Open Access is an initiative that aims to make scientific research freely available to all. To date our community has made over 100 million downloads. It’s based on principles of collaboration, unobstructed discovery, and, most importantly, scientific progression. As PhD students, we found it difficult to access the research we needed, so we decided to create a new Open Access publisher that levels the playing field for scientists across the world. How? By making research easy to access, and puts the academic needs of the researchers before the business interests of publishers.
We are a community of more than 103,000 authors and editors from 3,291 institutions spanning 160 countries, including Nobel Prize winners and some of the world’s most-cited researchers. Publishing on IntechOpen allows authors to earn citations and find new collaborators, meaning more people see your work not only from your own field of study, but from other related fields too.
To purchase hard copies of this book, please contact the representative in India:
CBS Publishers & Distributors Pvt. Ltd.
www.cbspd.com
|
customercare@cbspd.com
Wearable devices seem to have great potential that could result in a revolutionary non-clinical approach to health monitoring and diagnosing disease. With continued innovation and intensive attention to the materials and fabrication technologies, development of these healthcare devices is progressively encouraged. This chapter gives a concise review of some of the main concepts and approaches related to recent advances and developments in the scope of wearable devices from the perspective of emerging materials. A complementary section of the review linking these advanced materials with wearable device technologies is particularly specified. Some of the strong and weak points in development of each wearable material/device are clearly highlighted and criticized.
Environmental Monitoring and Sensing Technology Laboratory, School of Information Science and Engineering, Ningbo University, Ningbo, P.R. China
Qinghui Jin
Environmental Monitoring and Sensing Technology Laboratory, School of Information Science and Engineering, Ningbo University, Ningbo, P.R. China
Jiawen Jian
Environmental Monitoring and Sensing Technology Laboratory, School of Information Science and Engineering, Ningbo University, Ningbo, P.R. China
*Address all correspondence to: jinhan@nbu.edu.cn
1. Introduction
Wearable devices are apparatuses that can be worn directly on the skin in different parts of the body. These devices have gained considerable attention owing to their ease of collection of crucial information in real-time regarding a wearer’s health, both continuously and non-invasively [1, 2, 3, 4, 5, 6, 7]. In contrast to the traditional non-invasive methodologies (e.g., X-ray, type-B ultrasonic, nuclear magnetic resonance (NMR)), non-invasive diagnosis implemented by wearable devices creates new opportunities for remote and continuous healthcare monitoring in non-clinical settings [5, 8], with the ability to detect developing diseases at intervals between routine examinations. The use of wearable healthcare devices also encourages people to take greater interest in their own healthcare in a more convenient and cheaper way, thereby improving their compliance.
Wearable devices are becoming smaller and more mobile with time, opening new alternatives to traditional ways that providers have interacted with patients, carried out tests, collected data and delivered treatments [9, 10, 11]. Wearables come in many forms; there are smart wristbands [12], watches [13], shirts [14], shoes [15], headbands [16], eyeglasses [17] and necklaces [18], among others. Most of them contain sensors that gather raw data that is fed into a database or software application for analysis. Analysis typically triggers a response that would, for example, alert a physician to contact a patient who is experiencing abnormal symptoms, or might send an appreciative message when an individual achieves a fitness or diet goal [11].
Recent insights into wearable devices have resulted in rapid development of various kinds monitoring different parameters, such as pressure/strain, body vital signs (e.g. heartbeat rate, respiration rate and temperature) and biomarkers that can be found either in body fluids (e.g. saliva, sweat and tears) or in skin odor and breath [19, 20, 21, 22, 23]. Among these reported wearable healthcare settings, skin-based wearable devices have considerable advantage in allowing simultaneous monitoring of multi-physiological indexes and biomarkers [19, 21, 22, 24]. By continuous or frequent detection of the level of physical markers (e.g. pressure pulses and body temperature) using different sensing techniques, wearable devices can provide one of the most comprehensive feedbacks on human health states [7, 8, 19, 22, 24, 25, 26, 27, 28, 29, 30, 31, 32]. Perspiration contains a diversity of chemical species (e.g. minerals, trace elements, lactic acid, urea and volatile organic compounds) that can be detected and monitored in sweat, which can assist in determining the health status of a person. A good example of the potential of sweat diagnostics is the detection of cystic fibrosis [33, 34], currently in clinical use [35]. However, small sample volumes, dilution of analytes and variability in the measurement of concentration are some of the challenges that are encountered using sweat for diagnosis. Table 1 summarizes the different physiological indexes and diagnostic biomarkers using wearable devices. As inferred in Table 1 for gathering broad information of the health status, the skin-based wearable devices should integrate multiple sensing features. Consequently, the performance of a wearable device depends heavily on the proper sensing characteristics of its integrated sensing parts.
Although many challenging issues are of concern, developing technologies proposed in the last few years significantly enhance the practicability of applying these wearable devices in clinic use and academic research. One of the main reasons for the advance of wearable devices is covertly the rapid emergence of new materials with interesting properties [19, 23, 24]. As a result, modern wearable devices are becoming cheaper, smaller and more reliable. Several reviews on these aspects have already been published [38, 39, 40]. The aim here is to present the current state of non-invasive skin-based wearable devices from the perspective of materials science and engineering. In view of these emerging skin-based wearable devices that are expected to meet diverse healthcare applications, smart materials used for these purpose should be compliant with the following basic criteria:
Highly sensitivity to subtle changes in the physiological index: subtle change in the physiological index, for example, body temperature, may be an early indicator of health problems. For the purpose of catching very slight changes in the human body, stimuli-responsive materials used in skin-based wearables should be very sensitive. For instance, thermal-responsive materials used in body temperature sensors should sense down to 0.1°C variation, obviating missing any sign of health risk.
Adequate detection limit: stimuli and/or analytes obtained from human skin are generally at low level regime, for example, pressure of pulsing artery is typically within several hundreds of Pa, which gives the benchmark for designing wearable pulsing sensors. Consequently, an adequate detection limit is particularly required by the responsive materials for these stimuli and/or analytes. Since the level of analytes coming from human skin is usually lower than that of analytes obtained directly from tissue or body fluid (excluding sweat), detection limit of these smart materials dedicated in the design of skin-based wearables is even lower compared with those wearable counterparts developed for other uses (e.g. implantable devices, breath sample analyzer, etc.).
Compatible with human skin: materials that are featured as flexible, stretchable and less irritating are especially to be considered in these skin-based wearables, largely due to their unique biocompatibility with soft and moist skin. Flexible and stretchable elastomers are more suitable for fabricating strain-responsive devices.
Owing to the continuous increase in the essential requirements of and interest in, the skin-based wearables, an update on the relevant materials is presented in this review, with an emphasis on novel materials and sensing techniques, with regard to their advantages, limitations, challenges and future trends. A comprehensive description of the unique technique in accessing health conditions by monitoring skin biomarkers and physiological parameters has also been also particularly in focus.
2. Smart materials for wearable healthcare devices
2.1. Wearable temperature sensors
To meet the requirements of non-invasive and quantitative monitoring of body temperature, wearable temperature devices are required to be flexible, biocompatible, lightweight and highly sensitive within temperatures range 35–42°C [41, 42]. Several temperature-sensing devices have been invented in the past, but flexible wearable temperature-sensing devices (Table 2) are mainly limited to thermally resistance (resistometric) and thermally sensitive field-effect transistor (FET) transduction mechanisms. While the operation mode previously relied on resistance changes of the sensing elements according to the temperature change, the latter relies on detecting FET transfer characteristics as a result of thermal effects [7, 41, 42, 43, 44, 45, 46, 47, 48, 50, 53].
The attractive advantages of low-cost, simple configuration and ease of mass production made thermal-resistance-based temperature devices the most popular as skin-mounted sensors [43, 44, 45, 46, 48]. These devices relay thermal expansion and/or variation in temperature coefficient of resistance (TCR) induced by changing temperature [46]. An acceptable thermal expansion coefficient and TCR of metal and metal oxides have become available and have been used in the past decades [45, 48, 54]. A representative metal-based sensing element for tracking skin temperature is gold foil or nanoparticles that possess a desirable TCR [45, 54]. For TCR-based devices, gold has good ductility and can be deposited on a range of substrates, including glass, SiO2 and flexible polymer [45, 54]. This allows the design of ultrathin [55, 56] and flexible wearable temperature devices suitable for a range of complex measurement under different conditions.
An example of a successful wearable temperature sensor is the epidermal device designed for continuous thermal monitoring of human skin (Figure 1a). With integrating 4 × 4 TCR (Cr/Au) sensor array on a thin elastomeric substrate, this ultrathin conformal device gives reliable performance and minor deviations [45, 61]. However, the sensing range for most of metal-based sensing elements is limited to relatively high temperatures (normally >42°C) because of their low thermal-response [46]. Inadequate sensitivity of metal-based sensing elements also brings the need for sophisticated measurement electronics for increasing signal-to-noise ratio. Hence, to increase the thermal response range and improve the sensitivity of these devices, and increase efficiently the signal-to-noise ratio, alternatives have been devised, one being based on carbon nanotubes (CNTs) where resistance is strictly dependent on temperature. By dispersing proportional amounts of CNTs into specific polymer such as poly(styrene-b- (ethylene-co-butylene)-b-styrene; SEBS), flexible sensing composites had very high sensitivity to changes in ambient temperature from 20 to 40°C [43, 44].
