Flexible and Stretchable Bioelectronics

Abstract

Medical science technology has improved tremendously over the decades with the invention of robotic surgery, gene editing, immune therapy, etc. However, scientists are now recognizing the significance of ‘biological circuits’ i.e., bodily innate electrical systems for the healthy functioning of the body or for any disease conditions. Therefore, the current trend in the medical field is to understand the role of these biological circuits and exploit their advantages for therapeutic purposes. Bioelectronics, devised with these aims, work by resetting, stimulating, or blocking the electrical pathways. Bioelectronics are also used to monitor the biological cues to assess the homeostasis of the body. In a way, they bridge the gap between drug-based interventions and medical devices. With this in mind, scientists are now working towards developing flexible and stretchable miniaturized bioelectronics that can easily conform to the tissue topology, are non-toxic, elicit no immune reaction, and address the issues that drugs are unable to solve. Since the bioelectronic devices that come in contact with the body or body organs need to establish an unobstructed interface with the respective site, it is crucial that those bioelectronics are not only flexible but also stretchable for constant monitoring of the biological signals. Understanding the challenges of fabricating soft stretchable devices, we review several flexible and stretchable materials used as substrate, stretchable electrical conduits and encapsulation, design modifications for stretchability, fabrication techniques, methods of signal transmission and monitoring, and the power sources for these stretchable bioelectronics. Ultimately, these bioelectronic devices can be used for wide range of applications from skin bioelectronics and biosensing devices, to neural implants for diagnostic or therapeutic purposes.

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Citation: Chitrakar, C.; Hedrick, E.;
Adegoke, L.; Ecker, M. Flexible and
Stretchable Bioelectronics. Materials
2022,15, 1664. https://doi.org/
10.3390/ma15051664
Academic Editors: Zhou Li and
Bojing Shi
Received: 20 December 2021
Accepted: 8 February 2022
Published: 23 February 2022
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materials
Review
Flexible and Stretchable Bioelectronics
Chandani Chitrakar, Eric Hedrick, Lauren Adegoke and Melanie Ecker *
Department of Biomedical Engineering, University of North Texas, Denton, TX 76203, USA;
chandanichitrakar@my.unt.edu (C.C.); erichedrick@my.unt.edu (E.H.); laurenadegoke@my.unt.edu (L.A.)
*Correspondence: melanie.ecker@unt.edu
Abstract:
Medical science technology has improved tremendously over the decades with the inven-
tion of robotic surgery, gene editing, immune therapy, etc. However, scientists are now recognizing
the significance of ‘biological circuits’ i.e., bodily innate electrical systems for the healthy functioning
of the body or for any disease conditions. Therefore, the current trend in the medical field is to
understand the role of these biological circuits and exploit their advantages for therapeutic purposes.
Bioelectronics, devised with these aims, work by resetting, stimulating, or blocking the electrical
pathways. Bioelectronics are also used to monitor the biological cues to assess the homeostasis of the
body. In a way, they bridge the gap between drug-based interventions and medical devices. With
this in mind, scientists are now working towards developing flexible and stretchable miniaturized
bioelectronics that can easily conform to the tissue topology, are non-toxic, elicit no immune reaction,
and address the issues that drugs are unable to solve. Since the bioelectronic devices that come in
contact with the body or body organs need to establish an unobstructed interface with the respective
site, it is crucial that those bioelectronics are not only flexible but also stretchable for constant moni-
toring of the biological signals. Understanding the challenges of fabricating soft stretchable devices,
we review several flexible and stretchable materials used as substrate, stretchable electrical conduits
and encapsulation, design modifications for stretchability, fabrication techniques, methods of signal
transmission and monitoring, and the power sources for these stretchable bioelectronics. Ultimately,
these bioelectronic devices can be used for wide range of applications from skin bioelectronics and
biosensing devices, to neural implants for diagnostic or therapeutic purposes.
Keywords:
flexible and stretchable bioelectronics; stretchable polymer; conductive polymers; stretch-
able sensors; flexible and stretchable power sources; stretchable batteries; supercapacitors; fabrication
of stretchable bioelectronics
1. Introduction
The field of bioelectronics, which goes by the name of neuromodulation, bio-stimulation,
electroceuticals, wearables, implantables, etc., is an emerging field either as an alternative
or as an add-on to chemical and biologic drugs [
1
]. With pacemakers and deep-brain
stimulators, the idea of bioelectronics is not new [
1
]. Bioelectronics can be miniaturized
to micro and nano scales with the help of Micro-Electro-Mechanical-Systems (MEMS) and
nanotechnology [
2
], without compromising the efficacy of the devices. They can also be
designed to elicit a low level of inflammatory response. Although a great improvement
in miniaturized, precise, and accurate bioelectronic devices has been achieved, the next
generation of bioelectronics (implantable or wearable) will require incorporating mechan-
ical flexibility and stretchability into the device for the improved mechanical compliance
between the tissue and the implanted device, and to minimize the foreign body response
that limits the lifetime of these devices [3].
With the understanding that the human body functions and communicates through
biophysical cues such as electrical, thermal, mechanical, and topographic signals [
4
], tremen-
dous advances in developing tools have been achieved to sense and acquire those physi-
ological signals for diagnostic purposes or to introduce physical stimuli for preventative
Materials 2022,15, 1664. https://doi.org/10.3390/ma15051664 https://www.mdpi.com/journal/materials

Materials 2022,15, 1664 2 of 36
and therapeutic purposes. Those tools/devices utilize substrates or an encapsulation layer,
semiconductors as a functional interface with the biological material, and a power sup-
ply [
5
]. Conventional metal and silicon-based devices are bulky and rigid [
6
]. They are
not applicable as wearable devices and as implantable devices. On the contrary, those
inorganic materials offer the advantages of having high charge carrier mobilities and help
in signal transduction mechanisms at the bio-interface, thereby enabling a fast response
with high sensitivity [
4
]. Hence, significant interest lies in an emerging field of flexible and
stretchable bioelectronics to integrate with the soft and complicated surface topography
of the biological tissue. Flexible/stretchable bioelectronic devices are defined as those
devices that can bend and undergo mechanical deformation with the ability to conform
to biological tissue while maintaining electrical integrity under deformation [
7
]. In addi-
tion, the flexible/stretchable bioelectronic devices should be stable, biodegradable, and
biocompatible. Several measures have been explored to add the flexibility and stretchability
properties to biomedical devices, ranging from using elastomeric polymers as substrates
and organic materials as electrodes, to providing a specific stretchable architecture to the
inorganic conductor materials [
8
]. Engineering flexible and stretchable bioelectronics has a
broad field of biomedical applications such as monitoring electrophysiological (electroen-
cephalogram (EEG), electrocardiogram (ECG)), physiological (temperature, heart rate, etc.),
mechanical (strain, blood pressure, etc.), and biochemical (glucose level, enzyme level, etc.)
signals [
9
]. In this paper, we provide a complete review from the fabrication perspective by
summarizing the (i) materials for substrates and stretchable electrodes; (ii) structural design
and fabrication methods to adapt to large deformations without considerable damage to
the device and biological structure; (iii) methods of signal transduction and communica-
tion between the stretchable device and recording device; and (iv) power sources for the
operation of the fabricated stretchable devices.
2. Opportunities and Limitations in Bioelectronic Devices
Medical science has progressed tremendously in recent years. The previous gold
standards in medical science were therapeutic drugs for the treatment of diseases or
disorders, and imaging, blood testing, biopsies, etc. for diagnostics. With the increase
in research in gene therapy, gene silencing via CRISPR, and siRNA, gene editing has
shifted the theme of medical science from the earlier gold standards to these technologies.
However, the bioelectronic approaches for delivering therapeutic benefits are gradually
being recognized because the electrical systems of the body contribute to the healthy
functioning of the body. The potential of bioelectronics to be used in bodily conditions
controlled by biological circuits is one of the aspects that make it a competitor to the
conventional method of using drugs. Another advantage is that these bioelectronic devices
can facilitate precise treatment, meaning the therapeutic effect can be delivered locally only
to the part that needs the treatment. In contrast, drugs affect the whole body, which often
raises the issue of harmful side effects. Bioelectronics can also be utilized to realize smart
drug delivery systems for localized targeted delivery to reduce the side effects of the drugs.
Bioelectronics have also been implemented to stimulate nerves and modulate the neuronal
signals for treating diseases or disorders such as drug-resistant epilepsy [
10
], rheumatoid
arthritis [11], treatment-resistant depression [12], and neuropathic pain [13].
Bioelectronic devices are devices that use electricity to regulate biological processes,
treat diseases, or restore lost functionality [
14
]. They can be classified into wearable or
implantable devices depending on their placement. Skin mountable bioelectronics devices
have gained significant traction due to their non-invasiveness in continual monitoring
of physiological signals or for minimally-invasive localized drug delivery. Some of the
examples of skin mountable bioelectronics include: (i) skin-mountable strain or pressure
sensors for monitoring various human motions and physiological parameters [
15
]; (ii) elec-
trophysiological sensors for electromyography (EMG) and electroencephalography (ECG)
for the diagnosis of diseases such as epilepsy, dysphasia, and dementia; (iii) electrochemical
and microneedle-based biosensors that allow noninvasive and continuous monitoring of

Materials 2022,15, 1664 3 of 36
biomarkers [
16
,
17
]; and (iv) smart theranostic drug delivery systems, such as smart contact
lenses for diabetic diagnosis and therapy [18].
Bioelectronic devices also include devices that work on the principle of neuromodu-
lation, which modify neural signals via deep-brain stimulation, vagus nerve stimulation,
spinal nerve stimulation, and retinal and auditory implants [
19
]. Koopman et al. re-
ported that vagus nerve stimulation (VNS) inhibits tumor necrosis factor, an inflammatory
molecule that is a major therapeutic target in rheumatoid arthritis (RA), and which atten-
uates disease severity [
20
]. FDA Investigational Device Exemption (IDE) for the novel
bioelectronic device to treat RA has allowed the device to be used in a clinical trial to collect
safety and effectiveness data [
11
]. VNS has been established as a treatment method for
drug-resistant epilepsy to reduce the number, length, and severity of seizures. Similarly,
VNS has shown its efficacy for treatment-resistant depression [
21
]. Deep-brain stimulation
has become a routine treatment for disorders such as Parkinson’s disease, tremor, epilepsy,
and obsessive-compulsive disorder for patients whose symptoms are not treated with
medications. Recently, a pioneering “brain-to-spine” wireless implant, also known as a
stimulation movement overground (STIMO) device, has enabled paraplegic patients to
walk with the help of crutches or a walker [
22
]. All these stimulation treatment methods
utilize a pacemaker-like device that is implanted on the chest or abdomen area. A wire con-
nected to the device is then routed to the electrodes placed in areas of brain, nerve, or spinal
cord. The implanted electrodes produce electrical impulses that modulate the abnormal
impulses to alleviate the symptoms associated with the disorders. Stimulation parameters
such as frequency, amplitude, pulse width, and duration are adjusted to maximize the
therapeutic effect while minimizing the side effects for facilitating individualized neuro-
modulation therapy [
23
]. Retinal implants (Argus II Retinal prosthesis system, Second
Sight medical Technology) have enabled patients to regain some lost vision. These devices
consist of a receiver, transmitting coil, and an electrode array that stimulate the retina to
induce vision in patients with retinal degeneration [
24
]. Auditory implant types, such
as cochlear implants and auditory brainstem implants (ABIs), have a neuromodulatory
function [
19
]. Cochlear implants stimulate auditory nerve fibers via electrodes implanted
in the cochlea to help patients with sensorineural hearing loss using functional auditory
nerve fibers. ABIs stimulate the hearing pathways in the brain stem, bypassing the inner
ear and auditory nerve to help patients with missing or small auditory nerves [
25
]. One of
the breakthroughs in the field of bioelectronics is the development of the first bioelectronic
implantable device that stimulates nerve regeneration by the application of electrical pulses.
Following the healing of the nerve, the device disintegrates and disappears without any
harmful side effects [26].
Following the above introduction to promising diagnostic and therapeutic bioelec-
tronic devices, we now discuss several challenges that exist when fabricating or using
bioelectronic devices. First, a challenge exists in creating better sensors to increase band-
width [
27
] and in the development of novel fabrication techniques. Second, maintaining
long-term biocompatibility, reliability, and stability of bioelectronic devices
in vivo
for
chronic application are key challenges. For skin-mounted devices or implantable devices,
it is also critical to maintain long-term conformability between the device and the elas-
tic, wrinkled skin, or tissue. The stability of bioelectronic devices can deteriorate due to
corrosion in the metal electrodes upon contact with biological electrolytes, susceptibility
of hydrolysis and oxidation to specific bonds in polymer structures, and delamination
providing access to deeper structures with the disruption in insulation barriers [
28
]. This
delamination due to disruption in insulation barriers or due to repetitive stretching of the
skin or tissue may cause a lack of conformability, which increases the impedances and
decreases the signal-to-noise ratio (SNR). Miniaturized ultra-small implantable technol-
ogy, such as “neural dust” [
29
] and “Stim dust” [
30
], enables access to small target areas
and can be implanted anywhere in the body, and interfaces directly with the tissue or
organ. Miniaturized biosensors also require small volumes for biological assays. However,
with miniaturization comes the issue of optimizing the configuration of electrodes and

