**3.3 Enzyme-related aspects**

The selection of enzymes is a primary subject which should be discussed. Enzymes must be selected by considering their particular reactions to target

*Wearable Devices - The Big Wave of Innovation*

**3. Challenges and possible solutions**

electrodes for glucose biosensors [40].

**3.2 Powering wearable devices**

**3.1 Mechanical properties**

tance value.

relied on an organic polyterthiophene semiconductor, which drove a reduction reaction under illumination (wavelengths of 350 nm to over 600 nm). This system presented an

Additional efforts have been made to explore new biomedical applications of BFCs. **Figure 3F** shows an integrated fructose/O2 BFC patch that was conjugated with transdermal iontophoresis [33]. The current generated by the BFC was used to drive an osmotic flow from the anode to the cathode, resulting in the net ionic movement of small-molecule drug into the skin. The level of transdermal current to control the drug administration could be adjusted by connecting a thin poly(3,4-ethylenedioxythiophene)/PU resistor of a programmable resis-

Young's modulus of the human skin is in a range of 10–500 kPa [35, 36], while the moduli of common electronic materials, such as silicon and gold, are much higher (high GPa), indicating significant mechanical mismatch when integrating with the skin. Therefore, functionalities of non-stretchable electrodes will deteriorate after multiplex deformations commonly experienced by daily life activities. Furthermore, such rigidity and bulkiness of traditional devices also restrict the wearability and comfortability [14]. Non-compliant electrochemical devices will limit continuous long-term functions due to cracking and increasing of material resistance. This increasing of resistivity, which opposes the current flow in bioelectronics, causes poor electron communication at the enzyme-electrode interface. This major challenge of skin-integrated electronics can be addressed by exploring stretchable materials which display mechanical properties in a similar range of skin's modulus. One approach is using polymers due to their low mechanical toughness. For example, conducting materials with high moduli can be blended with soft polydimethylsiloxane or Ecoflex materials (Young's moduli of 0.4–3.5 MPa and 125 kPa, respectively) in order to tune the mechanical properties while keeping good electrochemical functions [37]. CNT-based materials, which are powerful for electrochemical devices [38], are used to combine with soft elastomers, such as PU and styrene-butadiene-styrene (SBS) [29, 39]. PU and SBS composites have moduli of ~700–800 kPa. As shown in **Figure 3C**, CNT filler (with the highaspect ratio ∼1300) was combined with PU [30], achieving stretchable conductive electrode materials. The percolation of dispersed CNTs can facilitate the electric flow in stretchable bioelectronics. Combining the intrinsic stretchability of this engineered inks with the structural stretchability of the serpentine design allows the device to tolerate strains as high as 500% with a small effect on its electrochemical performance [29]. This concept can be expanded by adding new functionalities into electrodes. For example, platinum-decorated graphite was mixed with PU to obtain stretchable electrocatalytic materials, allowing the fabrication of stretchable

Growing demand of wearable technologies has stimulated the need of the development of viable energy sources. The lack of anatomically power sources becomes a key bottleneck for the progress in wearable bioelectronics. Skin-worn bioelectronics mandates the compliant and efficient energy sources to supply multitasks, including

attractive example of on-skin autonomous power sources and sensors.

**38**

analytes or biofuels for electroanalytical monitoring and energy harvesting, respectively. One of the most predominant enzymes used to develop wearable bioelectronics is GOx from *Aspergillus niger*. It represents an example of commercially available biocatalyst that has good stability, substrate specificity, and electron turnover rate [3, 4, 45]. It is a powerful biorecognition element for glucose biosensors, the most widely interesting devices for diabetes health management. As shown in **Figure 4** (A–C), the enzyme is immobilized on the electrode, establishing a biosensor. GOx contains two 80 kDa subunits. Each holds a tightly bound flavin adenine dinucleotide (FAD) cofactor, the important redox center which has a redox potential −0.32 V (vs Ag/AgCl) at pH 7. This redox center is crucial to transfer electrons and specifically oxidize β-D-glucose to gluconolactone. However, this FAD is shielded by the protein and a glycan structure, hindering electron exchange at the enzyme-electrode interface. Inevitably, this requires research efforts to address this roadblock [46, 47]. FAD plays an important role as a common cofactor for glucose oxidation biocatalysis. The redox process for FAD/FADH2, involving two electrons, is shown in **Figure 4D**, where the R group represents adenosine diphosphate and ribitol connected with the flavin. However, it is O2-dependent; accordingly, O2 fluctuations can vary the performance of this type of oxidase-based bioelectronics. Although alternative O2-independent electrodes utilized NAD-dependent electrodes can be used, they require a diffusional cofactor, not simple for wearable applications. Hence, FAD-dependent dehydrogenases are becoming interesting choices since they are O2-independent and do not depend on diffusional mediators [48, 49].