Figure 1.
Schema of temperature sensors. (a) Image and the sensing performance of a 4 × 4 TCR sensor array after application to the skin using a water-soluble adhesive tape based on polyvinyl alcohol. (b) Cross-section SEM image of GNWs on PDMS and its “current and resistance versus temperature” curves. (c) Schematic of the TS-gated device and its response to temperature. (d) Photograph of a film of copolymer with graphite filler, with the temperature dependence of the resistivity of the temperature sensor. Reconstructed from Refs.[57, 58, 59, 60].
Compared to thermal resistance-based devices, FET-based temperature devices have greater sensitivity, and are easily incorporated into integrated circuits of convenience in signal amplification [7, 42, 50]. By applying thermal-responsive materials as an active channel, electrode or gate dielectric layer into the FET structure, promising results in determining skin temperature have been achieved. Few examples of FET-based temperature sensors include graphene-based sensing elements, metal-polymer hybrids, inorganic-polymer hybrids [7, 42, 47, 50, 51, 53]. Owing to the appropriate and adjustable thermal activation energy of the charger carrier for these thermal-responsive materials, the detection limit of the FET-based temperature devices can reach milli-Kelvin [47]. The inherent characteristics of FET-based devices also help in the precise sensing signals, with a deviation of <0.1°C [47].
Inadequate thermal-response, unsatisfactory stability and significant interference from the ambient environment are all challenging issues for wearable temperature devices [43, 46, 50]. Latterly, graphene nanowalls (GNWs) in combination with polydimethylsiloxane (PDMS) have been used as the sensing chip in a thermal-resistive sensor (Figure 1b) [46]. In comparison to conventional temperature sensors, this device has a sensitivity over 108.6 Ω °C−1 between 35 and 45°C, which is sufficient to monitor human body temperature. Further evidence proves that the high response of the sensor can be attributed to the excellent TCR (0.214°C−1) of GNWs and large expansion coefficient of PDMS [46]. Stability of these temperature devices has been many times reported, who found that the transition of morphology and phase of sensing elements during cooling-heating cycle is the main reason for irreproducibility of a thermal response [44, 48]. Adsorption of water vapor and/or ambient gas species is another crucial reason that generates hysteresis and causes significant interference [50]. Consequently, thermally stable materials are screened and extra encapsulation layers are added to decrease hysteresis, thereby improving stability and enhancing reproducibility (Figure 1c) [44, 48, 50]. Taking these materials from the lab to the technological/industrial benchmark would require advanced printing techniques. A graphite-polymer-based ultrathin thermometer of 15 μm thickness that could be accomplished by such simple print and press method is a good example for fabrication of large-area stretchable and transparent wearable temperature devices in the future. These fabricated thermometers should be able to measure temperature precisely between 25 and 50°C (Figure 1d) [47].
2.2. Wearable strain sensors
Wearable strain sensors could be useful for the detection and monitoring of movement-based signals, such as heartbeat and respiration rate. Generally, wearable strain devices should be lightweight, reliable, flexible and stretchable to match the mechanical properties of human skin. Sensitivity of these strain sensors must also be in line with their diverse healthcare applications. For instance, mean value of feet pressure for adults mostly ranges from 140 to 868 kPa [62], whereas systolic pressure (artery pulse) for healthy persons is within 0.67–5.32 kPa (5–40 mmHg) [63]. This suggests that high sensitivity in a low-pressure regime (perhaps <10 kPa [64]) is required for strain sensor trying to detect arterial pulse. To date, many working models for strain devices have been developed [65]. Among them are flexible and stretchable devices based on piezoresistive, piezocapacitive and piezoelectric principles seem to be closer to practical implementation, due to their relatively simple sensory configuration, uncomplicated read-out systems and acceptable dynamic performance [24, 66, 67, 68].
2.2.1. Piezoresistive-based strain sensors
Piezoresistive-based strain devices are typically composed of electrically conductive sensing films coupled with flexible substrates. When the structure of the device is deformed, changes in the microstructure of sensing films lead to changes in electrical resistance as a function of applied strain. Metal-based foil and graphitic-based sensing networks (e.g. InGaZn and graphene) are widely used in pressure/strain sensing [73, 74, 75]. In conductive networks, electrons can pass through overlapped nanomaterials within the percolation network. Stretching or bending in the conductive network results in disconnection between overlapped areas and thereby loss electrical connection; consequently, there is an increase of the electrical resistance of sensing film. Another known strain-responsive mechanism is crack propagation, that is, cracks originate and propagate in brittle thin films coated on the top of soft polymer layers upon stretching [20]. Generally speaking, the sensitivity (gauge factor) of the conventional metal foil-based strain device is in the range of 2–5 (a.u.) and semiconductor-based strain sensors might be 100 or greater. The value of sensitivity for flexible strain devices can be in the range of 1–100 s, depending on the sensing mechanisms, materials and micro/nanostructures [20].
To meet the demand of practical healthcare applications, a proof-of-concept wearable pressure device based on ultrathin Au nanowires (NWs) has been developed (Figure 2a) [76]. The Au NWs directly coated on a tissue paper can detect pressure forces as low as 13 Pa with high sensitivity (41.1 kPa−1). Furthermore, the devices also provide fast response times (<17 ms) and high stability (450,000 loading–unloading cycles) [76]. A similar strategy has also been used in the design of piezoresistive strain devices comprised of other sensing elements, for example, graphene [74]. Furthermore, these prototypes give excellent performance in real-time monitoring of human blood pulses and detection of music motion to confirm their usefulness in healthcare applications [74, 76].
Figure 2.
Strain-responsive devices based on emerging materials and fabricating technology. (a) AuNWs-coated tissue paper and (b) piezocapacitive-based strain device composed with AgNWs. (c) Three-dimensional (3D) microporous dielectric elastomer for a piezocapacitive strain device. (d) Piezoelectric devices based on aligned arrays of nanofibers of PVDF. Reconstructed from Refs. [69, 70, 71, 72].
Strain sensors based on the piezoresistive conductive polymer (CP) and their composites have also been receiving considerable attention due to their excellent flexibility and biocompatibility [77]. These CP composites are usually based on polymeric matrix filled with stainless steel fibers (SSFs) and carbon nanotubes (CNTs) [57]. CP and their composites generally have robust piezoresitive behavior for a long period time under the temperature ranging from −50 to +200°C [57]. Note that their piezoresistivity is closely dependent on the conductive phase content. When this is low, piezoresistivity monotonically increases; in contrast, the piezoresistivity is monotonically decreasing at high conductive phase content [58]. Such flexible piezoresistive CP or CP composites are attractive candidates for wearable devices, for instance, through coating poly(3,4-ethyl-enedioxythiophene–poly(styrenesulfonate) (PEDOT:PSS)/polyurethane composite polymer on the spring-like compressible micro-pyramid-structured substrate, a stretchable strain sensor with high sensitivity has been developed [59]. The sensor had a sensitivity of 10.3 kPa−1 when stretched by 40%, which is sufficiently sensitive to detect human pulse pressure (23 Pa) [59]. Furthermore, its suitability in non-invasive healthcare applications is validated by showing excellent piezoresistivity and a fast response rate in measuring the primary features of the waveform in the human pulse.
It should be noted that beyond those conventional strain-responsive materials, stretchable silicon has received considerable attention from researchers because of its superior response to strain [60]. This silicon can be reversibly stretched and compressed to a large extent without damage, a new finding that might provide opportunities to design tunable structures that are compatible with mechanical strain [60]. Miniaturization and simplicity of configuration of the devices, as well as decreasing power consumption, is another challenging issue. Hence, a high-resolution unpixelated smart device was proposed for sensing stimuli. The sensing film of the device is constructed from two parallel GNP strips with antiparallel sensitivity gradients [78]. This design enables the device not only to detect the loaded strain, but also to discriminate the spatial position of the strain [78]. The gradient design of the sensing film can remarkably decrease the complexity and power consumption of a future strain-responsive device.
2.2.2. Piezocapacitive-based strain devices
Piezocapacitive-based strain devices use a highly compliant dielectric layer sandwiched between a pair of stretchable electrodes. Applied strain changes the distance between two electrodes, which results in a change of capacitance [24]. For instance, a piezocapacitive strain sensor composed of silver nanowire (AgNWs), Ecoflex, liquid metal and PDMS was built for finger flexing and knee motions among other body movement tracking (Figure 2b) [79]. Quick response (<40 ms) and good pressure mapping function was discovered during the examine process; and the designed piezocapacitive strain device has satisfactory wearability when mounted on the human body.
Besides these conventional strain devices, a transparent flexible pieocapactive pressure/strain sensor has been developed by Lipomi et al. [80]. The pieocapactive sensor, made by spraying transparent carbon nanotube on PDMS, gave a desirable performance in detecting pressure down to ~50 kPa [80]. Note that this transparent piezocapacitive skin-like strain sensor is extremely compliant mechanically, physically robust and easily fabricated, although it is less sensitive than its other counterparts.