Materials 2022,15, 1664 4 of 36
compromising the available power due to the need to reduce the battery size for minia-
turization [
31
]. In addition to these challenges, a key issue is the mechanical mismatch,
particularly between the bioelectronic device and the biological tissue or organ that it con-
tacts. The mechanical mismatch arises due to the material properties or the geometry of the
device. Conventional metal or silicon-based bioelectronics are stiff. As a result, they are not
only incapable of making tight contact with the tissue, but also cause irritation to the tissue.
Therefore, it is critical to fabricate a bioelectronic device that has mechanical compliance
with the tissue. This means the device needs to be soft to adapt to the complicated surface
curvature of tissue, and stretchable to comply with the repetitive stretching of the skin
or tissue (such as muscle tissue, gut tissue, blood vessel, and ducts). Various fabrication
techniques have been established for fabrication of hard and stiff bioelectronic devices,
which can be used as a foundation to develop soft and stretchable bioelectronics. However,
advances in fabrication technologies are required to potentially solve the issues associated
with the fabrication of soft, stretchable bioelectronics. Hence, the main focus of this paper
is the realization of soft and stretchable bioelectronics, by reviewing recent advances in
materials, fabrication technologies for sensors and transducers, communication methods,
and power sources.
3. Materials
In general, the ideal materials for bioelectronics are nontoxic, prevent direct harm
to the body, and minimize the initiation of the immune response. For wearable and
implantable devices, biocompatibility requirements differ. Table 1compares the different
biocompatibility requirements between the two categories of bioelectronics.
Table 1.
A table comparing the biocompatibility requirements (ISO 10993) for both implanted and
wearable devices [32].
Wearable Devices Implantable Devices
Nonirritating to Dermal tissue Nonirritating, compliant to match
surrounding tissue
Not cytotoxic, via leachable substances Not cytotoxic
Either non degrading, or degraded substances
are safe Degraded substances are safe
Hemocompatibility (Indirect contact) Hemocompatibility (direct contact)
Minimize/eliminate biofouling
Bioelectronics that interact directly or indirectly with blood must be hemocompatible;
that is, the device must prevent hemolysis, or the lysis of red blood cells. Both types of
devices must be nontoxic to cells. For wearables, cytotoxicity can occur when substances
leach into the skin. With implanted devices, cell damage can be caused by leaking sub-
stances, and physically by the device. Wearable bioelectronic devices that are applied to
the skin need to be made from non-irritating biomaterials, and devices that are implanted
must be manufactured to minimize the risk of irritating and damaging the surrounding
tissue. Although irritating the nearby tissue is inevitable for implants, there are ways to
decrease inflammation and prevent long-term initiation of the immune system. Another
biocompatibility issue for implants is biofouling. This occurs when the implanted device ac-
tivates the complement system, and macrophages and foreign body giant cells attach to the
implanted device and grow, inevitably disrupting the normal operation of the device [33].
This response makes the devices cytotoxic and unstable, and renders bioelectronic devices,
such as biosensors, unsuitable for long-term implantation [34].
Two ways to minimize the inflammatory response is to create ultrathin devices and/or
devices with a low Young’s modulus to closely match the Young’s/elastic modulus of the
tissue. Both increase the conformability of the device, allowing for accurate recordings [
35
].
Flexible and stretchable materials are less likely to inflame the nearby tissues and cause
local swelling, which also lead to inaccurate measurements. Close contact with tissue

Materials 2022,15, 1664 5 of 36
also increases the surface area for the recording devices, which increases the accuracy of
bioelectronics [5].
Bioelectronics are usually made with flexible substrates and are encapsulated to
protect the nearby tissue. To create compliant bioelectronics, the materials must have a
low modulus. Human tissue has a modulus range between 10 GPa for bone to 1 kPa for
some soft tissues. Other soft tissue, such as brain tissue, falls below 10 kPa [
35
]. Materials
with a lower modulus allows for a better signal transduction [
35
]. Stretchable electronics
must have high flexibility and should tolerate a tensile strain of at least 10% [
36
]. Plastics
and polymeric materials show the most promise for creating clinically relevant flexible
bioelectronics; however, there is still a mechanical mismatch between tissues and these
flexible materials. To bridge that gap, research has shifted to using polymers with hard
and soft segments and elastomers. For example, regular silicon (Si), a common substrate
for bioelectronics, can have a modulus of about 130 GPa, as shown in Figure 1a. Silicone
elastomers, such as PDMS (polydimethylsiloxane), have a lower Young’s modulus that is
suitable for soft tissue (Figure 1b).
Materials 2022, 15, x FOR PEER REVIEW 5 of 38
ways to decrease inflammation and prevent long-term initiation of the immune system.
Another biocompatibility issue for implants is biofouling. This occurs when the implanted
device activates the complement system, and macrophages and foreign body giant cells
attach to the implanted device and grow, inevitably disrupting the normal operation of
the device [33]. This response makes the devices cytotoxic and unstable, and renders bio-
electronic devices, such as biosensors, unsuitable for long-term implantation [34].
Two ways to minimize the inflammatory response is to create ultrathin devices
and/or devices with a low Young’s modulus to closely match the Young’s/elastic modulus
of the tissue. Both increase the conformability of the device, allowing for accurate record-
ings [35]. Flexible and stretchable materials are less likely to inflame the nearby tissues
and cause local swelling, which also lead to inaccurate measurements. Close contact with
tissue also increases the surface area for the recording devices, which increases the accu-
racy of bioelectronics [5].
Bioelectronics are usually made with flexible substrates and are encapsulated to pro-
tect the nearby tissue. To create compliant bioelectronics, the materials must have a low
modulus. Human tissue has a modulus range between 10 GPa for bone to 1 kPa for some
soft tissues. Other soft tissue, such as brain tissue, falls below 10 kPa [35]. Materials with
a lower modulus allows for a better signal transduction [35]. Stretchable electronics must
have high flexibility and should tolerate a tensile strain of at least 10% [36]. Plastics and
polymeric materials show the most promise for creating clinically relevant flexible bioe-
lectronics; however, there is still a mechanical mismatch between tissues and these flexible
materials. To bridge that gap, research has shifted to using polymers with hard and soft
segments and elastomers. For example, regular silicon (Si), a common substrate for bioe-
lectronics, can have a modulus of about 130 GPa, as shown in Figure 1a. Silicone elasto-
mers, such as PDMS (polydimethylsiloxane), have a lower Young’s modulus that is suit-
able for soft tissue (Figure 1b).
Figure 1.
(
a
) Young’s modulus of multiple biomaterials compared to cells [
37
]. (
b
) Young’s modulus
scale comparing multiple types of bioelectronics and the moduli of multiple tissue types in the human
body [38].
However, these materials are still not suitable for long-term implantation due to the
mechanical mismatch between the non-stretchable polymers and the softer tissues. This
mismatch causes inflammation that results in swelling, and thereby decreases the accuracy
of the recording device. To combat this, research is aimed at creating hybrid materials with
suitable chemical and mechanical properties by changing the chemical structure of these
materials, i.e., adding hydroxyl groups to increase stretchability [
39
]. Mixing materials can
result in a material with favorable characteristics from both materials. The most common
example is adding an elastomer to a material to increase the stretchability
[39,40]
. Another
method involves using a material with the wanted bulk properties and treating the surface

Materials 2022,15, 1664 6 of 36
of the material to include characteristics that immediately effect the surrounding environ-
ment. Surface treatments are performed to increase adhesion to other materials [
41
], to
increase electrochemical properties, and to improve biocompatibility. One example of a
surface treatment is the addition of a hydrophilic coating to a device to increase cell adhe-
sion [
42
]. This section focuses on common materials and hybrids used to create the different
components in flexible and stretchable bioelectronics, the substrate and encapsulation
layers, and the conducting portions such as traces or recording electrode arrays.
3.1. Polymeric Substrates
Substrate and encapsulation materials are nonconducting materials. Substrates are
used as the first layer for the electronics to be printed on and can be used to encapsulate in-
ternal components and or protect the entire device. These materials protect the components
from mechanical deformation or environmental erosion caused by the body. In addition,
the materials can provide adhesion for the device via van der Waals forces, covalent bonds,
or hydrogen bonds. Polymers are commonly used as substrates because they are cheap to
manufacture and are usually highly tunable.
Thin-filmed parylene-C is a flexible polymer with an elastic modulus of about 3.2 GPa
and a Poisson’s ratio of 0.4 [
43
]. Therefore, it is a good material for hydrophobic coatings
around bioelectronics for
in vivo
and
in vitro
tests. Ji et al. used parylene-c to coat their
electrocorticogram (ECoG) electrodes, which were constructed from nanostructured gold
and conductive PEDOT [
43
]. A schematic of the design is shown in Figure 2. The conducting
traces were made into a serpentine structure to increase its flexibility and stretchability,
which is discussed in more detail in Section 3of this review.
Materials 2022, 15, x FOR PEER REVIEW 6 of 38
Figure 1. (a) Young’s modulus of multiple biomaterials compared to cells [37]. (b) Young’s modulus
scale comparing multiple types of bioelectronics and the moduli of multiple tissue types in the hu-
man body [38].
However, these materials are still not suitable for long-term implantation due to the
mechanical mismatch between the non-stretchable polymers and the softer tissues. This
mismatch causes inflammation that results in swelling, and thereby decreases the accu-
racy of the recording device. To combat this, research is aimed at creating hybrid materials
with suitable chemical and mechanical properties by changing the chemical structure of
these materials, i.e., adding hydroxyl groups to increase stretchability [39]. Mixing mate-
rials can result in a material with favorable characteristics from both materials. The most
common example is adding an elastomer to a material to increase the stretchability [39,40].
Another method involves using a material with the wanted bulk properties and treating
the surface of the material to include characteristics that immediately effect the surround-
ing environment. Surface treatments are performed to increase adhesion to other materi-
als [41], to increase electrochemical properties, and to improve biocompatibility. One ex-
ample of a surface treatment is the addition of a hydrophilic coating to a device to increase
cell adhesion [42]. This section focuses on common materials and hybrids used to create
the different components in flexible and stretchable bioelectronics, the substrate and en-
capsulation layers, and the conducting portions such as traces or recording electrode ar-
rays.
3.1. Polymeric Substrates
Substrate and encapsulation materials are nonconducting materials. Substrates are
used as the first layer for the electronics to be printed on and can be used to encapsulate
internal components and or protect the entire device. These materials protect the compo-
nents from mechanical deformation or environmental erosion caused by the body. In ad-
dition, the materials can provide adhesion for the device via van der Waals forces, cova-
lent bonds, or hydrogen bonds. Polymers are commonly used as substrates because they
are cheap to manufacture and are usually highly tunable.
Thin-filmed parylene-C is a flexible polymer with an elastic modulus of about 3.2
GPa and a Poisson’s ratio of 0.4 [43]. Therefore, it is a good material for hydrophobic coat-
ings around bioelectronics for in vivo and in vitro tests. Ji et al. used parylene-c to coat
their electrocorticogram (ECoG) electrodes, which were constructed from nanostructured
gold and conductive PEDOT [43]. A schematic of the design is shown in Figure 2. The
conducting traces were made into a serpentine structure to increase its flexibility and
stretchability, which is discussed in more detail in Section 3 of this review.
Figure 2.
Stretchable Parylene-C electrodes. (
a
) A schematic of stretchable parylene-C electrodes
with serpentine interconnects to increase the stretchability of the electrodes. A silicone rubber
adhesive was used to stick the electrodes onto the elastomer substrate. Attaching the electrodes
to an elastic substrate limited external deformation [
43
]. (
b
) A diagram of the strain applied to
the electrodes when stretched [
43
]. Reprinted with permission from Ref. [
43
]. Copyright 2020 The
Chinese Ceramic Society.
To support insertion, biodegradable silk fibrin was added to the stretchable parylene-C
electrodes for easier surgery afterwards; the fibrin degraded over time. After six months of
in vivo
implantation in rat brains, 75% of these devices still showed a stable impedance for
accurate recordings [
44
]. Currently, parylene-c is used as a substrate for neural electrodes
due to its low modulus and chemical stability. However, in chronic implants, parylene-c
can result in unfavorable biocompatibility. One study found that in
in vitro
samples aged
one year, a tissue capsule around the device kept increasing in size for up to 16 weeks. This
tissue capsule revealed macrophages and fibroblasts were in the area, but no cytokines
were present after week 8 when initial inflammation was resolved [
45
]. The researchers
of this study concluded that parylene-c induced more tissue deposition in the local area

Materials 2022,15, 1664 7 of 36
than the polymer polyimide and, for chronic implants, parylene-c was ruled out as a
favorable substrate. However, their analysis also stated that the implant did not harm
neural tissues nor was the efficiency of the neural device decreased after encapsulation. This
encapsulation occurred because parylene-c is relatively more hydrophilic than polymide,
which attracted more matrix deposition [45].
Silicones are polymeric siloxanes with a wide variety of use, from medical equipment
and implants to houseware. Medical grade silicones and silicone rubbers are bioinert, water
resistant, and biocompatible. Low-density silicone provides elasticity and flexibility with
simple processing methods. Silicone can be used as a material for encapsulation. This
material is most common for flexible and stretchable bioelectronics due to its diverse uses
and easy fabrication method. It can also be bought commercially as Ecoflex, which has a
low enough modulus for coating bioelectronic devices.
Another common silicone used in flexible bioelectronics is poly(dimethylsiloxane)
(PDMS). PDMS is a biostable polymer that can be easily modified by embedding other
substrates or manipulating its molecular structure [
46
]. PDMS is a favorable biomaterial
because its surface and bulk modifications can result in several changes: changes in stiff-
ness, the creation of anti-bacterial properties, and a short-term change into a hydrophilic
polymer [
47
]. Conductive materials can be used inside of etched tracks to create flexible
circuits. PDMS has a stretchability of up to 1000% [
47
]. Its flexibility is determined by its
Si–O siloxane backbone. The average Young’s modulus is ~1–3 MPa. Super soft PDMS
has a lower modulus of 0.1 kPa–2.3 MPa, which is favorable for soft tissues. Conductive
materials such as silver, gold, carbon nanotubes, graphene, or aluminum oxide can be
added into etched tracks within PDMS to create a functioning flexible and stretchable circuit.
At the National Taiwan University, Chun-Yi Li and Ying-Chih Liao used plasma-treated
PDMS to create a flexible, stretchable, and printable thin-filmed pattern. The film was
plasma treated and a silica-like layer was added to the surface of the PDMS to increase
the wettability. Finally, epoxy binder and silver paste were added to the silica layer as a
printed pattern. These printed conductive films were bended, twisted, and stretched, all
while retaining conductivity. Their fabrication method created a thin film with excellent
mechanical stability [41].
The biomedical use of polyurethane (PU), specifically ether based [
42
], has gained
more traction in this century and has become a common material for vascular grafts, due
to its high tear and abrasion resistance. PU contains three components: a diisocyanate or
triocyanate, a chain extender, and a macrodiol, as shown in Figure 3[48].
Materials 2022, 15, x FOR PEER REVIEW 8 of 38
Figure 3. Chemical structure of polyurethanes. (a) The basic components of polyurethanes [49]. (b).
Structure of the elastomer thermoplastic polyurethane which contains hard and soft segments [50].
Reprinted with permission from Ref. [50]. Copyright 2020 Springer Science Business Media.
PU is highly tunable when mixed with other polymers. Pure PU has a Young’s mod-
ulus of ~91 MPa and a stretchability of 290% [51]. As an elastomer, it has a high flex-fatigue
life and a better mechanical degradation resistance compared to silicone elastomers [48].
The mechanical behavior of the elastomer is also tunable because its Young’s modulus is
dependent on the temperature and humidity [48]. This further proves that PU is another
organic material with multiple applications for flexible and stretchable bioelectronics.
Elastomeric PU has even been advanced to be self-healing, with an elongation at break of
1900% [52]. Like PDMS, PU can be used as a base insulator to layer conductive tracts upon
it. By adding silver flakes to pristine PU, Zhou Li and his colleagues were able to increase
the conductivity of the flexible polymer for an adhesive [53]. One research group used
polyurethane as a scaffold to hold 3D networks of the conductive material boron nitride.
They used PU to increase the flexibility and stretchability of their thermally conductive
composite film. This film did not hold electrical conductivity above strains of 40% after
100 stretching cycles, indicating breaks in the conductive paths [54]. The conductive ma-
terials chosen are an important part of any stretchable and flexible device. The type of
material and its manufactured shape will determine how stretchable and reliable the de-
vice will be.
3.2. Conductive Materials
Conductive materials are imperative for bioelectronics to create conductive traces
and the recording portion of the device. Metallic materials such as silver (Ag), copper
(Cu), gold (Au), and carbon-based materials are highly conductive materials. However,
these traces and electrodes must be flexible and stretchable to undergo repeated excessive
strains without break. There are multiple avenues to create promising flexible and stretch-
able conductive materials: (1) hard metals can be manufactured into thin nanowires, (2)
flexible materials can be infused in an elastic matrix, (3) conductive polymers or carbon-
based materials can be used [55], and (4) by adding metallic flakes in a polymeric sub-
strate, manufacturing ultrathin metallic plates, and liquifying metals [55]. Regardless of
the method chosen, these materials must be able to remain electrically conductive and
stable after withstanding several cyclic strains.
3.2.1. Metal Nanowires
Creating nanowires is one way to make conductive materials stretchable and flexible
for bioelectronics. Silver, copper, and gold are all promising materials for bioelectronics.
Silver nanowires (AgNWs) have the highest electrical conductivity of all metal materials;
Figure 3.
Chemical structure of polyurethanes. (
left
) The basic components of polyurethanes [
49
].
(
right
). Structure of the elastomer thermoplastic polyurethane which contains hard and soft seg-
ments [
50
]. Reprinted with permission from Ref. [
50
]. Copyright 2020 Springer Science Business Media.