The first generation of biosensors relies on quantifying O2 generation or H2O2 depletion (**Figure 4A**). This leads to key drawbacks, such as low dynamic range, dependency to oxygen fluctuations, and interfering effects. For instance, for glucose amperometric sensors, the detection of H2O2 at common first-generation electrodes needs the high applied detection potential where interfering compounds existing in sweat, e.g., ascorbic acid, uric acid, and some drugs, are also electroactive. Lowering the applied potential for the detection is a strategy to minimize

#### **Figure 4.**

*Principles of interfacing the enzyme, such as glucose oxidase (GOx), with the electrode. Different generations of strategies (A–C: first, second, and third generations) are illustrated. (D) Reactions involving the glucose oxidase biocatalyst.*

**41**

*Wearable Skin-Worn Enzyme-Based Electrochemical Devices: Biosensing, Energy Harvesting…*

such electroactive interferences. One approach is to incorporate electrocatalysts in wearable electrodes, such as PB or Pt [17, 40]. This offers low-potential detection of

First, the MET strategy utilizes a redox mediator, acting as an electronshuttle assistant between the enzymatic active center and the electrode. The substrate level, such as glucose, can then be monitored by the redox process of the mediator. This results in the independence of oxygen and mitigating the interfering signals due to the operation at low potentials. The first consideration in electrically wiring the enzyme with the electrode is the choice of the mediator that should be close to the redox potential of the active center of the enzyme to facilitate efficient electron communication between the enzyme and the conductive electrode surface. In particular, for enzymatic BFCs, the selection of mediators is crucial to positively control the cell voltage and enhance heterogeneous electron transfer to the order of a homogeneous transfer [50]. However, challenges of using mediators, particularly for BFCs, are their stability and deviated cell voltage. In addition, biocompatibility is highly vital for skin-worn applications. In spite of the assistance of electron shuttle by redox mediators, major concerns are their biocompatibility. One possible solution is employing nanomaterials or highly biocompatible catalysts. For example, mushroom/plant extracts could be used to obtain efficient "green" bioelectrocatalytic reactions

Second, direct electron transfer is an ideal goal of electrical wiring. It can be achieved by employing nanomaterials which suggest the direct electron transfer between enzyme active site and electrode. This wiring strategy is based on the shortening of the electronic contact of the enzyme and electrode (a short distance of ~1.5 nm) where the redox center of the enzyme can be regenerated directly by the electrode [52]. Therefore, this strategy can maximize the performance of bioelectronics. The engineering needs to consider the position of the active site inside the protecting protein and the conformation of the protein in order to wire the conducting materials with the redox center. This still remains the most challeng-

Several variables also affect the response nature of enzyme bioelectronics. Consideration of the fundamental theory of their functions will help to improve their performances. A key well-known model of enzyme behaviors is Michaelis-

the reaction, the maximal rate of the reaction, the Michaelis-Menten constant, and the concentration of the substrate, respectively. In general, it is desirable to engineer the biointerface electrode system to obtain high *Vm* and low *Km* (good affinity). However, dynamic range is also a crucial characteristic for wearable biosensors. Traditionally, dilution or preconcentration can be used to adjust the level of the target to be fit in the linear range of the sensor; nonetheless, manipulating such processes for on-skin applications is sophisticated. Therefore, diffusion-limiting membranes may be a useful solution to tune the dynamic range. The linear range can be extended by coating a thin membrane over an active enzyme layer since the sensor response is controlled by the analyte diffusion and not by the nonlinear characteristic of enzyme kinetics. Nevertheless, it should be noted that coating may

, where *V0*, *Vmax*, *Km*, and *[S]* are the initial velocity of

Furthermore, researchers have developed two strategies to wire enzymes to the electrode interface (**Figure 4B** and **C**). These include (1) mediated electron transfer (MET) and (2) direct electron transfer (this may refer to mediatorless electron transfer between the enzyme and the electrode). Such new tactics are not only useful for enzymatic biosensors but also for enzymatic BFCs which also involve

*DOI: http://dx.doi.org/10.5772/intechopen.85459*

H2O2 to mitigate interference effects.

bioelectrocatalysis.

for ethanol BFCs [51].