In contrast to those piezoresistive counterparts, piezocapacitive strain devices have high linearity, stretchability and low hysteresis, but suffer from low sensitivity. Thus, a 3D microporous dielectric elastomer with giant piezocapacitive effect was used to construct a pressure device (Figure 2c) [81]. Due to the presence of micropores within the elastomeric dielectric layer, the resulting piezocapacitive pressure device is highly deformable by the minimal amount of pressure, leading to a marked increase in sensitivity. The gradual closure of micropores under compression also increased the effective dielectric constant, thereby enhancing sensitivity. The 3D microporous dielectric layer with serially stacked springs of elastomer bridges can cover a much wider range of pressure than those of previously reported micro−/nanostructured sensing materials. This proof-of-concept can help in monitoring both ultralow and high levels of human activity. Noteworthily, materials with unique microstructure will enhance their piezocapacitive performance. Compared with those piezocapacitive materials with flat surface, the pressure sensor comprised of pyramidal dielectric elastomer significantly enhanced blood pulse monitoring and force tracking [82]. Similar improvements have also been investigated for piezoresistive elastomer with hollow-sphere [83], thickness-gradient [84] and pyramidal microstructures [85], implying an efficient approach to improve the performance of strain sensors by well-designed strain-responsive materials.
2.2.3. Piezoelectric-based strain sensors
Piezoelectric-based strain sensors detect changes in pressure or strain; in particular, sensing signal (voltage) is only generated by dynamic changes of applied strain or force. Piezoelectric materials, especially certain crystals (e.g. ZnO and LiNbO3) and some ceramics (e.g. BaTiO3 and Pb(Zr11xTix)O3 (PZT)), generate a voltage potential when the crystal lattice is deformed [86, 87]. The sensitivity of piezoelectric materials depends on their crystal structure and the voltage generated is directly proportional to the applied strain. These sensors have very good high-frequency responses, which makes them ideal for measuring dynamic forces [88]. To match the mechanical properties of human skin, recent piezoelectric materials used in wearable devices are mechanically flexible [87, 89, 90]. These flexible piezoelectric materials can be based on either rigid piezoelectric materials embedded in a flexible substrate or a piezoelectric polymer film.
Compared with piezoresistive- and piezocapacitive-based strain sensors, strain devices comprised of piezoelectric materials have greater dynamic durability. The nature of good high-frequency response to an applied strain makes this device more suitable for measuring vibrations. For example, piezoelectric-based strain sensors could be installed in a smart watch for monitoring heartbeat rate. However, the dynamic piezoelectric sensing mechanism disallows them from measuring static forces. A classic piezoelectric material commonly used in flexible strain devices is poly(vinylidene fluoride; PVDF), a piezoelectric polymer that has considerable flexibility [90]. Its flexibility and high piezoelectric factor make PVDF an attractive material for flexible strain devices. For example, Persano et al. (Figure 2d) reported on aligned arrays of nanofibers of PVDF [91]. Free-standing 3D architectures of aligned arrangements of PVDF fibers have been designed (Figure 2d); furthermore, the polymer chains inside the fibers adopt strongly preferential orientations [91]. This microstructure enables ultra-high sensitivity in measuring pressure, even at exceptionally low values (0.1 Pa); in particular, exceptional piezoelectric characteristics offered by these devices indicate an application in monitoring human motion.
2.3. Wearable devices for recording electrical conductivity through the skin
Wearable devices that detect tiny electrical changes on the skin could be helpful for monitoring the electrical activity of heart to locate an arrhythmia (abnormal heartbeat). These results can help one or one’s doctor to decide whether you need medicine, a pacemaker, an implantable cardioverter defibrillator (ICD), cardiac neural ablation or some other surgery. For wearable and accurate monitoring of electrophysiological signals in a variety of everyday conditions, reliable biopotential-capturing electrodes that are reusable with long-term stability required, and they should also not cause skin irritations or allergic reactions.
To date, disposable silver/silver chloride (Ag/AgCl)-gelled electrodes are most commonly used, but these electrodes have limited storage time (<1 year) and are not reusable [96]. The best possible alternatives are flexible dry electrodes, which are usually made of polymeric- and textile-based materials [69, 96, 97, 98]. For instance, after patterning a metal layer on a thin PDMS substrate, a biocompatible and flexible PDMS-based (polymetric) electrocardiogram (ECG) recording device can be obtained (Figure 3a) [97]. In the case of textile-based electrodes, these are normally fabricated by knitting, weaving, embroidering and non-weaving methods. Figure 3b shows photographic images and comparison of the response of the textile-based dry electrode and (Ag/AgCl)-gelled electrode in wearable ECG recording devices. These flexible dry electrodes in general are more skin compatible compared with those of gelled electrodes (Figure 3) [69, 96, 97, 98]. However, disturbances to ECG signals caused by body movement can often appear. Despite larger disturbances due to movement, these flexible dry electrodes cause relatively little irritation or itches on human skin compared with (Ag/AgCl)-gelled electrodes. After wearing the polymeric dry electrodes for 7 days, no significant changes on skin were seen (Figure 3a); however, the skin under the (Ag/AgCl)-gelled electrode turned red and became itchy in two subjects. In conclusion, these newly proposed dry electrodes have comparatively good fidelity and are comfortable regarding body contact as well as being acceptable in long-term stability, indicating broad applicability to the ubiquitous field of biosignal monitoring.
Figure 3.
Typical examples of flexible dry electrodes for ECG signal recording. Photographs of (a) polymeric- and (b) textile-based electrodes, and their response to biosignals compared with those of commercial Ag/AgCl electrodes. Reconstructed from Refs. [92, 93, 94, 95].
2.4. Wearable sensors for analyzing sweat metabolites
Levels of sweat metabolites (such as lactate and uric acid) and electrolytes (metal ions, such as sodium and potassium), as well as skin humidity, are useful physiological parameter indirectly reflecting health status of an individual, and can potentially be used to probe body conditions by non-invasive monitoring [21, 23, 33, 34, 37, 99]. Table 1 summarizes the use of sweat body fluids as physiological biomarkers [7, 8, 27, 33]. The concentration of uric acid and urea in sweat is much higher in uremic patients [100]. Levels of skin humidity and pH of sweat, as well as its electrolytes, are important indicators of dehydration and electrolyte loss during exercise [101]. Accordingly, sweat-based wearable devices are required for monitoring these non-invasive health states.
To date, sweat-based healthcare devices are mainly based on an electrochemical sensing principle, owing to the low-cost, high-performance and excellent portability of electrochemical devices [21, 23]. Figure 4a shows the configuration of a generally reported electrochemical sweat-based device [102]. Three electrodes are printed on the flexible substrate and, as might be necessary, a fourth electrode can be present, that is, a working electrode (WE), a counter electrode (CE, sometimes as anode), a reference electrode (RE), and a cathode. WE and CE are normally composed of conductive materials, such as carbon black, CNTs, Pt black, modified Ag or Au, etc. [21, 102, 103]. Ag/AgCl is usually used as the RE. Conformal contact between the electrode surface and the skin is necessary to ensure efficient functioning of the devices. Accordingly, substrates for these devices should be made of fabric (e.g. wool, cotton or nylon) or flexible plastics [21]. Unique technology is required for fabricating firm electrodes/substrate configuration.
Figure 4.
General fabrication technology of sweat-based wearable devices. (a) Primary configuration of the sweat-based electrochemical wearable devices. (b) Steps involved in screen printing thick-film electrochemical devices. Left of Figure 4b: Representative design of flexible device fabricated by a screen-printing technique. (c) Preparation progress of stamp transfer electrodes. (d) Typical example of a tattoo-based wearable device. Reconstructed from Refs. [21, 23].
Figure 4b–d shows three typical approaches [21, 23, 37, 70, 103, 104, 105] for attachment of electrodes to a flexible substrate or skin [21, 23]. The first approach relies on screen-printed electrodes [23]. This approach has been massively employed in production due to its low-cost, industrial-scale fabrication of robust electrodes on different substrates. The geometries and thickness of the electrodes can be readily adjusted by changes of the screen mask to meet the design requirements of the wearable device. However, preparation of wearable devices by screen printing involves several complicating issues; for instance, the influence of manufacturing conditions and substrate properties, as well as the composition of the ink, need to be taken into consideration, since these factors significantly determine the performance of the devices. Generally, the natural properties of substrates must be compatible with the printing process and the specific operational environment. Some necessary modification, for example, doping of the noble catalyst, may be required for the inks to obtain specific functionality. The annealing temperature also requires optimization to avoid undesired deformation of the electrode and substrate. The left part of Figure 4b depicts the screen-printed microelectrodes on polyethylene naphthalate (PEN) for dopamine detection, suggesting the convenience of depositing responsive materials by this approach.
The second approach used for attachment of electrodes to a flexible substrate or skin relies on stamp transferred electrodes [23]. The pattern-transfer technique is very useful in fabricating sensing materials on a non-planar substrate, particularly on uniformly non-planar substrates. Figure 4c shows details of the stamp transfer process. It is known that, through pattern-transfer technique, responsive materials are highly compatible with irregular substrates possessing diverse surface morphologies (e.g. skin), without compromising the structural integrity of the pattern. Stamp transferred electrodes notably give better performance than those fabricated by screen-printing technique [23]. This fabricating technique provides a novel method to form the ink-based printable materials on non-planar and oversized surfaces incompatible with standard screen-printing methods.