Materials 2022,15, 1664 8 of 36
PU is highly tunable when mixed with other polymers. Pure PU has a Young’s modu-
lus of ~91 MPa and a stretchability of 290% [
51
]. As an elastomer, it has a high flex-fatigue
life and a better mechanical degradation resistance compared to silicone elastomers [
48
].
The mechanical behavior of the elastomer is also tunable because its Young’s modulus
is dependent on the temperature and humidity [
48
]. This further proves that PU is an-
other organic material with multiple applications for flexible and stretchable bioelectronics.
Elastomeric PU has even been advanced to be self-healing, with an elongation at break of
1900% [
52
]. Like PDMS, PU can be used as a base insulator to layer conductive tracts upon
it. By adding silver flakes to pristine PU, Zhou Li and his colleagues were able to increase
the conductivity of the flexible polymer for an adhesive [
53
]. One research group used
polyurethane as a scaffold to hold 3D networks of the conductive material boron nitride.
They used PU to increase the flexibility and stretchability of their thermally conductive
composite film. This film did not hold electrical conductivity above strains of 40% after
100 stretching cycles, indicating breaks in the conductive paths [
54
]. The conductive materi-
als chosen are an important part of any stretchable and flexible device. The type of material
and its manufactured shape will determine how stretchable and reliable the device will be.
3.2. Conductive Materials
Conductive materials are imperative for bioelectronics to create conductive traces and
the recording portion of the device. Metallic materials such as silver (Ag), copper (Cu),
gold (Au), and carbon-based materials are highly conductive materials. However, these
traces and electrodes must be flexible and stretchable to undergo repeated excessive strains
without break. There are multiple avenues to create promising flexible and stretchable
conductive materials: (1) hard metals can be manufactured into thin nanowires, (2) flexible
materials can be infused in an elastic matrix, (3) conductive polymers or carbon-based
materials can be used [
55
], and (4) by adding metallic flakes in a polymeric substrate,
manufacturing ultrathin metallic plates, and liquifying metals [
55
]. Regardless of the
method chosen, these materials must be able to remain electrically conductive and stable
after withstanding several cyclic strains.
3.2.1. Metal Nanowires
Creating nanowires is one way to make conductive materials stretchable and flexible
for bioelectronics. Silver, copper, and gold are all promising materials for bioelectronics.
Silver nanowires (AgNWs) have the highest electrical conductivity of all metal materials;
however, due to the high cost of silver, they are usually avoided for creating bioelectronics.
Gold nanowires (AuNWs) can also be used, but can also be costly. Copper nanowires
(CuNWs) are most used due to the lower price and excellent electrical conductivity of
copper. Conductive polymer nanowires (CPNWs) are also an excellent alternative to
expensive metal NWs; one example is discussed later in this section.
CuNWs are highly conductive and low in cost but have poor dispersity. To combat this,
Huang and his team infiltrated CuNWs into poly(styrene-block butadiene-block-styrene)
(SBS), to create printable CuNW-based composites [
55
]. These printable composites were
able to stretch and maintained conductivity after 1000 cycles of 4.0 mm bend radius. The
printable nanowires had high electrical conductivity with a break elongation up to 920%.
The elastic modulus of the wires was reported to be 0.081 GPa. The group only tested
the composites on paper to create flexible circuits; however, this material is an excellent
prospect for creating traces on stretchable and flexible substrates [56].
3.2.2. Conductive Polymers and Conductive Liquids
Conductive polymers are excellent materials for flexible bioelectronics; however, they
have limited stretchability because of their rigid backbones. One example is Poly(3,4-
ethylenedioxythiophene): poly(styrenesulfonate) (PEDOT:PSS). PEDOT:PSS is a polymer
with tunable electrical conductivity. PEDOT:PSS is the combination of the PEDOT monomer
and the dopant, PSS. By applying a current over the mixture, the monomer oxidizes and the

Materials 2022,15, 1664 9 of 36
PSS becomes electrostatically bound to the polymer backbone [
57
]. Organic solvent doping
treatments, ionic liquid treatments, strong acid soaking treatments, and acid-assisted
transfer-printing are all treatments that can be used to increase the electrical conductivity of
this polymer. PEDOT can be mixed with other polymers to create a flexible and stretchable
conductive polymer. Hansen and his team created a conductive polymer mix of PSS and
a PU elastomer polymer that increased in conductivity as it was being stretched, and
decreased in conductivity when relaxed [
58
]. A blend of PEDOT:PSS and PDMS also
proved to be a highly stretchable conductor [
59
]. Both these blends maintained electrical
stability after strains; however, creating these composites proved to be challenging due to
difficulty in phase separation [
35
]. Another way to increase the stretchability of PEDOT:PSS
is to add multiple hydroxyl groups to the polymer during synthesis; researchers stated an
increase in the elongation until break from <10% to >50% [39].
A number of problems remain to be addressed in the fabrication of printable circuitry
before advancing. Some problems include the decrease in measurement accuracy after
repeated stress and strain loads. This can be caused by the separation of the polymer
and conductive materials, or microcracks creeping into the conductive tracts of electrodes.
To combat the loss of conductivity in the electrodes, chemically stable liquid metals are
favored for flexible and stretchable devices. Gallium-indium alloys are a common liquid
metal to use for printable circuity. A self-healing, eutectic gallium-indium alloy (78 wt%
Ga and 22 wt% In) was reported to have a low viscosity for fabricating the device and
high conductivity [
60
]. However, using this alloy has a major problem during processing:
when in the presence of oxygen, it is incompatible with most standard liquid processing
techniques [60].
Polyaniline (PANI) is a conductive polymer popularized due to its low cost, tunability,
environmental stability, high temperature resistance, and high electrical conductivity [61].
Its structure consists of benzene rings in the backbone, making the polymer initially non-
stretchable. Pure PANI is highly soluble, and has poor mechanical properties, making
it hard to create sensors with [
62
]. PANI’s electrical conductivity decreases over time
and in physiological environments [
63
]. Despite this, PANI is an acceptable polymer for
conductive traces. By electrospinning PANI into one-dimensional nanowires, researchers
can increase its flexibility [
63
]. This polymer can also be woven into nanofibers to make the
polymer stretchable. One application of this is a stretchable temperature sensor [64].
3.3. Carbon-Based Materials
Carbon-based materials are other conductive materials that can be used for traces
or electrodes. Some carbon-based materials include graphene and carbon black. These
materials are known for their highly conductive properties and their unique molecular
structures.
3.3.1. Graphene
Graphene is characterized as carbon molecules stacked in a monolayer honeycomb.
Its Young’s modulus is 1 TPa. Graphene has excellent flexibility because the materials’
atoms are atomically thin; it also has some stretchability due to its hexagonal lattice struc-
ture [
65
]. However, the material’s conductivity can be destroyed by an elongation above
20%; thus, cross-linking graphene with an elastomer is important to enhance stretchabil-
ity [
66
]. Graphene is commonly grown by chemical vapor deposition (CVD), which is
discussed in the next section. Common metals used are nickel, iron, and copper. Different
metals, temperatures, gas pressures, and compositions affect the number of layers and
structure of the created graphene. For example, increasing solubility of carbon results
in a thicker graphene film. Once formed, graphene is scraped off the metal and used to
form different shapes: one-dimensional (fullerene), two-dimensional (nanotubes), and
three-dimensional (graphite) (Figure 4a) [
65
]. Carbon nanotubes and graphene flakes can
both serve as conductive fillers or base conductive material. Graphene can be woven into
textile materials to increase conductivity [47].

Materials 2022,15, 1664 10 of 36
Materials 2022, 15, x FOR PEER REVIEW 10 of 38
3.3. Carbon-Based Materials
Carbon-based materials are other conductive materials that can be used for traces or
electrodes. Some carbon-based materials include graphene and carbon black. These mate-
rials are known for their highly conductive properties and their unique molecular struc-
tures.
3.3.1. Graphene
Graphene is characterized as carbon molecules stacked in a monolayer honeycomb.
Its Young’s modulus is 1 TPa. Graphene has excellent flexibility because the materials’
atoms are atomically thin; it also has some stretchability due to its hexagonal lattice struc-
ture [65].
However, the material’s conductivity can be destroyed by an elongation above
20%; thus, cross-linking graphene with an elastomer is important to enhance stretchability
[66]. Graphene is commonly grown by chemical vapor deposition (CVD), which is dis-
cussed in the next section. Common metals used are nickel, iron, and copper. Different
metals, temperatures, gas pressures, and compositions affect the number of layers and
structure of the created graphene. For example, increasing solubility of carbon results in
a thicker graphene film. Once formed, graphene is scraped off the metal and used to form
different shapes: one-dimensional (fullerene), two-dimensional (nanotubes), and three-
dimensional (graphite) (Figure 4a) [65]. Carbon nanotubes and graphene flakes can both
serve as conductive fillers or base conductive material. Graphene can be woven into textile
materials to increase conductivity [47].
Figure 4. (a) Two-dimensional graphene sheets in various shapes: green shows 1D fullerene shapes,
red shows graphene rolled into nanotubes, blue shows graphene stacked into 3D graphite [67] Re-
printed with permission from Ref. [67]. Copyright 2007 Nature Publishing Group.; (b) atomic ar-
rangement of single- and double-layered graphene [68]. Reprinted with permission from Ref. [68].
Copyright 2010 Elsevier Ltd.
Graphene is a carcinogen and can be toxic to the liver, kidney, brain, and a few other
organs. It can also initiate chronic inflammation resulting in granulomas, or local buildup
of immune cells, once accumulated in the body [69]. Cytotoxicity has been found to be
related directly to the morphology and concentration of graphene [69]. Therefore, when
used in bioelectronics, graphene must be encapsulated or avoided. Another reason gra-
phene is not often used is its high cost. Further research and testing of different modifica-
tions are key to its increased use as a biomaterial [69].
Figure 4.
(
a
) Two-dimensional graphene sheets in various shapes: green shows 1D fullerene shapes,
red shows graphene rolled into nanotubes, blue shows graphene stacked into 3D graphite [
67
]
Reprinted with permission from Ref. [
67
]. Copyright 2007 Nature Publishing Group.; (
b
) atomic
arrangement of single- and double-layered graphene [
68
]. Reprinted with permission from Ref. [
68
].
Copyright 2010 Elsevier Ltd.
Graphene is a carcinogen and can be toxic to the liver, kidney, brain, and a few other
organs. It can also initiate chronic inflammation resulting in granulomas, or local buildup of
immune cells, once accumulated in the body [
69
]. Cytotoxicity has been found to be related
directly to the morphology and concentration of graphene [
69
]. Therefore, when used in
bioelectronics, graphene must be encapsulated or avoided. Another reason graphene is not
often used is its high cost. Further research and testing of different modifications are key to
its increased use as a biomaterial [69].
3.3.2. Carbon Black
Carbon black (CB) is a carbon-based material [
70
] that is preferred to graphene due to
its low cost and conductivity. It is found in soot and is biocompatible due to its spherical
shape [
71
]. CB is amorphous or shapeless carbon. Unlike graphene, it does not have
a defined crystal structure. Since CB does not have a structure, it can be mixed with
other polymers to make it stretchable and flexible. CB and Ecoflex can be mixed to create
electrodes. These electrodes have been printed on a barium titanate elastomer composite,
and maintained stability through 1000 cycles of 100% strain [
71
]. Another application used
CB nanoparticles mixed with Ecoflex to create conductive traces inside more Ecoflex to
create a stretchable nanocomposite piezo-resistor, decreasing resistance with increased
strain [72].
Although all the biomaterials mentioned in Table 2have been safely tested with
regards to the human body, surface modifications or encapsulation are still important to
increase device longevity. These favorable device characteristics mentioned are for both
implantable and wearable electronics. Wearable devices should have high mechanical
strength to withstand the outside wear and tear, and must be biocompatible to prevent skin
irritation. For implantable devices, however, the body’s environment can be detrimental
to any foreign object. Body temperature, pH, and blood components can all increase the
breakdown and release of ions from the implanted materials.