Menten kinetics, *<sup>V</sup>*<sup>0</sup> <sup>=</sup> *Vmax* \_\_\_\_\_\_ [*S*]

*Km* + [*S*]

lower the sensitivity and cause slow response time.

ing topic.

*Wearable Skin-Worn Enzyme-Based Electrochemical Devices: Biosensing, Energy Harvesting… DOI: http://dx.doi.org/10.5772/intechopen.85459*

such electroactive interferences. One approach is to incorporate electrocatalysts in wearable electrodes, such as PB or Pt [17, 40]. This offers low-potential detection of H2O2 to mitigate interference effects.

Furthermore, researchers have developed two strategies to wire enzymes to the electrode interface (**Figure 4B** and **C**). These include (1) mediated electron transfer (MET) and (2) direct electron transfer (this may refer to mediatorless electron transfer between the enzyme and the electrode). Such new tactics are not only useful for enzymatic biosensors but also for enzymatic BFCs which also involve bioelectrocatalysis.

First, the MET strategy utilizes a redox mediator, acting as an electronshuttle assistant between the enzymatic active center and the electrode. The substrate level, such as glucose, can then be monitored by the redox process of the mediator. This results in the independence of oxygen and mitigating the interfering signals due to the operation at low potentials. The first consideration in electrically wiring the enzyme with the electrode is the choice of the mediator that should be close to the redox potential of the active center of the enzyme to facilitate efficient electron communication between the enzyme and the conductive electrode surface. In particular, for enzymatic BFCs, the selection of mediators is crucial to positively control the cell voltage and enhance heterogeneous electron transfer to the order of a homogeneous transfer [50]. However, challenges of using mediators, particularly for BFCs, are their stability and deviated cell voltage. In addition, biocompatibility is highly vital for skin-worn applications. In spite of the assistance of electron shuttle by redox mediators, major concerns are their biocompatibility. One possible solution is employing nanomaterials or highly biocompatible catalysts. For example, mushroom/plant extracts could be used to obtain efficient "green" bioelectrocatalytic reactions for ethanol BFCs [51].

Second, direct electron transfer is an ideal goal of electrical wiring. It can be achieved by employing nanomaterials which suggest the direct electron transfer between enzyme active site and electrode. This wiring strategy is based on the shortening of the electronic contact of the enzyme and electrode (a short distance of ~1.5 nm) where the redox center of the enzyme can be regenerated directly by the electrode [52]. Therefore, this strategy can maximize the performance of bioelectronics. The engineering needs to consider the position of the active site inside the protecting protein and the conformation of the protein in order to wire the conducting materials with the redox center. This still remains the most challenging topic.

Several variables also affect the response nature of enzyme bioelectronics. Consideration of the fundamental theory of their functions will help to improve their performances. A key well-known model of enzyme behaviors is Michaelis-Menten kinetics, *<sup>V</sup>*<sup>0</sup> <sup>=</sup> *Vmax* \_\_\_\_\_\_ [*S*] *Km* + [*S*] , where *V0*, *Vmax*, *Km*, and *[S]* are the initial velocity of the reaction, the maximal rate of the reaction, the Michaelis-Menten constant, and the concentration of the substrate, respectively. In general, it is desirable to engineer the biointerface electrode system to obtain high *Vm* and low *Km* (good affinity). However, dynamic range is also a crucial characteristic for wearable biosensors. Traditionally, dilution or preconcentration can be used to adjust the level of the target to be fit in the linear range of the sensor; nonetheless, manipulating such processes for on-skin applications is sophisticated. Therefore, diffusion-limiting membranes may be a useful solution to tune the dynamic range. The linear range can be extended by coating a thin membrane over an active enzyme layer since the sensor response is controlled by the analyte diffusion and not by the nonlinear characteristic of enzyme kinetics. Nevertheless, it should be noted that coating may lower the sensitivity and cause slow response time.