The third approach used for attachment of electrodes to a flexible substrate or skin relies on epidermal suspended electrodes [21]. The rationale behind this approach is to deal with the major problem for traditional wearable device by restricted intimate contact with the skin at limited regions – something that can be solved if the sensing electrodes can be directly attached on the skin. Inspired by tattooing, the strategy of using elastomeric stamps or the tattoo technique to from electrodes directly on human epidermis has been proposed. The strategy includes steps of screen printing of conductive and insulating inks on commercial, temporary tattoo-base paper, to form the electrodes and/or devices, and then flip and apply the fabricated tattoo or printed electrodes to the skin. To adapt for mechanical stress and overcome the potential deformations due to bodily movements, it was suggested that finely dispersed carbon fibers should be incorporated into the tattoo-based devices. Numerous results showed the success and convenience of applying this new technique in preparing electrodes and/or wearable devices on the skin. Among these examples, a classic prototype reported by Jia et al. [71] was noteworthy. In the prototype (Figure 4d), an enzymatic tattoo amperometric biosensor was built for continuously monitoring lactate levels in the perspiration as a biomarker of physical stress. This flexible, printed, temporary-transfer tattoo, electrochemical biosensor adhered well to the skin and had acceptable selectivity to lactate, with linearity up to 20 mmol/L. In particular, the device had commendable resilience against continuous the mechanical deformation expected from epidermal wear.
Simultaneous screening of target biomarkers is helpful in ensuring the accuracy of measurements. Consequently, multifunctional wearable devices combining sensing and therapy have been particularly a focus of attention. Through providing sufficient information of heater, temperature, humidity, glucose and pH, as well as several chemicals in sweat, these multifunctional wearable devices (e.g. graphene-based electrochemical device) give a comprehensive profile of health status [27, 106]. On account of these attractive results, it is expected that, through combination of advanced materials science and these innovative preparation techniques, sweat-based wearable devices could possibly alert traditional disease prevention.
Spontaneously measuring chemical and ECG parameters will provide more comprehensive information about health status; however, most of the reported wearable healthcare sensors can only measure chemical or ECG parameters. To obtain fuller information of the human body, these monofunctional sensors need to be integrated into a sensor matrix, which increases the complexity of fabrication. An alternative strategy to resolve this issue is to design sensors that can be used for monitoring chemical and ECG parameters spontaneously. Based on this strategy, a wearable chemical-ECG hybrid biosensing system is proposed for real-time health and fitness monitoring [107]. The proposed system should give a desirable performance in simultaneously sensing lactate (0–28 mmol/L) and monitoring ECG parameters [107]. Such proof-of-concept opens a new way in designing flexible smart sweat-based healthcare devices.
2.5. Wearable sensors for detection of volatile biomarkers
Numerous experimental data have verified the robust correlation between volatile organic compounds (VOCs; organic compounds that have a high vapor pressure under ordinary room temperature conditions) and health status [72, 92, 93, 94, 95, 108, 109, 110, 111, 112, 113, 114, 115, 116, 117, 118, 119, 120, 121, 122, 123, 124, 125, 126, 127, 128, 129, 130, 131, 132, 133, 134, 135, 136, 137, 138, 139, 140, 141, 142, 143], indicating that changes in VOC profiles (amount and/or species) of human skin can be used for early diagnosis and warning of potential health risks [109, 110, 112]. The ideal approach to track VOCs exuded from human skin is expected to be non-invasive, real-time and operatable in non-clinical settings.These wearable devices can be managed either autonomically or remotely by a general practitioner.The concentration of these VOCs is normally at tens of ppb, suggesting high sensitivity and low LODs is required for these skin VOCs tracking sensors.
2.5.1. Metal nanoparticle-based resistometric sensors
In view of confirmation of the usability of metal nanoparticles [110, 113, 144, 145] in trace VOC detection, and other physical and environmental parameter monitoring (e.g. pressure, relative humidity and temperature) [146], Segev et al. [78] developed a smart patch based on gold nanoparticle (GNP) film which gave a satisfactory performance. Based on this strategy, they developed series of analogous GNP-based devices for tracking VOC biomarkers extracted from human skin [54, 111]. Figure 5a shows 2 typical examples of a GNP-based device and its performance relative to several kinds of VOCs found in human skin/breath. Interestingly, an acceptable discrimination capability was observed for the device after combination with specific algorithms [54, 111]. To prove the feasibility of the GNP-based devices in practical healthcare application, Broza et al. [93] used the approach developed to monitor VOCs of skin samples (Figure 5b); they included a heterogeneous group of 30 volunteers with no acute or active disease. Although their main goal was to associate VOCs from skin and breath of specific volunteers, the second goal was to achieve a chemical profiling based on the VOC samples from either skin or breath. This method successfully created a unique chemical barcoding enabling classification and clustering of the volunteers, for example, clustering based on age-related VOCs.
Figure 5.
Typical examples of the practical and potential sensing elements for skin VOCs detection. (a) Functional group (at different chain-length) modified GNPs-based flexible device. (b) Combined volatolomics monitoring by nanomaterial-based devices. (c) Fiber-like SnO2 for tracking trace amounts of VOCs. (d) A colorimetric VOCs device composed of different odor-responsive pigments. Reconstructed from Refs. 71, 78, 128 and [153, 163].
2.5.2. Potential candidates for tracking skin VOCs
Several materials appear to be sensitive to VOCs from human breath and odor, as well as air pollutants. Although there is no direct evidence of the use of these materials to track VOCs from the skin, they have no great potential for the near future. Thus, a general description of these candidates is introduced in the following sub-sections.
2.5.2.1. Metal oxides
Gas-sensitive metal oxides can be divided into transition- and non-transition metal oxides [147]. It is generally believed that the former are more gas-sensitive than the non-transition metal oxides because of their more active chemical structure. Primary factors that influence sensing behavior of metal oxides include chemical components, surface area (surface-to-volume ratio), microstructures of sensing layers, humidity and temperature [148].
Generally, very few unitary metal oxides possess favorable properties for sensing VOCs, especially in tracing gas species at ppb level. For this level of sensitivity, recent studies have focused mainly on composites, that is, binary, ternary, quaternary or even more complex metal oxides (such as noble metal decorated metal oxides) [147, 148]. Details of the enhancement mechanism have been published [147, 148, 149]. Besides the chemical composition, materials with high surface-to-volume ratio (high surface area) provide more reaction sites. Fabrication of nanofibers by electrospinning (Figure 5c), generation of thin film by sputtering, as well as creating porous structures with help of pore forming agents, are the most popular ways of obtaining sensing candidates with a high surface area [150, 151, 152, 153]. It has now been shown that, when the grain size is less than twice the thickness of surface charge layers, the grain is fully involved in the space-charge layer; consequently, gas sensitivity of the oxides can be significantly increased by adjusting grain size [149, 151]. Synthesis of oxides with unique shape and morphology is a very desirable method of enhancing the sensitivity of metal oxide gas sensors, as the shape and morphology of the oxides determines the crystallographic facets exposed on the surface of a nanocrystal, and will therefore determine the number of atoms located at the edges or corners [151]. By controlling the shape and morphology using different synthetic routes, the reactivity and selectivity of a nanocatalyst can be manually tailored.
One of the main challenges in metal oxide sensors is that water adsorption appreciably lowers the sensitivity of metal oxide gas sensors, since water adsorption on the metal oxide surface prevents donation of electrons to the sensing layers. Furthermore, prolonged exposure to humid environments leads to gradual formation of stable chemisorbed OH− on the surface, resulting in progressive degradation. However, surface hydroxyls start to desorb at ~400°C so that the hydroxyl ions can be removed by heating to temperatures >400°C [151, 159]. Thus, an optimal operating temperature will essentially eliminate humidity interference. As the balance of adsorption/desorption is markedly affected by temperature, optimization of operating temperature improves the magnitude of response/recovery speed and response. Table 3 lists several typical examples covering the above-mentioned strategies using SnO2 as the sensing element [147, 150, 152, 154, 155, 156, 157, 158]. It is clear that shape/morphology, surface area and operating temperature influence the performance.
Comparison of ethanol sensing performance, based on different SnO2-based sensing materials.
2.5.2.2. Conducting polymers
Conducting polymers (CPs) and their derivatives have been used as the active layers of gas sensors over the last century [160]. Compared to metal oxides sensors, devices made of conducting polymers have high sensitivities and short response durations at room temperature [161]. Conducting polymers are easily synthesized and their molecular chain structure can be conveniently modified by copolymerization or structural derivations [148, 161]. Furthermore, conducting polymers have good mechanical properties, which allow easy fabrication of sensor devices, and therefore considerable attention has been paid to these sensors.
Some of the earliest studies indicated that the sensing mechanism of CPs to VOCs involves swelling, namely the adsorbed VOC leads to the separation of polymer chains and increases the hopping distance of charge career for charge transport. As a result, the electrical resistance (response signal) increases [148, 160, 161]. In some cases, polar VOC molecules behave as electron acceptors or donors contributing a charge carrier to the polymer, and subsequently participate as dopants that increase or decrease carrier concentration [148, 160]. Hybridization of CPs with inorganic and graphitic nanoscale building blocks, such as metal oxides, CNTs or graphene, are more advantageous in polymer processing [148]. These hybridized CPs can also provide tunable sensing performance to VOCs through hybridization of CPs with optimal nanoscale building blocks, furthering the design of advanced sensing devices for specific monitoring of VOC biomarkers.
