
through the incorporation of wearables in clinical studies. Forming the basis of wear-
able technologies, transducers for biosensing may serve as a key factor to overcome the
current drawbacks of public health strategies. To that end, we believe that our review
and perspectives on the biosensors with highlighted breakthrough works provide a com-
plete framework and guide the readers in the construction of biosensing devices and their
wearable applications.
Author Contributions:
E.O.P. conceptualized, supervised, and majorly contributed to the writing
of the whole manuscript. E.B.G. majorly contributed to the writing and reviewing of Section 1;
M.M.C. majorly contributed to the writing and reviewing of Section 2; B.Ö.U. majorly contributed to
the writing and reviewing of Section 2.1.3; A.F.T. and T.A. majorly contributed to the writing and
reviewing of Section 3. A.K. organized the manuscript and partially contributed to the writing. H.H.
and S.B.G. helped in the organization of the article and partially contributed to the writing. All
authors have read and agreed to the published version of the manuscript.
Funding:
This work is funded by BAGEP 2021 Physics awards of Bilim Akademisi (E.O.P.) and 2022
Personal Research Funds of Kadir Has University (E.O.P., M.M.C., A.F.T., E.B.G., B.Ö.U. and T.A.).
Institutional Review Board Statement: Not applicable.
Informed Consent Statement: Not applicable.
Acknowledgments:
E.O.P. thanks Nihat Berker and Sondan Durukano˘glu Feyiz for their support in
the interdisciplinary collaboration and for providing personal research funds for open access research.
E.O.P and A.F.T. thank KHAS 2021 Summer School Student Selin Sönmez; M.M.C. thanks KHAS 2021
Summer School Students: Elif Serap Gurler, Zeynep Eylul Yagcıo ˘glu, Lila Kayıran, Alperen Gür, Baha
Eren Ertekin, Aslı Tavaslı, and Selen Melis Battal for their contribution in the literature categorization.
Conflicts of Interest: The authors declare no conflict of interest.
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