Materials 2022,15, 1664 11 of 36
Table 2.
Common bioelectronic materials with characteristics related to flexibility and stretchability.
(all these values are for the listed applications, as each of the materials are highly tunable.).
Materials Company Young’s Modulus Stretchability Poisson’s Ratio Applications
Silicone Azo Materials 0.001–0.05 GPa 1000%–2000% [73] 0.47–0.49 [74] Insulation
PDMS Dow’s Sylgard 184 ~1–3 MPa [75] ~1000% [47] ~0.5 [76]
Encapsulation/
Nonconductive
polymer
Polyimide Dupont 2.5 GPa 260% [77] 0.34 at 23 ◦CNeurocortical
electrode arrays
Parylene-C - 2.8 GPa, 3.2 GPa,
4.5 GPa [78,79]20–200% [78] 0.4 [79]Neural electrode
substrate
PEDOT-PSS - 1–7.5 GPa ~10% [80] 0.34 Conductive tracts
Polyurethane
elastomer -4.7 MPa, 5.3 MPa,
7.4 MPa [48]~1890% [51] 0.45–0.5 Vascular grafts,
blood bags
Graphene - ~1 TPa [81] ~30% [82] 0.456 ±0.008 [83]
Logic gates, transistors
4. Engineering Organic/Inorganic Functional Materials in Stretchable
Bioelectronic Devices
Flexible/stretchable devices for wearable and bio-integrated electronics not only re-
quire a flexible/stretchable substrate, but also electronic systems that possess mechanical
flexibility, stretchability, and conformal features, in addition to excellent electrical proper-
ties. High-throughput, large-scale, and cost-efficient fabrication of flexible and stretchable
bioelectronic devices requires a novel approach to material design, sensing material, fabrica-
tion technique, and sensing mechanism [
84
]. To achieve stretchable bioelectronic interfaces,
several attempts to merge the properties of high stretchability of elastomer and high con-
ductivity of organic/inorganic conductors have been made. Flexibility and stretchability in
conductive materials can generally be achieved by using (1) geometrically soft conductors
through the use of different structural configurations, (2) conductive polymer composites,
or (3) conductive liquids [85].
The conductive materials listed in Section 2can be either hard materials or intrinsically
stretchable materials. As mentioned earlier, hard materials such as pure metals (Ag, Cu, Au,
etc.) offer the advantage of having high conductivity for efficient signal transduction and
high sensitivity. However, those inorganic materials are rigid and do not possess mechanical
flexibility and stretchability. Processing these materials into thin films and nanowires (NWs)
using methods such as chemical vapor deposition (CVD) and lithography add the property
of flexibility, but not stretchability [
86
,
87
]. Hence, researchers developed a strategy where
they couple mechanically guided structural designs on those materials to accommodate
large deformation, and embed those patterned materials in elastic substrate. Common
mechanically guided structural configurations, i.e., serpentine geometry, wavy/wrinkled
structures, and fractal motifs, are discussed in Section 3.1.
Polymer composites with conductive fillers (e.g., carbon black, graphene flakes, metal
NWs, carbon nanotubes, metal nanoparticles), and liquid metals, when processed into
thin films, can absorb mechanical deformation by themselves and are considered to be
intrinsically stretchable materials [
88
,
89
]. To process those conductive polymer composites
and liquid materials into thin films, several fabrication techniques, such as soft lithography,
printing, and wet spinning, are utilized, and are discussed later in the section. In addition,
coating, dipping, and transfer printing can also be utilized to achieve thin films from
conductive liquids. A chart summarizing the processing of these conductive materials into
a flexible/stretchable format is shown in Figure 5.

Materials 2022,15, 1664 12 of 36
Materials 2022, 15, x FOR PEER REVIEW 12 of 38
The conductive materials listed in Section 2 can be either hard materials or intrinsi-
cally stretchable materials. As mentioned earlier, hard materials such as pure metals (Ag,
Cu, Au, etc.) offer the advantage of having high conductivity for efficient signal transduc-
tion and high sensitivity. However, those inorganic materials are rigid and do not possess
mechanical flexibility and stretchability. Processing these materials into thin films and
nanowires (NWs) using methods such as chemical vapor deposition (CVD) and lithogra-
phy add the property of flexibility, but not stretchability [86,87]. Hence, researchers de-
veloped a strategy where they couple mechanically guided structural designs on those
materials to accommodate large deformation, and embed those patterned materials in
elastic substrate. Common mechanically guided structural configurations, i.e., serpentine
geometry, wavy/wrinkled structures, and fractal motifs, are discussed in Section 3.1.
Polymer composites with conductive fillers (e.g., carbon black, graphene flakes,
metal NWs, carbon nanotubes, metal nanoparticles), and liquid metals, when processed
into thin films, can absorb mechanical deformation by themselves and are considered to
be intrinsically stretchable materials [88,89]. To process those conductive polymer com-
posites and liquid materials into thin films, several fabrication techniques, such as soft
lithography, printing, and wet spinning, are utilized, and are discussed later in the section.
In addition, coating, dipping, and transfer printing can also be utilized to achieve thin
films from conductive liquids. A chart summarizing the processing of these conductive
materials into a flexible/stretchable format is shown in Figure 5.
Figure 5. A chart summarizing the engineering of conductive material into flexible and stretchable
format.
4.1. Structural Configurations to Render Stretchability to Hard Inorganic Materials
To transform the non-stretchable electrically conductive material into a stretchable
format, researchers use the strategies of mechanically-guided structural configurations.
Creation of different structural configurations, such as serpentine geometry, wrinkled na-
nomembranes, wavy structures for geometrically soft conductors, and fractal layouts,
have aided in dramatically improving the stretchability of devices. Conductive traces can
be deposited into different shapes, such as curvilinear shapes with arc sections. Integrat-
ing arc sections instead of sharp turns into an electrical wire improves the elastic mechan-
ics [90]. One example of these curvilinear shapes is the serpentine structure. Serpentine
Figure 5.
A chart summarizing the engineering of conductive material into flexible and stretch-
able format.
4.1. Structural Configurations to Render Stretchability to Hard Inorganic Materials
To transform the non-stretchable electrically conductive material into a stretchable
format, researchers use the strategies of mechanically-guided structural configurations.
Creation of different structural configurations, such as serpentine geometry, wrinkled
nanomembranes, wavy structures for geometrically soft conductors, and fractal layouts,
have aided in dramatically improving the stretchability of devices. Conductive traces can
be deposited into different shapes, such as curvilinear shapes with arc sections. Integrating
arc sections instead of sharp turns into an electrical wire improves the elastic mechanics [
90
].
One example of these curvilinear shapes is the serpentine structure. Serpentine geometry
(Figure 6B) enables the in-plane stretchability by minimizing the maximum tensile strain
(
εmax
), which is strongly dependent on the adhesion with a stretchable substrate [
91
]. In a
serpentine structure, four geometric parameters—width (w), radius (R), arc angle (
α
), and
arm length (l) (Figure 6A)—dictate the stretchability. Lee et al. concluded that stretchability
increases with a small w/R ratio, large l/R ratio, and large
α
[
91
]. Additionally, researchers
have created stretchable electronics by placing the semiconductor material on a pre-strained
elastomeric substrate. Such devices undergo a lateral compression upon removing the pre-
strain, and therefore attain out-of-plane wavy or wrinkled structures. This approach has
been exploited to fabricate stretchable organic photovoltaics [91]. Fan et al. demonstrated
a promising approach to integrate electronic material in soft elastomeric substrates [
90
].
The method involves patterning the metal wires in various fractal motifs. The topology
types selected for creating those fractal motifs were lines (Peano), loops (Vicsek), and
branch-like meshes (Greek cross) (Figure 6C). These fractal constructs consist of a high-
order iteration of lines, loops, or meshes that are linked together. Those fractal constructs,
unlike the serpentine geometry, allow the researchers to control and tailor the stretchability
of the electronic devices depending on their applications. The orientation of the pattern
enhances the device’s elastic strain in the selected direction, thereby providing the ability
to accommodate different types of deformation modes from uniaxial and biaxial, to radial
deformation modes. The enhancement of stretchability can be attributed to factors such as
the geometric scaling of the arc sections, increase in the length of the wires, and addition
of high-order spring-like motifs. These fractal-based layouts have the potential to be
used for radio frequency-based applications, and offer a promising approach to devise
MRI-compatible skin-mountable or implantable devices.

Materials 2022,15, 1664 13 of 36
Materials 2022, 15, x FOR PEER REVIEW 14 of 38
Figure 6. Representative geometrical configurations for hard–soft material integration. (A). Sche-
matic representation of a serpentine unit cell. (B). Example of serpentine structured traces on an
electrode array [91]. Reprinted with permission from Ref. [91]. Copyright 2015. American Chemical
Society. (C). Different patterns of fractal-inspired layouts of metal wires (top), and FEM images of
each structure under elastic tensile strain [90]. Reprinted with permission from Ref. [90]. Copyright
2014 Nature Publishing Group.
4.2. Chemical Vapor Deposition for Creating Thin Films
Chemical vapor deposition (CVD) is a widely used manufacturing process in the
semiconductor industry to deposit or grow thin solid film on the substrate held in a reac-
tion chamber. This process has been extensively used to grow large-area and high-quality
graphene film on metal [92]. During CVD, the metal substrate to be coated is exposed to
a gaseous carbon source, such as methane, acetylene, or methanol. The decomposed car-
bon radicals from these gases precipitate out on the metal substrate to form layered gra-
phene. Commonly used metal substrates are Ni and Cu, which also act as a catalyst for
the reaction. Properties such as the number of layers, size, morphology, orientation, dop-
ing, defect, and grain boundaries can be controlled by manipulating parameters such as
the temperature, chamber pressure, gas flow rate, and source–substrate distance [93].
Figure 6.
Representative geometrical configurations for hard–soft material integration. (
A
). Schematic
representation of a serpentine unit cell. (
B
). Example of serpentine structured traces on an electrode
array [
91
]. Reprinted with permission from Ref. [
91
]. Copyright 2015. American Chemical Society.
(
C
). Different patterns of fractal-inspired layouts of metal wires (top), and FEM images of each
structure under elastic tensile strain [
90
]. Reprinted with permission from Ref. [
90
]. Copyright 2014
Nature Publishing Group.
4.2. Chemical Vapor Deposition for Creating Thin Films
Chemical vapor deposition (CVD) is a widely used manufacturing process in the
semiconductor industry to deposit or grow thin solid film on the substrate held in a reaction
chamber. This process has been extensively used to grow large-area and high-quality
graphene film on metal [
92
]. During CVD, the metal substrate to be coated is exposed
to a gaseous carbon source, such as methane, acetylene, or methanol. The decomposed
carbon radicals from these gases precipitate out on the metal substrate to form layered
graphene. Commonly used metal substrates are Ni and Cu, which also act as a catalyst for
the reaction. Properties such as the number of layers, size, morphology, orientation, doping,
defect, and grain boundaries can be controlled by manipulating parameters such as the
temperature, chamber pressure, gas flow rate, and source–substrate distance [
93
]. High-

Materials 2022,15, 1664 14 of 36
quality and sensitive 2D materials require few grain boundaries or large single crystals to
reduce charge scattering at the grain boundary and enhance the electronic, mechanical,
and thermal properties of the 2D material [
93
]. This challenge of growing large-grain
crystals was addressed by Rhyee and team using modified chemical vapor deposition
(m-CVD) for the fabrication of high-mobility 2D-layered transistors based on large-grain
and highly crystalline Molybdenum Diselenide (MoSe
2
) films grown onto SiO
2
substrates
(Figure 7b) [
94
]. The mechanical flexibility of these thin film transistors was realized by
transferring the MoSe
2
onto the flexible polyimide (PI) substrate. The fabricated transistors
have applications in human robotics and human-centered soft electronics. Other challenges
in growing high-quality large film include the random layer distribution, variation in
growth at different areas, transfer of the 2D film produced from the growth substrate to the
target device substrate, and unavoidable damage and contamination during transfer of the
film to the target [93].
Materials 2022, 15, x FOR PEER REVIEW 15 of 38
High-quality and sensitive 2D materials require few grain boundaries or large single crys-
tals to reduce charge scattering at the grain boundary and enhance the electronic, mechan-
ical, and thermal properties of the 2D material [93]. This challenge of growing large-grain
crystals was addressed by Rhyee and team using modified chemical vapor deposition (m-
CVD) for the fabrication of high-mobility 2D-layered transistors based on large-grain and
highly crystalline Molybdenum Diselenide (MoSe
2
) films grown onto SiO
2
substrates (Fig-
ure 7b) [94]. The mechanical flexibility of these thin film transistors was realized by trans-
ferring the MoSe
2
onto the flexible polyimide (PI) substrate. The fabricated transistors
have applications in human robotics and human-centered soft electronics. Other chal-
lenges in growing high-quality large film include the random layer distribution, variation
in growth at different areas, transfer of the 2D film produced from the growth substrate
to the target device substrate, and unavoidable damage and contamination during trans-
fer of the film to the target [93].
Figure 7.
Fabrication methods for stretchable electronics. (
a
–
c
) Chemical vapor deposition (CVD) (
a
).
Schematics of the synthesis mechanism of CVD graphene on Cu foil (left) and the roll-to-roll transfer

Materials 2022,15, 1664 15 of 36
process of graphene (right) [
95
]. Reprinted with permission from Ref. [
95
]. Copyright 2013 American
Chemical Society. (
b
). Modified CVD (mCVD). Schematic representation for synthesizing multilayer
MoSe
2
film. Schematic cross-section of a collection of MoSe
2
transistors on a flexible PI/PET sub-
strate [
94
]. Reprinted with permission from Ref. [
94
]. Copyright 2016 Wiley-VCH Verlag GmbH & Co.
(
c
). Fabrication process of the wrinkled graphene-AgNW hybrid electrode [
96
]. (
d
–
f
) Lithography
method. (
d
). UV Nanolithography (top) and soft lithography (bottom) [
97
] (
e
). Photolithographic
methods using masked irradiation and a negative photoresist material: (i) Patterning by single
exposure, (ii) patterning by layer-by-layer coating and exposure, (iii) tilted patterning by single
inclined exposure, (iv) patterning by double inclined exposure, (v) tapered patterns by rotating tilted
exposure [
97
]. Reprinted with permission from Ref. [
97
]. Copyright 2011 Elsevier Ltd. (
f
). Fabrication
of highly stretchable and transparent nanomesh electrodes via grain boundary lithography [
98
]. A
sheet of paper is laser-cut using the magnified image of Au nanomesh. The left image shows without
any cutting. The middle image shows the stretching of the structure after cutting a few ligaments.
The right image shows the stretching of the structure after cutting more ligaments. Reprinted with
permission from Ref. [98]. Copyright 2014 Nature Publishing Group.
CVD aids in growing thin film on substrate. Flexibility and stretchability can be
realized by transferring the thin film onto flexible/stretchable substrate. To realize the
stretchability on the device, it is ideal if the graphene layer can be grown directly on
the stretchable polymer substrate. However, one of the challenges of using the polymer
substrate for direct growth of graphene is that it possesses low processing temperature,
high surface roughness, and a low thermal expansion coefficient; therefore, it is not an
appropriate substrate for conventional CVD. Unlike CVD, plasma-enhanced CVD (PECVD)
possesses a low processing temperature, thereby providing an opportunity for direct
growth of graphene on flexible/stretchable substrate. However, its disadvantage is the
production of inferior graphene. Due to the inferior graphene qualities resulting from
PECVD, Jang et al. reported coupling CVD for growing graphene film on Cu with the
roll-to-roll transfer method to polymer substrate (Figure 7a) [
92
]. Similarly, atmospheric
pressure chemical vapor deposition (APCVD) coupled with the transfer method and
inkjet printing was also used for the fabrication of a conductible vertically aligned carbon
nanotubes (VACNTs)/PDMS/Graphene Oxide (GO) stretchable electrode [
99
]. It was
observed that after multi-cycle stretching the electrode up to 100% tensile strain, it did
not show an increase in sheet resistance [
99
]. Fabricating a highly stretchable wrinkled
graphene for a sensitive strain sensor and a smart window was also reported using the
CVD approach for growing graphene, and the transfer printing method for transferring
the graphene onto a pre-stretched elastomer (Figure 7c) [
96
,
100
]. Chen et al. incorporated
crumples and buckles on graphene by attaching the graphene film to a uniaxially pre-
stretched very-high-bond (VHB) film to support the stretchability [
100
]. The gauge factor,
defined as the ratio of
∆
R/R
0
to
ε
(resistance variation to applied strain), was 20.1 at the
working range of 105% tensile strain. Additionally, with small strains (
ε
) < 105%, the
resistance increased slowly due to the buckled graphene structure. However, increasing
the strain >105% led to formation of cracks, which resulted in a rapid increase in resistance.
In summary, CVD coupled with the transfer method is generally utilized to engineer
flexible/stretchable devices with high graphene quality.
4.3. Lithography Method for Creating Thin Films and Nanowires
Lithography is a popular key technology in the semiconductor industry used for the
fabrication of integrated circuits and microchips. This approach is utilized to fabricate a
desired pattern ranging from nanometer to micrometer size on the underlying substrate.
This fabrication approach has been extended to fabricate soft and stretchable electronics for
biomedical applications, such as skin-mountable and wearable electronics. Photolithog-
raphy, soft lithography, nanoimprint lithography, e-beam lithography, focused ion beam
lithography, scanning probe lithography, etc., are some of the widely used lithography
techniques. Commonly used photolithography is a masked UV lithography that uses a
positive or negative photoresist (Figure 7d). Upon exposing the negative photoresist to UV