2.5.2.3. Responsive dyes
Theoretically, if a VOC molecule interacts with a responsive dye, a visual response can be obtained by a change in color. Based on this working principle, colorimetric VOCs devices have been designed based on two fundamental requirements: (1) each chemically responsive dye must contain a center that strongly interacts with the analyte; (2) each interaction center must be strongly coupled to an intense chromophore. Metalloporphyrins (e.g. Cu(II), Zn(II), Mn(II), Co(III), Cr(III), Sn(IV)-based porphyrins, etc.) and some pH sensitive dyes (e.g. methyl red, bromophenol blue, chlorophenol red, etc.) are among the primary compositions for sensors based on responsive dyes [162]. Similarly, colorimetric devices showed their merits in identifying odorous biomarkers [163]. Following the interaction of VOCs with an array of responsive dyes, a unique color map was generated for each VOC and different color patterns could be seen at different concentrations of VOCs (Figure 5d) [163]. Combined with a specific algorithm, excellent discrimination was obtained. LODs of the colorimetric devices made of responsive dyes have generally ranged from hundreds of ppb to several ppm. Due to the strong interaction between VOC molecules and responsive dyes, this kind of device can often be very sensitive to VOC biomarkers. In summary, these disposable responsive dye-based colorimetric devices have given interesting results for consideration in the future of skin-based wearable devices.
The rapid development of material science and engineering essentially stimulates progress in the development of wearable healthcare devices. Introduction of innovative sensing techniques has recently reformed our approach to disease diagnosis. It is undoubtedly reasonable to expect that real-time monitoring devices can give a much greater comprehensive assessment of the wearer’s health status and improve his or her quality of life. Indeed, many more organizations are paying attention to this field since continuous personalized health monitoring using wearable healthcare devices that can be expected to provide low-cost solutions especially for remote monitoring, and will be particularly useful in indicating efficient treatment where early diagnosis can be made. To meet the market demands, extensive progress must be made in this field, which relies on the collaboration between material sciences and electronics, as well as sensing signal acquisition and processing. The requirement for highly active sensing materials has spurred tremendous interests in developing monodisperse, single-component small-size NPs and encouraged researchers to put greater effort in searching for creative strategies in designing and synthesizing sensing elements with novel properties. Considering this goal, a better understanding of sensing mechanism can provide deep insight into the connection between sensing characteristics and materials composition/structure/morphology that will be instructive to scientists and engineers for exploring high-performance healthcare devices.
Although encouraging results have been achieved with a wide variety of wearable sensors, major obstacles remain that prevent the routine use of wearable devices by the general population. First, there is a lack of “bioaffinity” protocols [21], namely the relationship between these biomarkers and physiological index or a specific disease. Although some reports discuss this issue, the intrinsic connections remain ambiguous. Second, the precision, sensitivity, long-term stability and biocompatibility of devices must improve to meet the standard requirements from diagnostic devices, as defined by regulatory agencies. Third, there is a need to effectively reduce power consumption of these devices. In most situations, a sensing matrix should be built into these wearable devices for recording comprehensively body information, leading to considerable power consumption. Even though fabrication of self-powered devices suggests the way to solve this issue, power generation efficiency of this kind of devices remains questionable.
Although the importance of smart materials have been validated in designing high-performance, multifunctional, compact skin-based wearables, updates on the materials sciences still lags behind progress being made in the relevant fields. A typical example is that the biomaterials are less often reported in wearable healthcare devices, whereas this type of candidate will probably improve the current status. To date, biomaterials have been widely directed at their therapeutic applications; in particular they have shown unique features in the application of implantable healthcare devices [164, 165]. Owing to their natural biocompatibility and biodegradability, it is expected that skin-based wearable devices made of biomaterials will achieve even more unexpected results. One possible promotion is to replace the frequently used polymeric elastomer with more biocompatible biomaterials based substrates, for example, protein films, which will eliminate completely any irritation caused by the wearables. Their biodegradability will also be especially useful in designing environmental friendly products. As well as the more usual biomaterials, multifunctional materials also need more attention. As discussed in the section on multifunctional devices, these multifunctional materials that are spontaneously energy harvesting/self-healing and stimuli-responsive, or multi stimuli-responsive, will probably decrease the size, cost and the complexity of fabricating these healthcare devices. Although until now few relevant multifunctional candidates have been successfully devised, we speculate that wearables with smart multifunctional materials will lead to further innovative developments.
Overall, with advances in sensor sophistication and miniaturization, and battery solutions as well as materials sciences, much is expected to be packed into a patch. As with any wearable solution, there may need to be trade-offs in the number of data-points collected, the period of use and the processing power. Generally speaking, the closer the connection between the wearable device and the signal it is collecting, the greater will be the accuracy of the data. Because patches provide direct contact with the skin, they can detect very subtle changes in the patient’s physiological condition and give comprehensive profile of one’s health status.
References
1.Stoppa M, Chiolerio A. Wearable electronics and smart textiles: A critical review. Sensors. 2014;14:11957-11992
2.Tamsin M. Wearable biosensor technologies. International Journal of Scientific Research. 2015;13:697-703
3.Zeng W, Shu L, Li Q, Chen S, Wang F, Tao XM. Fiber-based wearable electronics: A review of materials, fabrication, devices, and applications. Advanced Materials. 2014;26:5310-5336
4.Ajami S, Teimouri F. Features and application of wearable biosensors in medical care. Journal of Research in Medical Sciences. 2015;20:1208-1215
5.Appelboom G, Camacho E, Abraham ME, Bruce SS, Dumont EL, Zacharia BE, et al. Smart wearable body sensors for patient self-assessment and monitoring. Archives of Public Health. 2014;72:1-9
6.Esteban M, Castaño A. Non-invasive matrices in human biomonitoring: A review. Environment International. 2009;35:438-449
7.Trung TQ, Ramasundaram S, Hwang BU, Lee NE. An all-elastomeric transparent and stretchable temperature sensor for body-attachable wearable electronics. Advanced Materials. 2016;28:502-509
8.Kantoch E, Augustyniak P. Human activity surveillance based on wearable body sensor network. Computing in Cardiology. 2012;39:325-328
9.Comina G, Suska A, Filippini D. Autonomous chemical sensing Interface for universal cell phone readout. Angewandte Chemie, International Edition. 2015;54:8708-8712
10.Munos B, Baker PC, Bot BM, Crouthamel M, Vries GD, Ferguson I, et al. Mobile health: The power of wearables, sensors, and apps to transform clinical trials. Annals of the New York Academy of Sciences. 2016;1375:3-18
11.Patel S, Park H, Bonato P, Chan L, Rodgers M. A review of wearable sensors and systems with application in rehabilitation. Journal of Neuroengineering and Rehabilitation. 2012;9:1-17
12.Nelson EC, Verhagen T, Noordzij ML. Health empowerment through activity trackers: An empirical smart wristband study. Computers in Human Behavior. 2016;62:364-374
13.Chuah HW, Rauschnabel PA, Krey N, Bang N, Ramayah T, Lade S. Wearable technologies: The role of usefulness and visibility in Smartwatch adoption. Computers in Human Behavior. 2016;65:276-284