*Wearable Devices - The Big Wave of Innovation*

analytes or biofuels for electroanalytical monitoring and energy harvesting, respectively. One of the most predominant enzymes used to develop wearable bioelectronics is GOx from *Aspergillus niger*. It represents an example of commercially available biocatalyst that has good stability, substrate specificity, and electron turnover rate [3, 4, 45]. It is a powerful biorecognition element for glucose biosensors, the most widely interesting devices for diabetes health management. As shown in **Figure 4** (A–C), the enzyme is immobilized on the electrode, establishing a biosensor. GOx contains two 80 kDa subunits. Each holds a tightly bound flavin adenine dinucleotide (FAD) cofactor, the important redox center which has a redox potential −0.32 V (vs Ag/AgCl) at pH 7. This redox center is crucial to transfer electrons and specifically oxidize β-D-glucose to gluconolactone. However, this FAD is shielded by the protein and a glycan structure, hindering electron exchange at the enzyme-electrode interface. Inevitably, this requires research efforts to address this roadblock [46, 47]. FAD plays an important role as a common cofactor for glucose oxidation biocatalysis. The redox process for FAD/FADH2, involving two electrons, is shown in **Figure 4D**, where the R group represents adenosine diphosphate and ribitol connected with the flavin. However, it is O2-dependent; accordingly, O2 fluctuations can vary the performance of this type of oxidase-based bioelectronics. Although alternative O2-independent electrodes utilized NAD-dependent electrodes can be used, they require a diffusional cofactor, not simple for wearable applications. Hence, FAD-dependent dehydrogenases are becoming interesting choices since they

are O2-independent and do not depend on diffusional mediators [48, 49].

The first generation of biosensors relies on quantifying O2 generation or H2O2 depletion (**Figure 4A**). This leads to key drawbacks, such as low dynamic range, dependency to oxygen fluctuations, and interfering effects. For instance, for glucose amperometric sensors, the detection of H2O2 at common first-generation electrodes needs the high applied detection potential where interfering compounds existing in sweat, e.g., ascorbic acid, uric acid, and some drugs, are also electroactive. Lowering the applied potential for the detection is a strategy to minimize

*Principles of interfacing the enzyme, such as glucose oxidase (GOx), with the electrode. Different generations of strategies (A–C: first, second, and third generations) are illustrated. (D) Reactions involving the glucose oxidase* 

**40**

**Figure 4.**

*biocatalyst.*

In addition, extra membranes can be a biocompatible barrier to address challenges from biofouling and interferents, especially when electrochemical operations are made in real matrices, samples, such as sweat. A perfluorinated sulfonated membrane (Nafion®) is an example membrane, which is also easy to drop-cast. This coating membrane can protect the enzymatic layer and also prevent anionic interferents, such as ascorbate [53].

Shelf life and operational stabilities of enzymatic electrodes are among the most critical challenges. The enzyme and active materials, such as mediators, can also leach during operations. Extensive studies have been made to improve enzyme bioelectrodes, such as by crosslinking hydrogels in the presence of the enzyme [54, 55]. Such crosslinking can entrap the enzyme to be more stable; moreover, this way enhances the loading of the enzyme, while the three-dimensional structure can facilitate the transport of analytes or biofuels, improving bioelectrode functions. Nevertheless, crosslinking enzyme or covalent binding of the enzyme can change the conformation of the enzyme and thus affect the activity [56]. Furthermore, one alternative to stabilize the enzyme electrode is the addition of stabilizers, such as polyelectrolytes, dextrans, glycerol, polyethyleneimine, and hydrophobic oils [57–59]. For instance, hydrophobic mineral oil or silicone grease can be used to minimize enzyme denaturation [58, 59]. The pasting liquid helps to lower protein mobility, maintain conformational rigidity of enzymes, and barrier to hydronium ions from acid environments. This strategy can stabilize many enzymes, such as GOx, LOx, AOx, horseradish peroxidase, amino acid oxidase, and polyphenol oxidase.

Increasing enzyme loading can also improve the performance of biocatalytic devices. Employing high surface nanomaterials is useful to enhance the surface loading of the target catalyst. A graphene-based electrode is a good example platform to offer a high enzyme loading (1.1 nmol cm<sup>−</sup><sup>2</sup> ); in addition, it offers a fast heterogeneous electron transfer rate (*k*s) of 2.8 s<sup>−</sup><sup>1</sup> [60]. Moreover, CNTs, which have high conductivity and specific surface, represent an outstanding candidate nanomaterial for electrochemical wiring [38, 61]. The thin nanoscale structure can intimately incorporate with the active enzyme. Adsorption of GOx on CNTs provides the apparent *k*s, of 1.5 s<sup>−</sup><sup>1</sup> [62]. The *k*s of GOx at the hybrid biocomposite can be as high as 11.2 s<sup>−</sup> 1 [63]. Therefore, mediatorless bioelectrodes with excellent electron transfer could be demonstrated. Their high three-dimensional architecture also offers an enhanced loading of enzyme and/or redox mediator immobilizations. As a result, this can enlarge the current output from the biosensor or BFCs. Importantly, for BFCs, the maximized OCV and current density could be observed [43]. This BFC consists of a GOx/catalase/CNT bioanode and laccase/CNT biocathode without additional mediators. The CNT/enzyme matrix was compressed together under high hydraulic pressure (10 kN). The resulting output in an airsaturated electrolyte (200 mM glucose in 0.2 M phosphate buffer solution, pH 7 at room temperature) after 3 days displayed a high maximum OCV of 1 V. Note that the GOx/catalase/CNT bioanode and the laccase/CNT biocathode showed OCV values of −0.35 and +0.6 V, respectively.