Materials 2022,15, 1664 16 of 36
light using a photomask, the UV-light-exposed areas become polymerized, and therefore
become insoluble in developer solution. On the contrary, upon UV exposure using a mask,
the positive photoresist areas become depolymerized and are therefore easily dissolved
using developer solution. Generally, the photolithography process comprises of several
steps, such as photoresist application, prebake, mask alignment and exposure, develop-
ment, post bake, etching, and resist stripping (Figure 7e) [
97
]. Electron beam lithography,
in contrast, is mainly used to produce a photomask. It incorporates a high-resolution direct
writing approach using an electron beam on an electron-sensitive resist, and is therefore
not a suitable approach for mass production [
97
]. The soft lithography process enables
fabrication of structures or the transfer of material using an elastomeric stamp, mold, or
photomask (Figure 7d). This process is easy, fast, low-cost, mass reproducible, and can be
used for thin film fabrication.
Although lithography enables the deposition of metal on flexible/stretchable soft
substrates, the problem of crack formation on conductive metal conduits upon stretching
or bending the device remains. In addition to this technique, researchers use the strategy of
structural configurations to realize a highly stretchable device while keeping the conductive
materials intact to maintain the device’s conductivity. When engineered with serpentine-
shaped ligaments, metal films tend to become highly stretchable, as mentioned in Section 3.1.
This strategy of enhancing stretchability was presented by Guo et al. by fabricating Au
metal nanomesh electrodes using a method called grain boundary lithography [
98
]. This
lithography method involves depositing Au metal in undercuts created in the SiO
x
and
In
2
O
3
layer by isotropic etching followed by the lift-off process (Figure 7f). Au nanomesh
on PDMS substrate resulted in only a slight increase in resistance when it was stretched to
about 160% strain, or after 1000 cycles at a strain of 50%. This high stretchability due to the
serpentine-shaped ligaments is attributed to the distributed rupture of the Au ligaments
and the out-of-plane deflection of the serpentine ligaments. The nanowire approach of
fabricating stretchable sensors has also been exemplified by the research conducted by
Tiefenauer. Tiefenauer’s team utilized UV interference lithography for fabrication of a
nanowire array template on a PEN/PVA sheet, followed by gold evaporation onto those
templates and template stripping-based nano-transfer printing of stretchable plasmonically
coupled graphene structures on an elastomer for Surface Enhanced Raman Scattering
(SERS)-based sensing applications [
101
]. Another approach for highly stretchable low-
impedance electrode fabrication comprises microcracks, as demonstrated by Decataldo et al.
The team reported microcracked Ti/Au films covered with stretchable conducting polymer
composite fabricated via the method of lithography and electrodeposition for bioelectronic
recording from small peripheral nerves. The composite displayed high conductivity under
tensile strain exceeding 40% [85].
This section has discussed the deposition of hard metal onto a substrate using lithog-
raphy, and the realization of stretchable devices using either structural configurations or
transfer printing on a stretchable substrate. An alternative approach to realize soft and
stretchable microelectronics using lithography is the use of liquid metals such as eutectic
gallium-indium (eGaIn) alloy. EGaIn alloy is nontoxic, possesses excellent mechanical,
thermal, and electrical properties, and facilitates a broad range of patterning methods,
including soft lithography based on stamping, stencil printing, injection, and additive
and subtractive patterning processes. Lithography-enabled stencils have a low resolution
(
≈
20–200
µ
m) and pose challenges such as having a rough surface and loss of material
during lift-off. The microfluidic injection technique provides a higher resolution (>10
µ
m)
but requires microchannel thickness >50
µ
m. The subtractive technique is a low-cost and
facile method, but the material removal process is slow in the case of patterning small
features on large substrates. Additive patterning, like direct writing, has a low resolution
(
≈
100
µ
m) due to the size limitation of the injection nozzle. On the contrary, feature sizes
>2
µ
m can be achieved via soft lithography [
102
]. Therefore, a good selection of lithography
method is required to achieve high-quality stretchable devices using liquid metals. A
potential solution to overcome patterning challenges is employing wetting/nonwetting

Materials 2022,15, 1664 17 of 36
physical or chemical surface modifications, which help by providing smooth and uniform
deposition of thin films for multiscale patterning and functional integration [
102
]. By
combining additive and subtractive soft lithography with surface modification and 3D
heterogenous integration, Kim and team successfully fabricated eGaIn thin film-based
functional soft microsystems such, as a soft LC (inductor-capacitor) sensing platform, a
fingertip-mountable biological sensing platform, and soft heaters with heating capability.
The fabricated soft and stretchable circuit exhibited bending and strain deformation up to
50% while maintaining electrical conductivity.
4.4. Printing for Creating Thin Films of Conductive Liquids
Chemical vapor deposition (CVD) aids in fabrication of high-quality thin films of
graphene with low flexural rigidity on a metallic substrate. PECVD, by comparison,
can be used for direct fabrication of thin films of graphene on polymer for engineering
stretchable devices; however, it cannot produce high-quality graphene. Hence, in order
to engineer conductive materials to a stretchable format, researchers have to perform a
film transfer from the growth substrate to a target stretchable substrate. Therefore, this
does not represent an ideal fabrication procedure for the fabrication of a stretchable device.
Although lithography can be used to create a conductive circuit pattern having a size in
the range of nm to
µ
m on a soft stretchable substrate, it requires a planar substrate surface,
a photomask (for mask lithography), and expensive resources, such as a mask aligner, to
develop complicated structures. Moreover, due to phenomena such as diffraction and
the wavelength factor of optical sources, the lithography method may not produce high
resolution patterns on the substrate, and may also limit the method. Similarly, serpentine
and wavy structures of rigid materials limit the strain of devices and can lead to the
formation of cracks and fractures. Hence, conductive liquid metal is an emerging and
attractive field of research.
Printing is an additive manufacturing process that is a fast, accurate, and fully auto-
mated non-contact approach that uses liquid metal to create a conductive functional region
on a single platform [
60
]. In addition to the robustness, flexibility, and stretchability of
the device, printing also improves the repeatability and scalability of liquid metal process-
ing, and provides the capability of achieving electronics with complex patterns without
compromising the electrical conductivity of the devices [60].
To fabricate conductive electronics using printing techniques, it is pivotal to formu-
late conductive liquid metal ink with ideal features, such as the choice of the conductive
material and the solvent, and the appropriate viscosity, surface tension, and rheology for
achieving high printing performance and high accuracy [
103
]. A range of ink formulations
have manifested the potential to be employed for printing. Some of those formulations are
the commercially available AgNP ink [
104
,
105
], eGaIn liquid metal ink [
60
,
106
], toluene-
based ink composed of Semiconducting Single Walled Carbon Nanotubes (SC-SWCNTs)
and conjugated polymer (to prevent CNT aggregation) [
107
], poly (vinylidene fluoride-
co-hexafluoropropylene) (PVDF-HFP) in N-methyl-2-pyrrolidone anhydrous [
107
], and
poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) in solution of bis
(trifluoromethane) sulfonimide lithium [
107
]. These inks were prepared either by disper-
sion [
104
] or sonication [
60
,
107
]. Depending on its viscosity and rheology, the liquid ink
may tend to emerge as a ball rather than a stream. The rheology of printable ink can be
regulated for the jetting stream using additive agents or by optimizing the concentration of
the ink [
103
]. Surface energy mismatch between the substrate and the ink results in poor
wettability, which can be improved by decreasing the surface tension. The effective ways
to regulate surface tension are either by performing oxygen plasma or UV/ozone surface
treatment, by adding surfactants such as Triton X-100 or Zonyl FC-300 [
41
,
103
], or using
toluene [
101
]. Printing of conductive liquid is mostly followed by the sintering process to
create a conductive path. Sintering can be undertaken in several ways—self-sintering [
108
],
mechanical [
60
], thermal [
109
], chemical [
110
], or using pulsed light [
111
] or a focused laser
beam [
105
,
112
]. In addition to the robustness, flexibility, and stretchability of the device,

Materials 2022,15, 1664 18 of 36
printing also improves the repeatability and scalability of liquid metal processing and pro-
vides the capability of achieving electronics with complex patterns without compromising
the electrical conductivity of the devices [60].
Based on this knowledge about printing techniques, their advantages, and the consid-
erations to be made when formulating conductive inks, next we look at how researchers
have utilized this technique to fabricate stretchable devices. Mohammed G. Mohammed
combined inkjet printing for the extrusion printing of uncured base elastomer, i.e., Smooth-
Sil 950 layer and spray printing of eGaIn (78% Ga, 22% In) liquid metal ink prepared by
sonicating bulk liquid metal in a carrier solvent (Figure 8a) [
60
]. Selective mechanical sinter-
ing of slurry particles was performed to create a conductive path. Lastly, extrusion-printed
Sylgard 184 was used as an encapsulating elastomer. Lopes reported a novel method for
rapid fabrication of a thin-film circuit with integrated microelectronics using a printing
method and hydrographic transfer [
113
]. Highly conductive circuits were produced using
silver paste with eGaIn coating on a non-conductive circuit template printed on Transfer
Tattoo Paper (TTP) using a desktop printer (Figure 8b). Hydrographic transfer, also referred
as water transfer printing, was demonstrated to fabricate a stretchable circuit for its appli-
cation in human–machine interface devices (Figure 8b). Albrecht’s team fabricated highly
conductive and highly stretchable (up to 300% strain) wires on plasma-treated PDMS using
simple inkjet printing and commercial AgNP ink [104]. Lopez et al. fabricated stretchable
Field Effect Transistor (FET) arrays and interconnections that can bear in-plane stretching of
up to 20% strain by inkjet printing using PEDOT:PSS, PVDF-HFP, and SC-SWCNTs [107].
4.5. Wet Spinning Method for Flexible/Stretchable and Conductive Fiber Production
Due to a huge interest in flexible/stretchable electronics, the stretchable property of
everyday textiles has been exploited as a platform for wearable electronics. Fiber is the
fundamental unit of the textile. With the incorporation of conductive material, the fiber’s
functionality can be expanded to stretchable strain sensors, supercapacitors [
114
,
115
], nano-
generators [
116
,
117
], etc., for wearable electronics. These fiber-like electronics not only
provide excellent wearable properties, but also can withstand mechanical deformations
such as bending, folding, twisting, and stretching. Such wearable electronics can be easily
attached conformably onto rough and uneven surfaces such as human skin. As a result
of the benefits of fiber-like electronics, researchers developed several strategies, one of
which is the wet spinning method. The fiber wet spinning method is a facile and viable
strategy to fluidly spin the macroscopic fiber in a continuous way [
118
]. The overall
process of preparing conductive fibers includes preparation of dispersion, preparation
of the coagulation bath, and wet spinning. First, the dispersion solution is loaded into
a syringe pump or a spinneret. The dispersion is then extruded through the spinneret
into a non-solvent bath for coagulation. The fibers are collected in the roller and dried at
room temperature to remove the residual solvent in the fiber [
119
,
120
]. Post-processing
is generally carried out to enhance the conductivity, including modification of the for-
mulation with an organic solvent or conductive additives [
121
]. However, studies have
been conducted to develop a fast one-step process and eliminate the post-treatment [
121
].
Zhang et al. demonstrated the production of strong and conductive PEDOT:PSS fiber
in a one-step process by replacing the conventional coagulation bath with concentrated
sulfuric acid [
121
]. Cong et al. reported the processing of graphene oxide (GO) disper-
sion in a coagulation bath of hexadecyctrimethyl ammonium bromide (CTAB) solution
followed by chemical reduction to produce graphene fibers with mechanical strength of
182 MPa and conductivity of 35 S/cm (Figure 8c) [
118
]. The team also showed that soaking
GO fibers in epoxy resin solution can produce fibers with good rigidity, and integrating
poly(N-isopropylacrylamide) (PNIPAM) in GO dispersion can be undertaken to retain the
thermosensitive property of PNIPAM in the fibers. The assembly mechanism of GO fibers
is attributed to the electrostatic interaction between the CTAB and GO sheets. Tang et al.
expanded this fabrication strategy to develop one-step coaxial wet-spinning assembly to
fabricate silicone-elastomer-wrapped multiwalled CNT (MWCNT)-based core-sheath fiber