14.Finlay DD, Nugent CD, Donnelly MP, Mccullagh PJ, Black ND. Optimal electrocardiographic lead systems: Practical scenarios in smart clothing and wearable health systems. IEEE Transactions on Information Technology in Biomedicine. 2008;12:433-441
15.Jung PG, Oh S, Lim G, Kong K. A mobile motion capture system based on inertial sensors and smart shoes. Journal of Dynamic Systems Measurement and Control. 2013;136:692-697
16.Kim SH, Ryoo DW, Bae C. U-healthcare system using smart headband. Proceedings of 30th Annual International IEEE Engineering in Medicine and Biology Society Conference, Vancouver, British Columbia, Canada. 2008; 1557-1560
17.Constant N, Douglas-Prawl O, Johnson S, Mankodiya K. Pulse-glasses: An unobtrusive, wearable HR monitor with internet-of-things functionality. 12th IEEE International Conference on Wearable & Implantable Body Sensor Networks. 2015
18.Khandwalla RM, Birkeland K, Zimmer R, Banet M, Pede S, Kedan I. Predicting heart failure events with home monitoring: Use of a novel wearable necklace to measure stroke volume, cardiac output and thoracic impedance. Journal of the American College of Cardiology. 2016;67:1296
19.Konvalina G, Haick H. Sensors for breath testing: From nanomaterials to comprehensive disease detection. Accounts of Chemical Research. 2013;47:66-76
20.Amjadi M, Kyung KU, Park I, Sitti M. Stretchable, skin-mountable, and wearable strain sensors and their potential applications: A review. Advanced Functional Materials. 2016;26:1678-1698
21.Bandodkar AJ, Wang J. Non-invasive wearable electrochemical sensors: A review. Trends in Biotechnology. 2014;32:363-371
22.Vashist SK. Non-invasive glucose monitoring technology in diabetes management: A review. Analytica Chimica Acta. 2012;750:16-27
23.Windmiller JR, Wang J. Wearable electrochemical sensors and biosensors : A review. Electroanalysis. 2012;24:1-18
24.Chortos A, Bao ZN. Skin-inspired electronic devices. Materials Today. 2014;17:321-331
25.Oliver NS, Toumazou C, Cass AEG, Johnston DG. Glucose sensors: A review of current and emerging technology. Diabetes Medication. 2009;26:197-210
26.So CF, Choi KS, Wong TKS, Chung JWY. Recent advances in noninvasive glucose monitoring. Journal of Medical Devices. 2012;5:45-52
27.Gao R, Emaminejad S, Nyein HYY, Challa S, Chen K, Peck A, et al. Fully integrated wearable sensor arrays for multiplexed in situ perspiration analysis. Nature. 2016;529:509-514
28.Windmiller JR, Bandodkar AJ, Parkhomovsky S, Wang J. Stamp transfer electrodes for electrochemical sensing on non-planar and oversized surfaces. Analyst. 2012;137:1570-1575
30.Guinovart T, Bandodkar AJ, Windmiller JR, Andrade FJ, Wang J. A potentiometric tattoo sensor for monitoring ammonium in sweat. Analyst. 2013;138:7031-7038
31.Ruiz JG, Mas R, Haro C, Cabruja E, Camero R, Lomillo MAA, et al. Early determination of cystic fibrosis by electrochemical chloride quantification in sweat. Biosensors and Bioelectronics. 2009;24:1788-1791
32.Solovei D, Žák J, Majzlíková P, Sedláček J, Hubálek J. Chemical sensor platform for non-invasive monitoring of activity and dehydration. Sensors. 2015;15:1479-1495
33.Paliwal S, Hwang BH, Tsai KY, Mitragotri S. Diagnostic opportunities based on skin biomarkers. European Journal of Pharmaceutical Sciences. 2013;50:546-556
34.Mishra A, Greaves R, Massie J. The relevance of sweat testing for the diagnosis of cystic fibrosis in the genomic era. Clinical Biochemist Reviews. 2005;26:135-153
35.Vermeulen F, Lebecque P, Boeck KD, Leal T. Biological variability of the sweat chloride in diagnostics sweat tests: A retrospective analysis. Journal of Cystic Fibrosis. 2016;15:S30-S35
36.Koh A, Kang D, Xue YG, Lee S, Pielak RM, Kim J, et al. A soft, wearable microfluidic device for the capture, storage, and colorimetric sensing of sweat. Science Translational Medicine. 2016;8:1-13
37.Schazmann B, Morris D, Slater C, Beirne S, Fay C, Reuveny R, et al. A wearable electrochemical sensor for the real-time measurement of sweat sodium concentration. Analytical Methods. 2010;2:342-348
38.Sanders JP, Loveday A, Pearson N, Edwardson C, Yates T, Biddle SJ, et al. Devices for self-monitoring sedentary time or physical activity: A scoping review. Journal of Medical Internet Research. 2016;18:e90
39.Page T. A forecast of the adoption of wearable technology. International Journal of Technology Diffusion. 2015;6:12-29
40.Paul BK, Panat R, Mastrangelo C, Kim D, Johnson D. Manufacturing of smart goods: Current state, future potential, and research recommendations. Journal of Micro and Nano-Manufacturing. 2016;4:044001-044011
41.Kim J, Lee M, Shim HJ, Ghaffari R, Cho HR, Son D, et al. Stretchable silicon Nanoribbon electronics for skin prosthesis. Nature Communications. 2014;5:5747
42.Kaltenbrunner M, Sekitani T, Reeder J, Yokota T, Kuribara K, Tokuhara T, et al. An ultra-lightweight Design for Imperceptible Plastic Electronics. Nature. 2013;499:458-463
43.Giuliani A, Placidi M, Francesco FD, Pucci A. A new polystyrene-based Ionomer/MWCNT nanocomposite for wearable skin temperature sensors. Reactive and Functional Polymers. 2014;76:57-62
44.Matzeu G, Pucci A, Savi S, Romanelli M, Francesco FD. A temperature sensor based on a MWCNT/SEBS nanocomposite. Sensors and Actuators A. 2012;178:94-99
45.Webb RC, Bonifas AP, Behnaz A, Zhang YH, Yu KJ, Cheng HY, et al. Ultrathin conformal devices for precise and continuous thermal characterization of human skin. Nature Materials. 2013;12:938-944
46.Yang J, Wei DP, Tang LL, Song JZ, Luo W, Chu J, et al. Wearable temperature sensor based on graphene nanowalls. RSC Advances. 2015;5:25609-25615
47.Yokota T, Inoue Y, Terakawa Y, Reeder J, Kaltenbrunner M, Ware T, et al. Ultraflexible, large-area, physiological temperature sensors for multipoint measurements. Proceedings of the National Academy of Sciences of the United States of America. 2015;112:14533-14538
48.Jeon J, Lee HBR, Bao ZN. Flexible wireless temperature sensors based on Ni microparticle-filled binary polymer composites. Advanced Materials. 2013;25:850-855
49.Yang H, Qi DP, Liu ZY, Chandran BK, Wang T, Yu JC, et al. Soft thermal sensor with mechanical adaptability. Advanced Materials. 2016;28:9175-9181
50.Trung TQ, Tien NYT, Kim D, Jung JH, Yoon OJ, Lee NE. High thermal responsiveness of a reduced graphene oxide field-effect transistor. Advanced Materials. 2012;24:5254-5260
51.Tien NT, Jeon D, Kim DI, Trung DD, Jang JS, Hwang BH, et al. A flexible bimodal sensor Array for simultaneous sensing of pressure and temperature. Advanced Materials. 2014;26:796-804
52.Ren XC, Chan PKL, Lu JB, Huang BL, Leung DCW. High dynamic range organic temperature sensor. Advanced Materials. 2013;25:1291-1295
53.Hattori Y, Falgout L, Lee E, Jung JH, Poon E, Lee E, et al. Multifunctional skin-like electronics for quantitative clinical monitoring of cutaneous wound healing. Advanced Healthcare Materials. 2014;3:1597-1607
54.Huynh TP, Haick H. Self-healing, fully functional, and multiparametric flexible sensing platform. Advanced Materials. 2015;28:138-143
55.Haick H, Ambrico M, Ghabboun J, Ligonzo T, Cahen D. Contacting organic molecules by metal evaporation. Physical Chemistry Chemical Physics. 2004;6:4538-4541
56.Haick H, Ghabboun J, Cahen D. Pd versus au as evaporated metal contacts to molecules. Applied Physics Letters. 2005;86:042113/1-042113/3
57.Wang LH, Li YL. A review for conductive polymer piezoresistive composites and a development of a compliant pressure transducer. IEEE Transactions on Instrumentation and Measurement. 2013;6:495-502
58.Zhang MH, Chen JK, Lu CS. Influence of strain rate on the piezoresistive behavior of conductive polyamide composites. Composites Science and Technology. 2016;133:1-6
59.Choong CL, Shim MB, Lee BS, Jeon S, Ko DS, Kang TH, et al. Highly stretchable resistive pressure sensors using a conductive elastomeric composite on a micropyramid Array. Advanced Materials. 2014;26:3451-3458
60.Khang DY, Jiang HQ, Huang Y, Rogers JA. A stretchable form of single-crystal silicon for high-performance electronics on rubber substrates. Science. 2006;311:208-212
61.Webb RC, Pielak RM, Bastien P, Ayers J, Niittynen J, Kurniawan J, et al. Thermal transport characteristics of human skin measured in vivo using ultrathin conformal arrays of thermal sensors and actuators. PLoS One. 2015;10:e0118131
62.Cavanagh PR, Rodgers MM, Iiboshi A. Pressure distribution under symptom-free feet during barefoot standing. Foot & Ankle International. 1987;7:262-276