Importantly, biofluids from the skin (such as sweat and extracted interstitial fluids) contain a variety of chemicals that can inhibit enzyme activity, reflecting challenges in biosensing and BFC functions in real-time on-body applications. For instance, heavy metals can be found in sweat as the body expels chemicals or balances the charges. One example is Cu2+ which has been reported as an inhibitor to deactivate the enzyme. The Cu2+ in sweat can be in a range of 1.6–16 μM [11]. 0.1 μM Cu2+ could decrease the OCV value of the glucose BFC [64]. However, this enzyme-inhibitor electrochemical behavior is analytically attractive toward the development of self-powered biosensors, such as for direct heavy metal screening

**43**

**Figure 5.**

*Wearable Skin-Worn Enzyme-Based Electrochemical Devices: Biosensing, Energy Harvesting…*

or indirect cysteine monitoring. For example, cysteine prefers to bind with Cu2+ via the Cu-S bond; this superior conjugation between cysteine and Cu2+ removes metal

Since the O2 level in biofluids may vary, first-generation biosensors, employing O2-dependent mechanism, are subject to inaccuracy. This issue can be addressed by using fluorocarbon pasting liquids to supply internal O2 [65]. Using redox mediator as a second-generation sensor is another way to eliminate this error. Furthermore, FAD-dependent glucose dehydrogenase is an option to address O2-dependent problems due to its O2-insensitive nature, compared with GOx [49]. In addition, because of the high rate of homogeneous electron transfer rate between GOx and oxygen, GOx prefers to transfer electrons to oxygen rather than to the electrode, causing undesirable O2 competition effect [66]. Moreover, for BFCs and selfpowered sensors, the commonly used ORR cathode may cause the error under anaerobic conditions. The use of Ag2O/Ag redox cathode, which does not depend on ORR, can be used to operate BFCs, mitigating the possible O2 errors [30, 67]. Note that the reduction potential of Ag2O/Ag (0.342 V vs. SHE) is close to that of O2/ OH<sup>−</sup> (0.401 V vs. SHE) at pH 7. Moreover, using O2-rich cathode is another possible

*DOI: http://dx.doi.org/10.5772/intechopen.85459*

option to mitigate O2-deficit effects [68].

low-volume electroanalytical systems [70].

*iontophoresis and (B) reversed iontophoresis.*

**3.5 On-skin biofluid extraction: electrical-based approaches**

across the skin surface and subjects, ranging from 16 to 530 glands cm<sup>−</sup><sup>2</sup>

Normally, during exercise, sweat can be secreted around 20 nL gland<sup>−</sup><sup>1</sup>

example, the forehead or arm can generate sweat around 3 μL cm<sup>−</sup><sup>2</sup>

Each person has 2.03 million sweat glands; sweat gland densities vary broadly

fluctuation of sweat rate is also related to numerous factors, such as activity intensity and hydration level. Therefore, the limited volume of sweat causes a challenge in sweat analysis and operations. This leads to the development of miniaturized skin-worn electrochemical devices that can be practical in such small dead volume. For instance, the textile-based energy-harvesting BFC requires sweat volume per area of 40 μL cm<sup>−</sup><sup>2</sup> to deliver steady outputs [31]. Designing a capillary chamber is a possible route for

In addition to a passive way to collect sweat, one strategy is an active electrical-

based approach, called "iontophoresis" [71, 72]. This active strategy offers

*Electrical-based strategies using iontophoretic electrodes to extract biofluids, including (A) pilocarpine* 

[11, 13, 69].

or even lower. The

[11]. For

min<sup>−</sup><sup>1</sup>

ions from the bioanode, consequently turning on the OCV.

**3.4 Effects of oxygen fluctuations on electrochemical performances**

*Wearable Skin-Worn Enzyme-Based Electrochemical Devices: Biosensing, Energy Harvesting… DOI: http://dx.doi.org/10.5772/intechopen.85459*

or indirect cysteine monitoring. For example, cysteine prefers to bind with Cu2+ via the Cu-S bond; this superior conjugation between cysteine and Cu2+ removes metal ions from the bioanode, consequently turning on the OCV.