Materials 2022,15, 1664 19 of 36
with high stretchability above 300% and excellent stability (>1000 cycles) (Figure 8d) [
122
].
Seyedin et al. successfully demonstrated the wet spinning technique using polyurethane
(PU) elastomer, PEDOT:PSS-conducting filler, dimethyl sulfoxide (DMSO) solvent for stable
dispersion, and 80/20 v/visopropanol/water mixture as a coagulation bath [123]
Some of the considerations required during fabrication are that the spinning ink must
exhibit shear-thinning behavior for efficient flow, and an appropriate coagulation bath
should be chosen so that the extruded spinning ink is rapidly vulcanized without any
breakup [
122
]. The diameter of the fibers must be easily controlled by the nozzle size
of the spinneret [
119
]. The fibers intended to be used as strain sensors need to have a
high sensitivity for sensing tiny deformations during respiration, heartbeat, and speech
recognition, and a broad sensing range to detect large movements, such as finger bending,
walking, and running. In order to achieve both features, dip-coating of fibers has been
reported [
124
,
125
]. A summary of advantages and disadvantages of each of the fabrication
methods to create a flexible/stretchable device is presented in Table 3.
Materials 2022, 15, x FOR PEER REVIEW 20 of 38
isopropylacrylamide) (PNIPAM) in GO dispersion can be undertaken to retain the ther-
mosensitive property of PNIPAM in the fibers. The assembly mechanism of GO fibers is
attributed to the electrostatic interaction between the CTAB and GO sheets. Tang et al.
expanded this fabrication strategy to develop one-step coaxial wet-spinning assembly to
fabricate silicone-elastomer-wrapped multiwalled CNT (MWCNT)-based core-sheath fi-
ber with high stretchability above 300% and excellent stability (>1000 cycles) (Figure 8d)
[122]. Seyedin et al. successfully demonstrated the wet spinning technique using polyure-
thane (PU) elastomer, PEDOT:PSS-conducting filler, dimethyl sulfoxide (DMSO) solvent
for stable dispersion, and 80/20 v/v isopropanol/water mixture as a coagulation bath [123]
Some of the considerations required during fabrication are that the spinning ink must
exhibit shear-thinning behavior for efficient flow, and an appropriate coagulation bath
should be chosen so that the extruded spinning ink is rapidly vulcanized without any
breakup [122]. The diameter of the fibers must be easily controlled by the nozzle size of
the spinneret [119]. The fibers intended to be used as strain sensors need to have a high
sensitivity for sensing tiny deformations during respiration, heartbeat, and speech recog-
nition, and a broad sensing range to detect large movements, such as finger bending,
walking, and running. In order to achieve both features, dip-coating of fibers has been
reported [124,125]. A summary of advantages and disadvantages of each of the fabrication
methods to create a flexible/stretchable device is presented in Table 3.
Figure 8. Alternative fabrication methods. (a,b) Printing method. (a). Schematics of printing of liq-
uid metal (eGaIn) on an elastomer base layer [60]. (i) Extrusion printing of a base elastomer layer
(ii) Spray printing of liquid metal (iii) Selective activation of the electrical path (iv) Extrusion print-
ing of an encapsulation elastomer to seal the device. Reprinted with permission from Ref. [60]. Cop-
yright 2017 Wiley-VCH Verlag GmbH & Co. (b). Fabrication steps for hydroprinted electronics.
Figure 8.
Alternative fabrication methods. (
a
,
b
) Printing method. (
a
). Schematics of printing of liquid
metal (eGaIn) on an elastomer base layer [
60
]. (i) Extrusion printing of a base elastomer layer (ii)
Spray printing of liquid metal (iii) Selective activation of the electrical path (iv) Extrusion printing of
an encapsulation elastomer to seal the device. Reprinted with permission from Ref. [
60
]. Copyright
2017 Wiley-VCH Verlag GmbH & Co. (
b
). Fabrication steps for hydroprinted electronics. Process
for application of a hydroprinted electronic tattoo over the forearm for EMG signal acquisition [
113
].
Reprinted with permission from Ref. [
113
]. Copyright 2018 American Chemical Society. (
c
,
d
) The
wet spinning method of fabrication. (
c
). Schematic illustration of wet spinning GO fibers. Wet
spinning set-up [
118
]. (
d
). Schematic illustration of the coaxial spinning process for highly stretchable
fibers [122]. Reprinted with permission from Ref. [122] Copyright 2018 American Chemical Society.

Materials 2022,15, 1664 20 of 36
Table 3. Summary of fabrication methods with their advantages and disadvantages.
Fabrication Method Advantages Disadvantages
Chemical vapor deposition (CVD)
Application: For creating thin films
Grow thin solid film on the substrate.
Grow large-area and high-quality
graphene film on metal substrate.
Need a metal (Ni, Cu) substrate to
grow graphene layer.
Cannot use polymer substrate to grow
graphene layer.
Challenge of growing large-grain
crystal, which is required to enhance
the electronic, mechanical, and
thermal properties.
Alternative: Modified CVD can be
used to produce large-grain and
highly crystalline film.
Crystal growth variation at different
areas.
Requires transfer of the 2D film from
the growth substrate to the target
stretchable device substrate to add the
stretchability property to the device.
Possible contamination and damage
during transfer of the film to the
target.
Plasma-enhanced CVD (PECVD)
Application: For creating thin films on
polymer substrate.
Potential for direct growth of
graphene layer on polymer substrate.
Utilizes low processing temperature,
providing an opportunity to use
polymer as substrate.
Inferior graphene properties.
Formation of cracks in the graphene
structure may occur if repeatedly
strained, leading to an increase in
resistance.
UV lithography
Application: For creating patterns for
circuits for microelectronics, sensing,
and optoelectronics. Can also be used
to create high-aspect-ratio structures
and 3D nanostructures such as
nanowires (A novel top-down
fabrication process for a
vertically-stacked silicon-nanowire
array).
Fabricate desired pattern ranging from
nm to µm size on substrate.
Soft and stretchable substrate can be
used.
Can pattern only photoresist and may
require a photomask for masked
lithography [126,127].
Can be expensive if infrastructure
such as a mask aligner is needed to
develop complicated structures [128].
Resolution of the projection optics is
diffraction-limited [126,129].
Challenging to design complex optics.
Use of harsh processing conditions
such as ion, plasma, acid, or
temperature treatment to remove
barrier layer, which limits the use of
the substrate or photoresist that cannot
handle such harsh processing [130].
Electron beam (E-beam) lithography
Application: For producing patterns in
a photomask
Produce a photomask.
High-resolution approach to creating a
photomask.
High resolution approach and
therefore, time consuming.
Not suitable for mass production.
Soft lithography
Application: Thin film fabrication on
soft substrates.
Micro/nano-structures can be
fabricated on a soft substrate.
Can be used to fabricate a mold.
Easy, fast, low-cost, mass reproducible
process.
Does not require harsh processing
chemicals.
Requires a stamp or a mold.
Printing
Application: For creating thin film
circuits of conductive liquids.
Can be used with conductive polymer
composites and conductive liquid
metals, thereby aiding in the
fabrication of flexible and stretchable
devices.
Fast, accurate, and fully automated
non-contact approach.
Repeatable and scalable.
Capability of achieving electronics
with complex patterns without
compromising the electrical
conductivity of devices.
Requires good choice of conductive
material and a solvent with
appropriate viscosity, surface tension,
and rheology.
Requires the sintering process to create
a conductive path.

Materials 2022,15, 1664 21 of 36
Table 3. Cont.
Fabrication Method Advantages Disadvantages
Wet spinning method
Application: For fabricating stretchable
and electrically conductive fibers for
wearable electronics
(strain sensors, supercapacitors,
nanogenerators).
Good technology for wearable electronics
as the fibers can withstand mechanical
deformations such as bending, folding,
twisting, and stretching.
A facile and viable strategy to fluidly spin
the macroscopic fiber in a continuous way.
Ability to control the diameter of the fiber
by choosing an appropriate nozzle size of
the spinneret.
Shear-thinning spinning ink is required
for efficient flow.
Requires selection of an appropriate
coagulation bath so that the extruded
spinning ink is rapidly vulcanized
without any breakup.
5. Transducers and Communication
As technology advances, the demand for new and innovative forms of wireless com-
munication increases. These new forms of communication are being incorporated within
polymers to help aid the medical field with progressive monitoring and tasks. They operate
under different dynamic ranges, standards, and levels of repeatability, noise, and hysteresis,
depending on the location of the device in use. Throughout this section, various methods
of data collection and transition are reviewed in terms of the different types of transducers
and how they enable communication. Transducers enable the conversion of one form of
energy into another—usually electrical—to allows the devices to communicate. Various
types of transducers also have different dynamic ranges [
131
]; the higher the range, the
more precise and sensitive the transducer. The use of sensors for communication has aided
the fields of biomechanics, clinical medicine, ergonomics, and biomedicine.
When dealing with wireless forms of communication, electrical conductivity is essen-
tial for high gain with efficient bandwidth. The material used also needs to have resistance
to various forms of degradation from the mechanical stress of use. The proper base material
is also important regarding the use of the device. Does the situation call for a flexible or
stretchable substrate? Although there are various forms of transducers and communication
methods, the following sections discuss the prominent forms used, in addition to examples
being used currently in the field of flexible and stretchable bioelectronics.
5.1. Piezoelectric Transducers
Since being first reported in 1880 [
132
], piezoelectricity has been used successfully in
the design of therapeutic systems and widely used due to its transducer properties. To
obtain an electrical charge, a force is first exerted on a given solid material. The stress upon
the given material causes a deformation in the positive and negative subatomic particles of
the material, causing a polarization that generates an electric field that is used to transform
mechanical energy into electrical energy [
133
]; that is, a molecule is unperturbed with
no polarization until an outside force exhibits stress upon the molecule. When stress is
applied, it causes movement within the electrons and results in polarization. When the
force is removed from the material, the polarization also stops. Due to its small size and
high-frequency response, and versatile flexibility, the piezoelectric transducer is easy to use
and excellent for clinical applications. This has resulted in therapeutic applications in the
fields of neurosurgery and cancer treatment. In 1949, Dr. George Ludwig first reported the
use of ultrasound in animal tissue [132] due to the piezoelectricity effect.
As a result of new and emerging biomechanical technologies, piezoelectric sensors
are used to measure either kinetics or kinematics. Applications for the use of piezoelectric
transducers include their use in recording the forces between different parts of the body
and support surfaces [
134
], uses in clinical settings, helping to monitor human performance,
or even their use for leisure purposes. In the 1980s, due to their low cost and lightweight
properties, piezoelectric transducers were used in the measurement of vibration and impact
in activities. Pedobarography has grown significantly, in part because of the assistance
of accelerometers and gyroscopes, which have enabled ambulatory measurements, thus

Materials 2022,15, 1664 22 of 36
leading to new advances in the study of gait and pressure while moving [
134
]. Some of the
first uses of piezoelectrical materials in this fashion were first described by Hennacy and
Gunther in 1975 in the measurement of dynamic pressure distribution [
135
]. Iqra Choudhry
et al. described a flexible piezoelectric transducer that is capable of harvesting energy
from human kinematics using flexible nanocomposite piezoelectric nanogenerators [
136
]
(Figure 9). The group designed a shoe-insole nanogenerator with the aim of detecting
human movement. In the design, movement generated power which, in turn, created a
frequency that was transmitted to a digital storage oscilloscope and recorded through data
acquisition (DAQ).
Materials 2022, 15, x FOR PEER REVIEW 23 of 38
with no polarization until an outside force exhibits stress upon the molecule. When stress
is applied, it causes movement within the electrons and results in polarization. When the
force is removed from the material, the polarization also stops. Due to its small size and
high-frequency response, and versatile flexibility, the piezoelectric transducer is easy to
use and excellent for clinical applications. This has resulted in therapeutic applications in
the fields of neurosurgery and cancer treatment. In 1949, Dr. George Ludwig first reported
the use of ultrasound in animal tissue [132] due to the piezoelectricity effect.
As a result of new and emerging biomechanical technologies, piezoelectric sensors
are used to measure either kinetics or kinematics. Applications for the use of piezoelectric
transducers include their use in recording the forces between different parts of the body
and support surfaces [134], uses in clinical settings, helping to monitor human perfor-
mance, or even their use for leisure purposes. In the 1980s, due to their low cost and light-
weight properties, piezoelectric transducers were used in the measurement of vibration
and impact in activities. Pedobarography has grown significantly, in part because of the
assistance of accelerometers and gyroscopes, which have enabled ambulatory measure-
ments, thus leading to new advances in the study of gait and pressure while moving [134].
Some of the first uses of piezoelectrical materials in this fashion were first described by
Hennacy and Gunther in 1975 in the measurement of dynamic pressure distribution [135].
Iqra Choudhry et al. described a flexible piezoelectric transducer that is capable of har-
vesting energy from human kinematics using flexible nanocomposite piezoelectric nano-
generators [136] (Figure 9). The group designed a shoe-insole nanogenerator with the aim
of detecting human movement. In the design, movement generated power which, in turn,
created a frequency that was transmitted to a digital storage oscilloscope and recorded
through data acquisition (DAQ).
Figure 9. The proposed shoe-insole nanogenerator that can generate an open circuit voltage of ~27
V. The sensors used for movement detection responded to almost every joint movement [136]. Re-
printed with permission from Ref. [136]. Copyright 2020 American Chemical Society.
A recent monitoring device was developed specifically using the piezoelectric effect.
This device, developed by Dagdeviren et al., is used for monitoring the vital signs and
ingestion within the GI tract [137]. The device offers the flexibility and durability needed
in the continually mocking GI tract. This unique device offers an interesting means of im-
plantation—the user swallows an encapsulated device. Once inside the stomach, the cap-
sule dissolves, releasing the monitoring device, which then unfolds and positions itself
within the mucosa. Due to the piezoelectric nature of the device, as the mucosa causes the
Figure 9.
The proposed shoe-insole nanogenerator that can generate an open circuit voltage of ~27 V.
The sensors used for movement detection responded to almost every joint movement [
136
]. Reprinted
with permission from Ref. [136]. Copyright 2020 American Chemical Society.
A recent monitoring device was developed specifically using the piezoelectric effect.
This device, developed by Dagdeviren et al., is used for monitoring the vital signs and
ingestion within the GI tract [
137
]. The device offers the flexibility and durability needed
in the continually mocking GI tract. This unique device offers an interesting means of
implantation—the user swallows an encapsulated device. Once inside the stomach, the
capsule dissolves, releasing the monitoring device, which then unfolds and positions itself
within the mucosa. Due to the piezoelectric nature of the device, as the mucosa causes the
device to move, electricity is developed. The device uses hook-up wires to transmit the
recorded movements via signals to a computer and a USB multimeter.
Based on the understanding of the need for stretchable devices, researchers have
utilized either intrinsic piezoelectric materials (polyvinylidene fluoride (PVDF), polyvinyli-
dene fluoride-trifluoroethylene (PVDF-TrFE) and odd nylon), piezoelectric electret, or
piezoelectric fillers (barium titanate (BTO), lead zirconate titanate (PZT), (1
−
x)Pb(Mg,
Nb)O
3−x
PbTiO
3
(PMN-PT)) in a stretchable substrate such as PDMS or Ecoflex [
138
]. Piezo-
electric electric materials exhibit piezoelectric effects due to the polarity generated from
charged air bubbles or space charges injected by a corona charger at a high DC voltage [
139
].
Because these piezoelectric materials are flexible when processed into thin films, but not
stretchable, researchers employ the method of structure configurations to provide the
flexible materials with stretchability. Sun et al. reported utilizing the kirigami approach to
fabricate stretchable strain sensors from flexible PVDF film encapsulated in PDMS [
140
].
The strain sensor could be used on curved surfaces, including the heart and body joints. Sim-
ilarly, a highly stretchable PVDF vibration-sensing e-tattoo for siesmocardiography (SCG)
monitoring was created utilizing serpentine interconnects on Tegaderm substrate [
141
].