63.Wald H, Guernsey M, Scott FH. Some effects of alteration of posture on arterial blood pressure. American Heart Journal. 1937;14:319-330
64.Schwartz G, Tee BC-K, Mei J, Appleton AL, Kim DH, Wang H, et al. Flexible polymer transistors with high pressure sensitivity for application in electronic skin and health monitoring. Nature Communications. 2013;4:1859
65.Tiwana MI, Redmond SJ, Lovell NH. A review of tactile sensing technologies with applications in biomedical engineering. Sensors and Actuators. 2012;179:17-31
66.Amjadi M, Kyung KU, Park I, Sitti M. Stretchable, skin-mountable, and wearable strain sensors and their potential applications: A review. Advanced Functional Materials. 2016;26:1678-1698
67.Tai YL, Mulle M, Ventura IA, Lubineau G. A highly sensitive, low-cost, wearable pressure sensor based on conductive hydrogel spheres. Nanoscale. 2013;7:14766-14773
68.Boutry CM, Nguyen A, Lawal QO, Chortos A, Rondeaugagné S, Bao Z. A sensitive and biodegradable pressure sensor Array for cardiovascular monitoring. Advanced Materials. 2015;43:6954-6961
69.Park S, Jayraman S. Smart textile-based wearable biomedical systems: A transition plan for research to reality. IEEE Transactions on Information Technology in Biomedicine. 2010;14:86-92
70.Windmiller JR, Bandodkar AJ, Ramirez FH, Parkhomovsky S, Martinez AG, Wang B. Electrochemical sensing based on printable temporary transfer tattoos. Chemical Communications. 2012;48:6794-6796
71.Jia WZ, Bandodkar AJ, Ramírez GV, Windmiller JR, Yang ZJ, Ramírez JL, et al. Electrochemical tattoo biosensors for real-time noninvasive lactate monitoring in human perspiration. Analytical Chemistry. 2013;85:6553-6560
72.Amal H, Leja M, Funka K, Skapars R, Sivins A, Ancans G, et al. Detection of precancerous gastric lesions and gastric cancer through exhaled breath. Gut. 2016;65:400-407
73.Park IJ, Jeong CY, Cho IT, Lee JH, Cho ES, Kwon SJ, et al. Fabrication of amorphous InGaZnO thin-film transistor-driven flexible thermal and pressure sensors. Semiconductor Science and Technology. 2012;27:105019-105024
74.Wang Y, Wang L, Yang TT, Li X, Zang XB, Zhu M, et al. Wearable and highly sensitive graphene strain sensors for human motion monitoring. Advanced Functional Materials. 2014;24:4666-4670
75.Li YQ, Samad YA, Liao K. From cotton to wearable pressure sensor. Journal of Materials Chemistry A. 2014;3:2181-2187
76.Gong S, Schwalb W, Wang YW, Chen Y, Tang Y, Si J, et al. A wearable and highly sensitive pressure sensor with ultrathin gold nanowires. Nature Communications. 2014;5:163-180
77.Stassi S, Cauda V, Canavese G, Pirri CF. Flexible tactile sensing based on piezoresistive composites: A review. Sensors. 2014;14:5296-5332
78.Segev-Bar M, Konvalina G, Haick H. High-resolution unpixelated smart patches with antiparallel thickness gradients of nanoparticles. Advanced Materials. 2015;27:1779-1784
79.Yao SS, Zhu Y. Wearable multifunctional sensors using printed stretchable conductors made of silver nanowires. Nanoscale. 2014;6:2345-2352
80.Lipomi DJ, Vosgueritchian M, Tee BC-K, Hellstrom SL, Lee JA, Fox CH, et al. Skin-like pressure and strain sensors based on transparent elastic films of carbon nanotubes. Nature Nanotechnology. 2011;6:788-792
81.Kwon D, Lee TI, Shim J, Ryu S, Kim MS, Kim S, et al. Highly sensitive, flexible, and wearable pressure sensor based on a giant piezocapacitive effect of three-dimensional microporous elastomeric dielectric layer. ACS Applied Materials & Interfaces. 2016;8:16922-16931
82.Tee BC-K, Chortos A, Dunn RR, Schwartz G, Eason E, Bao Z. Tunable flexible pressure sensors using microstructured elastomer geometries for intuitive electronics. Advanced Functional Materials. 2014;24(34):5427
83.Pan LJ, Chortos A, Yu GH, Wang YQ, Isaacson S, Allen R, et al. An ultra-sensitive resistive pressure sensor based on hollow-sphere microstructure induced elasticity in conducting polymer film. Nature Communications. 2013;5:3002
84.Liu HY, Qi DP, Guo PZ, Liu Y, Zhu BW, Yang H, et al. Thickness-gradient films for high gauge factor stretchable strain sensors. Advanced Materials. 2015;27:6230-6237
85.Zhu BW, Niu ZQ, Wang H, Leow WR, Wang H, Li YG, et al. Microstructured graphene arrays for highly sensitive flexible tactile sensors. Small. 2014;10:3625-3631
86.Anton SR, Sodano HA. A review of power harvesting using piezoelectric materials. Smart Materials and Structures. 2007;16:R1-R21
87.Sun KT, Zhang HQ, Hu LZ, Yu DQ, Qiao SS, Sun JC, et al. Growth of Ultralong ZnO microwire and its application in isolatable and flexible piezoelectric strain sensor. Physica Status Solidi A. 2010;207:488-492
88.Dagdeviren C, Shi Y, Joe P, Ghaffari R, Balooch G, Usgaonkar K, et al. Conformal piezoelectric systems for clinical and experimental characterization of soft tissue biomechanics. Nature Materials. 2015;14:728-736
89.Gullapalli H, Vemuru VSM, Kumar A, Mendez AB, Vajtai R, Terrones M, et al. Flexible piezoelectric ZnO-paper nanocomposite strain sensor. Small. 2010;6:1641-1646
90.Sharma T, Je SS, Gill B, Zhang JXJ. Patterning piezoelectric thin film PVDF-TrFE based pressure sensor for catheter application. Sensors and Actuators A. 2012;177:87-92
91.Persano L, Dagdeviren C, Su YW, Zhang YH, Girardo S, Pisignano D, et al. High performance piezoelectric devices based on aligned arrays of nanofibers of poly(vinylidenefluoride-co-trifluoroethylene). Nature Communications. 2013;4:67-88
92.Amal H, Shi D-Y, Ionescu R, Zhang W, Hua Q-L, Pan Y-Y, et al. Assessment of ovarian cancer conditions from exhaled breath. International Journal of Cancer. 2015;136:E614-EE22
93.Amann A, Mochalski P, Ruzsanyi V, Broza YY, Haick H. Assessment of the exhalation kinetics of volatile cancer biomarkers based on their physicochemical properties. Journal of Breath Research. 2014;8:016003
94.Bachar N, Mintz L, Zilberman Y, Ionescu R, Feng X, Mullen K, et al. Polycyclic aromatic hydrocarbon for the detection of nonpolar analytes under counteracting humidity conditions. ACS Applied Materials & Interfaces. 2012;4:4960-4965
95.Barash O, Peled N, Tisch U, Bunn PA, Hirsch FR, Haick H. Classification of the lung cancer histology by gold nanoparticle sensors. Nanomedicine (New York, NY, US). 2012;8:580-589
96.Marozas V, Petrenas A, Daukantas S, Lukosevicius A. A comparison of conductive textile-based and silver/silver chloride gel electrodes in exercise electrocardiogram recordings. Journal of Electrocardiology. 2011;44:189-194
97.Baek JY, An JH, Choi JM, Park KS, Lee SH. Flexible polymeric dry electrodes for the long-term monitoring of ECG. Sensors and Actuators A: Physical. 2008;143:423-429
98.Xu PJ, Zhang H, Tao XM. Textile-structured electrodes for electrocardiogram. Tex Prog. 2008;40:183-213
99.Cizza G, Marques AH, Eskandari F, Christie IC, Torvik S, Silverman MN, et al. Elevated Neuroimmune biomarkers in sweat patches and plasma of premenopausal women with major depressive disorder in remission: The POWER study. Biological Psychiatry. 2008;64:907-911
100.Tamer YYA, Hadi EA, Badrani IEIA. Sweat urea, uric acid and Creatinine concentrations in uraemic patients. Urological Research. 1997;25:337-340
101.Costill DL. Water and electrolyte requirements during exercise. Clinics in Sports Medicine. 1984;3:639-648
102.Guinovart T, Parrilla M, Crespo GA, Riusa FX, Andrade FJ. Potentiometric sensors using cotton yarns, carbon nanotubes and polymeric membranes. Analyst. 2013;138:5208-5215
103.Weber J, Kumar A, Kumar A, Bhansali S. Novel lactate and pH biosensor for skin and sweat analysis based on single walled carbon nanotubes. Sensors and Actuators B. 2006;17:308-313
104.Kim J, Jeerapan I, Imani S, Cho TN, Bandodkar A, Cinti S, et al. Noninvasive alcohol monitoring using a wearable tattoo-based Iontophoretic-biosensing system. ACS Sensors. 2016;1:1011-1019
105.Yang YL, Chuang MC, Lou SL, Wang J. Thick-film textile-based amperometric sensors and biosensors. Analyst. 2012;135:1230-1234
106.Lee H, Choi1 TK, Lee YB, Cho HR, Ghaffari R, Wang L, et al. A graphene-based electrochemical device with thermoresponsive microneedles for diabetes monitoring and therapy. Nature Nanotechnology. 2016;11:566-572
107.Imani S, Bandodkar AJ, Mohan AMV, Kumar R, Yu SF, Wang J, et al. A wearable chemical-electrophysiological hybrid biosensing system for real-time health and fitness monitoring. Nature Communications. 2016;7:11650