Materials 2022,15, 1664 23 of 36
Enrico and team created a press sensor made of composite material with PDMS and PZT
fillers [142].
5.2. Ultrasonic Transducers
By interpreting reflected signals, similar to that of radar or sonar, ultrasonic transduc-
ers offer a low-power form of communication. Based on the properties of piezoelectric
transducers, ultrasonic transducers can convert alternating current into ultrasound, in
addition to reversing it. To produce an ultrasonic sound, piezoelectric crystals have an
electrical current applied to them so their shape changes; this causes a vibration that creates
sound waves in the same manner in which bats use echolocation. With the advent of
microelectromechanical systems, new capacitive micromachined ultrasonic transducers
(CMUTs) have emerged, helping replace some of the bulker piezoelectric transducers that
some traditional medical machines have relied on [
143
]. CMUTs offer little impedance of
the membrane, which vibrates to detect ultrasonic waves, and provide a wide bandwidth
and an excellent coupling efficiency. A unique use for CMUTs was developed by Ramanavi-
ciene et al., who created a biosensor that is able to sense the change in mass loading on the
top of a membrane, which was then used to detect bovine leukemia virus proteins [
144
].
The device read the normal frequency of the CMUT and, with any deviation caused by the
change in device, the results were then recorded and transmitted.
Based on the advantages of CMUTs, a novel transparent flexible CMUT was devel-
oped using a roll-lamination technique with properties that demonstrated transparency,
flexibility, and non-contact detection in display panels that offer a better human-to-machine
interface (Figure 10) [
145
]. This CMUT was developed on a flexible substrate at a tempera-
ture below 100
◦
C to help reduce the stress of high temperatures. The developed CMUT
is able to transmit pulse echoes, at various distances, which are then recorded given the
frequency at which the echoes are transmitted.
Materials 2022, 15, x FOR PEER REVIEW 25 of 38
Figure 10. Flexible CMUT step process by Pang et al. [145].
5.3. Wearable Antenna Communication
Wearable antennas are specifically designed to transmit and receive soft radio fre-
quencies while being worn. This area of communication has been widened with the ad-
vent of wireless body area networks, which provide applications both within the medical
and non-medical fields, such as medical sensing, body movement detection, skin moni-
toring, and other monitoring devices. Antennas help establish communication by linking
the wearable device to a targeted device. Two possibilities exist for wearable devices: off
the body and on the body. Off the body includes items such as bracelets, whereas on the
body includes clothes, medical tape, and bandages. Some notable examples of antennas
in various fields are fabric-based embroidered antennas, microfluidic antennas with injec-
tion alloys, and polymer-embedded antennas. Embroidered antennas include antennas
such as Metal Composite Embroidery Yarn, Shieldex thread, carbon nanotubes, and Am-
berstrand fibers. Examples of microfluidic antennas are EGaIn and Galinstan. Zhibo Chen
et al. showed the creation of a stretchable elastomer composed of Ag-PDMS for radio fre-
quency passive components for various wireless wearable communication applications
[148]. A wearable antenna is often the key component because it is responsible not only
for receiving, but also transmitting, the signal used for the network of the device [149,150].
Table 4 provides details of the known frequency bands and the use of those frequencies
in terms of applications [150].
Table 4. A wide variety of frequencies used when dealing with networks and antennas in wearable
devices [150].
Frequency Application
401–402 MHz
Medical Implant Communication Services
(MICS).
402–405 MHz
403.5–403.8 MHz
405–406 MHz
413–457 MHz Medradio Micropower Networks (MMNs),
transmit and relay data for implanted and
Figure 10. Flexible CMUT step process by Pang et al. [145].
Efforts have also been made to realize stretchable ultrasonic transducers. Most of the
fabrication techniques for realizing stretchable ultrasonic transducers include using (1) a
structural design for stretchable circuit patterning or a piezoelectric composite; (2) transfer
printing; and (3) soft elastomeric packaging [
146
]. An ultrasonic probe with an island-
bridge layout for biaxial stretchability was reported to have more than 50% stretchability, a
high signal-to-noise ratio, a wide bandwidth, and a high spatial resolution to allow imaging
through complex surfaces [
147
]. A similar device developed by Wang et al. to monitor the

Materials 2022,15, 1664 24 of 36
central blood pressure waveform was reported to be ultrathin, and to have stretchability of
up to 60% with high sensitivity [146].
5.3. Wearable Antenna Communication
Wearable antennas are specifically designed to transmit and receive soft radio frequen-
cies while being worn. This area of communication has been widened with the advent
of wireless body area networks, which provide applications both within the medical and
non-medical fields, such as medical sensing, body movement detection, skin monitoring,
and other monitoring devices. Antennas help establish communication by linking the
wearable device to a targeted device. Two possibilities exist for wearable devices: off the
body and on the body. Off the body includes items such as bracelets, whereas on the
body includes clothes, medical tape, and bandages. Some notable examples of antennas in
various fields are fabric-based embroidered antennas, microfluidic antennas with injection
alloys, and polymer-embedded antennas. Embroidered antennas include antennas such as
Metal Composite Embroidery Yarn, Shieldex thread, carbon nanotubes, and Amberstrand
fibers. Examples of microfluidic antennas are EGaIn and Galinstan. Zhibo Chen et al.
showed the creation of a stretchable elastomer composed of Ag-PDMS for radio frequency
passive components for various wireless wearable communication applications [
148
]. A
wearable antenna is often the key component because it is responsible not only for receiv-
ing, but also transmitting, the signal used for the network of the device [
149
,
150
]. Table 4
provides details of the known frequency bands and the use of those frequencies in terms of
applications [150].
Table 4.
A wide variety of frequencies used when dealing with networks and antennas in wearable
devices [150].
Frequency Application
401–402 MHz
Medical Implant Communication Services (MICS).
402–405 MHz
403.5–403.8 MHz
405–406 MHz
413–457 MHz
Medradio Micropower Networks (MMNs), transmit and
relay data for implanted and body-worn medical devices
for diagnostics and therapeutic functions.
2.4 and 5 GHz Wi-Fi, smart hospital beds, mobile nursing stations.
2.404–2.478 GHz
Bluetooth, indoor navigation for patients, connectivity
between the device and the smartphone for heath-data
monitoring.
2.36–2.4 GHz New Medical BAN (MBAN).
New Medical BAN (MBAN) Industrial, Scientific, and Medical (ISM).
3.1–10.76 GHz Ultra-Wideband (UWB).
57–64 G The band plans and rules at mmW-60 BAN.
59–66 GHz
5.4. Wireless Communication
The use of wireless devices for communication and receiving information is not only
limited to cellphones, but is now rising in significance within medical technology, from
operating room equipment to home-healthcare devices that can be used for mobile phones.
The use of interconnected wires limits the range of communication, in addition to posing
an issue in regard to processing information and data collection. In the process of creating
wireless forms of communication within biomedical devices, the general cost of creation
and their efficiency of communication have been investigated. The devices are advancing
from bulky models into soft deformable models that offer low cost and comfort for the user,

Materials 2022,15, 1664 25 of 36
combined with the freedom of not having to use bulky wires. A stretchable conductive
elastomer was created that not only offers wireless communication, but also offers low-
cost fabrication [
148
]. Zhibo Chen et al. designed a soft and deformable device with a
stretchable conductor made from a mixture of Ag and PDMS (Figure 11). This mixture
offered a high strain along with conductivity that enabled quality RF passive components
for wireless communication. This device was designed to be mounted on the skin, with a
flexible antenna having frequency ranges between 500 MHz and 3 GHz. The group’s data
showed that the Ag-PDMS elastomer with a wireless form of communication was able to
resist the signal loss of the RF passive front-end.
Materials 2022, 15, x FOR PEER REVIEW 27 of 38
Figure 11. The Ag-PDMS sample with transmission line adapters seen in (a), (b) the tensile test
fixture (c) graphing the different Ag volumes in comparison to the attenuation of transmission line,
and (d) showing the attenuation of the transmission [148].
6. Power Sources
Throughout this article, we reviewed various flexible and stretchable electronics,
their fabrication methods, and communication techniques. In addition, stretchable energy
sources play an equally salient role in any kind of flexible and stretchable electronic de-
vice. Developing these energy sources presents some rewarding challenges. For an energy
source to be used in flexible and stretchable electronics, several things must be considered,
including the compliance of the material, the mechanical endurance under repetitive load,
the amount of energy that it can provide to satisfy the application requirements, and the
safety of the device. For continuous power generation and supply, researchers have ex-
plored the use of energy-harvesting components based on piezoelectric, triboelectric, pho-
tovoltaic, and thermoelectric mechanisms of power generation. Briefly, piezoelectric en-
ergy harvesters rely on the induced strain or pressure for electricity generation, as dis-
cussed previously. Triboelectric energy harvesters utilize the physical contact and sepa-
ration between two dissimilar materials resulting from the biomechanical motion of the
human body to generate electricity [151]. Photovoltaics cells have also been incorporated
to convert light energy into electrical energy in self-powered ultra-flexible devices
[152,153]. Similarly, thermoelectric energy generators have also been widely used in flex-
ible devices to generate electrical power using the human body as the heat source
[154,155]. Batteries and supercapacitors are the most widely used power sources in
stretchable bioelectronics. Hence, we discuss these in further detail.
Batteries store energy via an electrochemical process. Some batteries, such as Li-ion
batteries, are widely used due to their design flexibility, ability to store large quantities of
energy, and long-term cycling life [156]. The drawback of these batteries is they contain a
non-aqueous electrolyte, from which arises the problem of flammability. When develop-
ing safe, stretchable batteries, one should also recognize that deformation may cause an
internal short circuit, thus leading to explosion. By comparison, aqueous batteries demon-
strate safety, high ionic conductivity, and cost effectiveness, and are therefore an ideal
Figure 11.
The Ag-PDMS sample with transmission line adapters seen in (
a
), (
b
) the tensile test
fixture (
c
) graphing the different Ag volumes in comparison to the attenuation of transmission line,
and (d) showing the attenuation of the transmission [148].
6. Power Sources
Throughout this article, we reviewed various flexible and stretchable electronics,
their fabrication methods, and communication techniques. In addition, stretchable energy
sources play an equally salient role in any kind of flexible and stretchable electronic device.
Developing these energy sources presents some rewarding challenges. For an energy
source to be used in flexible and stretchable electronics, several things must be considered,
including the compliance of the material, the mechanical endurance under repetitive load,
the amount of energy that it can provide to satisfy the application requirements, and
the safety of the device. For continuous power generation and supply, researchers have
explored the use of energy-harvesting components based on piezoelectric, triboelectric,
photovoltaic, and thermoelectric mechanisms of power generation. Briefly, piezoelectric
energy harvesters rely on the induced strain or pressure for electricity generation, as
discussed previously. Triboelectric energy harvesters utilize the physical contact and
separation between two dissimilar materials resulting from the biomechanical motion of the

Materials 2022,15, 1664 26 of 36
human body to generate electricity [
151
]. Photovoltaics cells have also been incorporated to
convert light energy into electrical energy in self-powered ultra-flexible devices [
152
,
153
].
Similarly, thermoelectric energy generators have also been widely used in flexible devices
to generate electrical power using the human body as the heat source [
154
,
155
]. Batteries
and supercapacitors are the most widely used power sources in stretchable bioelectronics.
Hence, we discuss these in further detail.
Batteries store energy via an electrochemical process. Some batteries, such as Li-ion
batteries, are widely used due to their design flexibility, ability to store large quantities of
energy, and long-term cycling life [
156
]. The drawback of these batteries is they contain a
non-aqueous electrolyte, from which arises the problem of flammability. When developing
safe, stretchable batteries, one should also recognize that deformation may cause an internal
short circuit, thus leading to explosion. By comparison, aqueous batteries demonstrate
safety, high ionic conductivity, and cost effectiveness, and are therefore an ideal alternative
to conventional batteries. An approach of incorporating hybrid carbon filler/polymer
(HCP) composite as a current collector to fabricate aqueous Li-ion batteries was presented
by the Song group (Figure 12a) [
156
]. They successfully demonstrated that such batteries
provide a maximum power of 1260 W kg
−1
and possess an excellent capacity retention
of 93% after 500 charge–discharge cycles. Additionally, the HCP in these batteries is
capable of straining up to 200%. Another similar effort was made by Liu’s team, in
which they introduced liquid metal alloy, i.e., eGaIn-MnO
2
, rechargeable batteries for
stretchable electronics [
157
]. The battery has a 3.76 mA h cm
−2
specific areal capacity
when fully discharged, exhibits stable discharge voltage within 100 cycles at 0.4 mA cm
−2
,
and supports tensile strain of 100% (Figure 12b). In addition to employing the concept of
aqueous batteries for stretchable bioelectronics, researchers have also implemented designs
using helical spring-like and serpentine structures as a structural support and backbone for
battery components. Zamarayeva et al. successfully implemented these designs to fabricate
battery sources that have high specific capacity (~1.25 mA h cm
−1
) when charged at 0.25 C
and discharged at 0.5 C [
158
]. Although the helical spring batteries showed that they can
withstand flexing over 17,000 times to the bending radius of 0.5 cm, they found that these
batteries cannot be stretched due to the mechanical properties of the polymer electrolyte,
cellophane layer, and silver electrode used for the battery. The serpentine-shaped batteries
overcame this issue because they can be readily stretched and can accommodate motions
in a biaxial direction while retaining their electrochemical performance when stretched to
100% (Figure 12c).
Supercapacitors can survive millions more charging cycles than a rechargeable battery.
However, they charge and discharge rapidly with massive bursts of energy while discharg-
ing. Generally, carbon nanomaterials such as CNTs and graphene have been explored for
use in supercapacitors due to their low cost, high surface area, and long cycle lifetime [
159
].
Chen et al. demonstrated that highly stretchable supercapacitors with high electrochemical
performance can easily be fabricated using a method involving the wrapping of CNT thin
film synthesized by chemical vapor deposition around pre-stretched elastic wires, and
twisting two of these CNT-wrapped elastic wires precoated with poly(vinyl alcohol) pow-
der (PVA)/H
3
PO
4
/water hydrogel as the electrolyte (Figure 13a(i)) [
160
]. A PEDOT-PSS
coating on CNT-wrapped elastic wires exhibited improved conductive and capacitive per-
formance (Figure 13a(iii)). Supercapacitors based on CNT-wrapped wire showed excellent
capacitance up to 30.7 F g
−1
and high stretchability of up to a strain of 350%. Cao and
Zhou’s team demonstrated a facile approach of fabricating crumpled CNT forest-based
supercapacitors on a pre-strained elastomer substrate (Figure 13c) [
159
]. Such a crumpled
CNT forest has the advantages of easy access due to its pore structure, short ion transport
time, and low ionic diffusion resistance. With the deposition of Au on the CNT forest, they
were able to demonstrate superior electrochemical performance (specific capacitance of
∼
26 mF cm
−2
for a biaxially stretchable Au-CNT forest) under a large deformation of 800%
area strain regardless of charging and discharging rates.

Materials 2022,15, 1664 27 of 36
Materials 2022, 15, x FOR PEER REVIEW 28 of 38
alternative to conventional batteries. An approach of incorporating hybrid carbon
filler/polymer (HCP) composite as a current collector to fabricate aqueous Li-ion batteries
was presented by the Song group (Figure 12a) [156]. They successfully demonstrated that
such batteries provide a maximum power of 1260 W kg
−1
and possess an excellent capacity
retention of 93% after 500 charge–discharge cycles. Additionally, the HCP in these batter-
ies is capable of straining up to 200%. Another similar effort was made by Liu’s team, in
which they introduced liquid metal alloy, i.e., eGaIn-MnO
2
, rechargeable batteries for
stretchable electronics [157]. The battery has a 3.76 mA h cm
−2
specific areal capacity when
fully discharged, exhibits stable discharge voltage within 100 cycles at 0.4 mA cm
−2
, and
supports tensile strain of 100% (Figure 12b). In addition to employing the concept of aque-
ous batteries for stretchable bioelectronics, researchers have also implemented designs
using helical spring-like and serpentine structures as a structural support and backbone
for battery components. Zamarayeva et al. successfully implemented these designs to fab-
ricate battery sources that have high specific capacity (~1.25 mA h cm
−1
) when charged at
0.25 C and discharged at 0.5 C [158]. Although the helical spring batteries showed that
they can withstand flexing over 17,000 times to the bending radius of 0.5 cm, they found
that these batteries cannot be stretched due to the mechanical properties of the polymer
electrolyte, cellophane layer, and silver electrode used for the battery. The serpentine-
shaped batteries overcame this issue because they can be readily stretched and can accom-
modate motions in a biaxial direction while retaining their electrochemical performance
when stretched to 100% (Figure 12c).
Figure 12. Stretchable batteries for powering bioelectronics. (a). Schematic showing a stretchable
aqueous batteries configuration in which the HCP composite was used as a current collector. Long-
term cycle performance and coulombic efficiency of the full cell at a rate of 20 C over 500 cycles
Figure 12.
Stretchable batteries for powering bioelectronics. (
a
). Schematic showing a stretchable
aqueous batteries configuration in which the HCP composite was used as a current collector. Long-
term cycle performance and coulombic efficiency of the full cell at a rate of 20 C over 500 cycles [
156
].
Reprinted with permission from Ref. [
156
]. Copyright 2018 Wiley-VCH Verlag GmbH &Co. (
b
).
Schematic showing the composition of a stretchable EGaIn battery [
157
] A. Photographs of the stretch-
able EGaIn-MnO
2
battery array in series of two stretched under 0, 50, and 100% strain integrated with
LEDs. B. Photograph of battery-powered strain sensor that is mounted on the wrist. Reprinted with
permission from Ref. [
157
]. Copyright 2019 Wiley-VCH Verlag GmbH & Co. (
c
). Schematics of simple
serpentine current collector and optical images of four serpentine-shaped batteries connected in series.
Batteries continuously power an OLED while being subjected to a uniaxial strain of 100%. Schematics
of a self-similar serpentine current collector and optical images of the full battery assembled around
such a current collector. Geometry of the battery facilitates biaxial stretching [158].

Materials 2022,15, 1664 28 of 36
Materials 2022, 15, x FOR PEER REVIEW 30 of 38
Figure 13. Supercapacitors for powering bioelectronic devices. (a) (i) Schematic illustration of fabri-
cating stretchable conducting wire by wrapping an aligned CNT sheet around a pre-stretched elastic
wire. CV curves of the supercapacitors based on (ii) the bare CNT-wrapped and (iii) CNT/PEDOT-
PSS-wrapped wires [160]. (b) Digital photograph of a typical wire-shaped supercapacitor with a
twisted structure after being stretched from strains of 0 to 370% [160]. Reprinted with permission
from Ref. [160]. Copyright 2015 Wiley-VCH Verlag GmbH & Co. (c). Stretchable Au-CNT forest
electrodes: (i) SEM image of the Au-CNT forest pattern morphology generated by a uniaxial pre-
strain of 300% and by applying a biaxial pre-strain of 200% × 200%. (ii) Capacitance retention of a
uniaxially stretchable Au-CNT forest electrode under mechanical stretching–relaxation cyclic defor-
mations to a strain of 200% and for 10,000 charge/discharge cycles at the relaxed state. Inset shows
the CV curves measured before and after the electrochemical stability test at the scan rate of 500 mV
s
−1
[159]. Reprinted with permission from Ref. [159]. Copyright 2020 Elsevier Inc. (d). Digital images
of stretchable rectangular-shaped supercapacitors (with geometric parameters of y = 0.7 cm, m = 0.2
cm, x = 194 µm, T = 0.5 cm) under different strain tests. The inset images (upper left) are the scheme
showing the expandable honeycomb structure and the hexagonal unit cell before and after being
stretched. Capacitance retention ratio of 3D stretchable supercapacitor based on PPy/BPO-CNT elec-
trodes tested at 7.8 mA cm
−2
under the cycling tensile strain of 2000% [161]. (e). Arched bridge-
shaped supercapacitors acting as a 3D helmet worn on the head of an owl toy model to power a 3.0
V flexible LED strip (right) [161]. Reprinted with permission from Ref. [161]. Copyright 2018 Wiley-
VCH Verlag GmbH & Co.
7. Conclusions
This review paper focused on various aspects required for the fabrication of a fully
functional stretchable bioelectronic device. Therefore, it discussed several flexible and
stretchable polymers, their mechanical and biocompatible properties, different strategies
for making a device that is not only flexible but also stretchable, signaling and communi-
cating methods, and stretchable power sources. The material section does not contain an
exhaustive list of flexible and stretchable biomaterials. Those listed in Section 2 were only
common biomaterials already deemed biocompatible with a multitude of applications.
These materials are highly moldable. This area of research is moving fast and adapting.
In fact, the use of organic materials such as proteins is a growing area for bioelectronics
and offers better biocompatibility than polymers or metallic-based materials. Next, we
discussed the common fabrication methods for manufacturing flexible and stretchable bi-
oelectronic devices. The flexibility and stretchability of the device not only depends on the
Figure 13.
Supercapacitors for powering bioelectronic devices. (
a
) (i) Schematic illustration of fabri-
cating stretchable conducting wire by wrapping an aligned CNT sheet around a pre-stretched elastic
wire. CV curves of the supercapacitors based on (ii) the bare CNT-wrapped and (iii) CNT/PEDOT-
PSS-wrapped wires [
160
]. (
b
) Digital photograph of a typical wire-shaped supercapacitor with a
twisted structure after being stretched from strains of 0 to 370% [
160
]. Reprinted with permission
from Ref. [
160
]. Copyright 2015 Wiley-VCH Verlag GmbH & Co. (
c
). Stretchable Au-CNT forest
electrodes: (i) SEM image of the Au-CNT forest pattern morphology generated by a uniaxial pre-
strain of 300% and by applying a biaxial pre-strain of 200%
×
200%. (ii) Capacitance retention
of a uniaxially stretchable Au-CNT forest electrode under mechanical stretching–relaxation cyclic
deformations to a strain of 200% and for 10,000 charge/discharge cycles at the relaxed state. Inset
shows the CV curves measured before and after the electrochemical stability test at the scan rate of
500 mV s
−1
[
159
]. Reprinted with permission from Ref. [
159
]. Copyright 2020 Elsevier Inc. (
d
). Digital
images of stretchable rectangular-shaped supercapacitors (with geometric parameters of y = 0.7 cm,
m = 0.2 cm, x = 194
µ
m, T = 0.5 cm) under different strain tests. The inset images (upper left) are the
scheme showing the expandable honeycomb structure and the hexagonal unit cell before and after
being stretched. Capacitance retention ratio of 3D stretchable supercapacitor based on PPy/BPO-
CNT electrodes tested at 7.8 mA cm
−2
under the cycling tensile strain of 2000% [
161
]. (
e
). Arched
bridge-shaped supercapacitors acting as a 3D helmet worn on the head of an owl toy model to power
a 3.0 V flexible LED strip (right) [
161
]. Reprinted with permission from Ref. [
161
]. Copyright 2018
Wiley-VCH Verlag GmbH & Co.
Most of the supercapacitors fabricated on pre-stretched elastomers or textiles are 2D
based. These supercapacitors have the limitations of the need for thin electrodes to achieve
higher mechanical strain. This requirement prevents these supercapacitors from active
material loading in the vertical direction, thereby resulting in low areal capacitance power.
With this in mind, Lv et al. proposed honeycomb-lantern-inspired 3D supercapacitors
(Figure 13d) [
161
]. They utilized the stretchability property of doped polypyrrole (PPy).
To enhance stress relief and specific areal capacitance via effective ion transfer, they elec-
trodeposited porous PPy and black phosphorus oxide (BPO) composites on CNT films.
With this approach, they were successful in achieving supercapacitors that could maintain
a capacitance ratio of 95% even when stretched at 2000% after 10,000 stretch and release
cycles (Figure 13d), and had an enhanced specific areal capacitance of 7.34 F cm−2.

Materials 2022,15, 1664 29 of 36
7. Conclusions
This review paper focused on various aspects required for the fabrication of a fully
functional stretchable bioelectronic device. Therefore, it discussed several flexible and
stretchable polymers, their mechanical and biocompatible properties, different strategies
for making a device that is not only flexible but also stretchable, signaling and commu-
nicating methods, and stretchable power sources. The material section does not contain
an exhaustive list of flexible and stretchable biomaterials. Those listed in Section 2were
only common biomaterials already deemed biocompatible with a multitude of applications.
These materials are highly moldable. This area of research is moving fast and adapting.
In fact, the use of organic materials such as proteins is a growing area for bioelectronics
and offers better biocompatibility than polymers or metallic-based materials. Next, we
discussed the common fabrication methods for manufacturing flexible and stretchable
bioelectronic devices. The flexibility and stretchability of the device not only depends
on the material properties of the substrate, but also on the mechanical properties of the
electrodes or the conducting pathways to transmit the signal. Flexibility and stretchability
in bioelectronic devices can be achieved using conductive polymer composites, the use of
different structural configurations, such as serpentine geometry, wrinkled nanomembranes,
or wavy structures, or the use of conductive liquids. Sensing the biochemical, biomechan-
ical, or bioelectric cues of the human body, and communicating with an external data
acquisition device, is an equally important aspect of the bioelectronic device. Piezoelectric
devices, ultrasonic transducers, wearable antennas, and wireless forms of communication
are a few of the trending methods of sensing and communication in the field of flexible and
stretchable bioelectronics. Finally, we talked about the two major types of power sources
(i.e., batteries and supercapacitors) that are widely used for stretchable bioelectronics.
Flexible/stretchable bioelectronics have attracted significant interest in health care
applications as an alternative for bulky health-monitoring devices. Although the advances
discussed above are noteworthy and provide a solid foundation for new classes of in-
telligent flexible/stretchable bioelectronics, several challenges remain to be addressed.
Most of the implants or devices for chronic use are encapsulated by a polymer to create
a barrier from biological fluids because these fluids can increase the rates of degradation
and corrosion, or cause biofouling. This not only increases the thickness of the device and
alters its mechanical properties, but also poses a challenge in the field of biosensing when
monitoring the interstitial fluid [
162
]. Flexible and stretchable devices are an excellent
choice when the device needs to be implanted on a flat surface, such as the skin or the heart,
for constant monitoring of biosignals. However, what if the device has to be implanted
on a rough surface or curvatures such as those of the brain or gastrointestinal (GI) tract?
Due to this need for the conformability of the devices, the polymer or substrate must be
conformable, so that the device adopts the curvature of the human body part where it is
patched or implanted. Device conformability also ensures that the device is able to acquire
changes in physiological cues or record bodily signals. As a result of the advance in highly
stretchable and conformable substrates for a better interface with the human body arises
the necessity of improved manufacturing technologies, such as low-temperature processes.
As a hybrid approach or hard–soft material integration would lead to unwanted stress,
localized strain, and hot spots in contact with human tissue [
163
], much effort is still needed
in expanding the micro- and nano-fabrication techniques to accommodate the integration
of hard semiconductors on soft, flexible, stretchable, and conformable devices. In addition
to the available conductive inks/polymers, exploration of biocompatible and bioresorbable
conductive inks/polymers is an active area of research. Furthermore, for most fabrication
processes, crack formation has been observed in electrical conduits upon repeated cycles of
straining the devices, and a tradeoff exists in mechanical and electrical properties in the case
of using conductive polymer composites. These factors lead to a decrease in the sensitivity
of the device; thus, research can be conducted to improve the device’s conductivity while
maintaining its capability of sustaining high mechanical deformation. As batteries offer
a high energy density and are an easy deployment option, efforts can be undertaken not

Materials 2022,15, 1664 30 of 36
only to reduce their size and increase their capacity [
164
], but also in creating batteries that
are flexible and biodegradable [
165
]. Future flexible and stretchable bioelectronic devices
can also be integrated with sensor arrays, on-site signal processing circuitry, and sensor
calibration mechanisms for accurate signal analysis [
166
]. However, the integration of
electronic components, especially power sources, into flexible/stretchable devices where
polymer is often used as a substrate or a form of encapsulation has a disadvantage, namely
that electronic components, including power sources, generate heat [
167
]. As the human
body is sensitive to heat, efforts can be made to explore and develop a good thermally
conductive material or a heat sink for bioelectronics [
167
]. If these challenges can be ad-
dressed, the next generation of bioelectronics will constitute a flexible, stretchable, and
intelligent/automated multifunctional diagnostic and therapeutic bioelectronic system.
Author Contributions:
Conceptualization, M.E. and C.C.; investigation, C.C., E.H. and L.A.; data
curation, C.C., E.H. and L.A.; writing—original draft preparation, C.C., E.H. and L.A.; writing—
review and editing, M.E., C.C., E.H. and L.A.; supervision, M.E. project administration, M.E. and C.C.
All authors have read and agreed to the published version of the manuscript.
Funding: This research received no external funding.
Institutional Review Board Statement: Not applicable.
Informed Consent Statement: Not applicable.
Data Availability Statement: The study did not report any data.
Conflicts of Interest: The authors declare no conflict of interest.
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