108.Broza YY, Mochalski P, Ruzsanyi V, Amann A, Haick H. Hybrid volatolomics and disease detection. Angewandte Chemie, International Edition. 2015;54:11036-11048
109.Broza YY, Zuri L, Haick H. Combined volatolomics for monitoring of human body chemistry. Scientific Reports. 2014;4:4611-4617
110.Haick H, Broza YY, Mochalski P, Ruzsanyi V, Amann A. Assessment, origin, and implementation of breath volatile cancer markers. Chemical Society Reviews. 2014;43:1423-1449
111.Jin H, Huynh TP, Haick H. Self-healable sensors based nanoparticles for detecting physiological markers via skin and breath: Toward disease prevention via wearable devices. Nano Letters. 2016;16:4194-4202
112.Nakhleh MK, Amal H, Jeries R, Broza YY, Aboud M, Gharra A, Ivgi H, Khatib S, Badarneh S, Har-Shai L, Glass-Marmor L, Lejbkowicz I, Miller A, Badarny S, Winer R, Finberg J, Cohen-Kaminsky S, Perros F, Montani D, Girerd B, Garcia G, Simonneau G, Nakhoul F, Baram S, Salim R, Hakim M, Gruber M, Ronen O, Marshak T, Doweck I, Nativ O, Bahouth Z, Shi DY, Zhang W, Hua QL, Pan YY, Tao L, Liu H, Karban A, Koifman E, Rainis T, Skapars R, Sivins A, Ancans G, Liepniece-Karele I, Kikuste I, Lasina I, Tolmanis I, Johnson D, Millstone SZ, Fulton J, Wells JW, Wilf LH, Humbert M, Leja M, Peled N, Haick H. Diagnosis and classification of 17 diseases from 1404 subjects via pattern analysis of exhaled molecules. ACS Nano. 2017;11:112-125
113.Peng G, Tisch U, Adams O, Hakim M, Shehada N, Broza YY, et al. Diagnosing lung cancer in exhaled breath using au nanoparticles. Nature Nanotechnology. 2009;4:669-673
114.Amal H, Leja M, Funka K, Lasina I, Skapars R, Sivins A, et al. Breath testing as potential colorectal cancer screening tool. International Journal of Cancer. 2016;138:229-236
115.Barash O, Zhang W, Halpern JM, Hua QL, Pan YY, Kayal H, et al. Differentiation between genetic mutations of breast cancer by breath volatolomics. Oncotarget. 2015;6:44864-44876
116.Cohen-Kaminsky S, Nakhleh M, Perros F, Montani D, Girerd B, Garcia G, et al. A proof of concept for the detection and classification of pulmonary arterial hypertension through breath analysis with a sensor Array. American Journal of Respiratory and Critical Care Medicine. 2013;188:756-759
117.Davies MP, Barash O, Jeries R, Peled N, Ilouze M, Hyde R, et al. Unique volatolomic signatures of TP53 and KRAS in lung cells. British Journal of Cancer. 2014;111:1213-1221
118.Hakim M, Billan S, Tisch U, Peng G, Dvrokind I, Marom O, et al. Diagnosis of head and neck cancer from exhaled breath. British Journal of Cancer. 2011;104:1649-1655
119.Hakim M, Broza YY, Barash O, Peled N, Phillips M, Amann A, et al. Volatile organic compounds of lung cancer and possible biochemical pathways. Chemical Reviews. 2012;112:5949-5966
120.Homede E, Abo Jabal M, Ionescu R, Haick H. Printing ultrasensitive artificially intelligent sensors array with a single self-propelled droplet containing nanoparticles. Advanced Functional Materials. 2016;26:6359-6370
121.Ionescu R, Broza Y, Shaltieli H, Sadeh D, Zilberman Y, Feng X, et al. Detection of multiple sclerosis from exhaled breath using bilayers of polycyclic aromatic hydrocarbons and Single-Wall carbon nanotubes. ACS Chemical Neuroscience. 2011;2:687-693
122.Karban A, Nakhleh MK, Cancilla JC, Vishinkin R, Rainis T, Koifman E, et al. Programmed nanoparticles for tailoring the detection of inflammatory bowel diseases and irritable bowel syndrome disease via Breathprint. Advanced Healthcare Materials. 2016;5:2339-2344
123.Konvalina G, Haick H. Sensors for breath testing: From nanomaterials to comprehensive disease detection. Accounts of Chemical Research. 2014;47:66-76
124.Marom O, Nakhoul F, Tisch U, Shiban A, Abassi Z, Haick H. Gold nanoparticle sensors for detecting chronic kidney disease and disease progression. Nanomedicine (London, UK). 2012;7:639-650
125.Nakhleh M, Broza YY, Haick H. Monolayer-capped gold nanoparticles for disease detection from breath. Nanomedicine (London, UK). 2014;9:1991-2002
126.Nakhleh MK, Amal H, Awad H, Gharra A, Abu-Saleh N, Jeries R, et al. Sensor arrays based on nanoparticles for early detection of kidney injury by breath samples. Nanomedicine (New York, NY, US). 2014;10:1767-1776
127.Nakhleh MK, Badarny S, Winer R, Jeries R, Finberg J, Haick H. Distinguishing idiopathic Parkinson's disease from other Parkinsonian syndromes by breath test. Parkinsonism & Related Disorders. 2015;21:150-153
128.Nakhleh MK, Baram S, Jeries R, Salim R, Haick H, Hakim M. Artificially intelligent nanoarray for the detection of preeclampsia under real-world clinical conditions. Advanced Materials Technology. 2016;1:1600132-1600136
129.Nakhleh MK, Jeries R, Gharra A, Binder A, Broza YY, Pascoe M, et al. Detecting active pulmonary tuberculosis by breath test using nanomaterial-based sensors. The European Respiratory Journal. 2014;43(5):1522
130.Nardi-Agmon I, Abud-Hawa M, Liran O, Gai-Mor N, Ilouze M, Onn A, et al. Exhaled breath analysis for monitoring response to treatment in advanced lung cancer. Journal of Thoracic Oncology. 2016;11:827-837
131.Peled N, Barash O, Tisch U, Ionescu R, Broza YY, Ilouze M, et al. Volatile fingerprints of cancer specific genetic mutations. Nanomedicine (New York, NY, US). 2013;9:758-766
132.Peled N, Hakim M, Bunn Jr PA, Miller YE, Kennedy TC, Mattei J, et al. Non-invasive breath analysis of pulmonary nodules. Journal of Thoracic Oncology. 2012;7:1528-1533
133.Peng G, Hakim M, Broza YY, Billan S, Abdah-Bortnyak R, Kuten A, et al. Detection of lung, breast, colorectal, and prostate cancers from exhaled breath using a single Array of nanosensors. British Journal of Cancer. 2010;103:542-551
134.Peng G, Tisch U, Adams O, Hakim M, Shehada N, Broza YY, et al. Diagnosing lung cancer in exhaled breath using gold nanoparticles. Nature Nanotechnology. 2009;4:669-673
135.Shehada N, Brönstrup G, Funka K, Christiansen S, Leja M, Haick H. Ultrasensitive silicon nanowire for real-world gas sensing: Noninvasive diagnosis of cancer from breath volatolome. Nano Letters. 2014;15:1288-1295
136.Shehada N, Cancilla JC, Torrecilla JS, Pariente ES, Bronstrup G, Christiansen S, et al. Silicon nanowire sensors enable diagnosis of patients via exhaled breath. ACS Nano. 2016;10:7047-7057
137.Shuster G, Gallimidi Z, Reiss AH, Dovgolevsky E, Billan S, Abdah-Bortnyak R, et al. Classification of breast cancer precursors through exhaled breath. Breast Cancer Research and Treatment. 2011;126:791-796
138.Tisch U, Haick H. Nanomaterials for cross-reactive sensor arrays. MRS Bulletin. 2010;35:797-803
139.Vishinkin R, Haick H. Nanoscale sensor technologies for disease detection via volatolomics. Small. 2015;11:6142-6164
140.Wang B, Cancilla JC, Torrecilla JS, Haick H. Artificial sensing intelligence with silicon nanowires for ultraselective detection in the gas phase. Nano Letters. 2014;14:933-938
141.Wang B, Huynh T-P, Wu W, Hayek N, Do TT, Cancilla JC, et al. Highly sensitive ambipolar field effect transistor-based diketopyrrolopyrrole copolymer for selective detection and discrimination of xylene isomers. Advanced Materials. 2016;28:4012-4018
142.Broza Y, Kremer R, Tisch U, Gevorkyan A, Abdah-Bortnyak R, Shiban A, et al. A nanomaterial-based breath test for short-term follow-up after lung tumor resection. Nanomedicine (New York, NY, US). 2013;9:15-21
143.Tisch U, Schlesinger I, Ionescu R, Nassar M, Axelrod N, Robertman D, et al. Detection of Alzheimer’s and Parkinson’s diseases from exhaled breath using nanomaterial-based sensors. Nanomedicine (London, UK). 2013;8:43-56
144.Dovgolevsky E, Konvalina G, Tisch U, Haick H. Monolayer-capped cubic platinum nanoparticles for sensing nonpolar analytes in highly humid atmospheres. Journal of Physical Chemistry C. 2010;114:14042-14049
145.Dovgolevsky E, Tisch U, Haick H. Chemically sensitive resistors based on monolayer-capped cubic nanoparticles: Towards configurable nanoporous sensors. Small. 2009;5:1158-1161
146.Zaretski AV, Root SE, Savchenko A, Molokanova E, Printz AD, Jibril L, et al. Metallic nanoislands on graphene as highly sensitive transducers of mechanical, biological, and optical signals. Nano Letters. 2016;16:1375-1380
147.Korotcenkov G. Metal oxides for solid-state gas sensors: What determines our choice? Materials Science and Engineering B. 2007;139:1-23
148.Hangarter CM, Chartuprayoon N, Hernández SC, Choa YH, Myung NV. Hybridized conducting polymer chemiresistive Nano-sensors. Nano Today. 2013;8:39-55
149.Yamazoe N, Sakai G, Shimanoe K. Oxide semiconductor gas sensors. Catalysis Surveys from Asia. 2003;7:63-75
150.Li ZP, Zhao QQ, Fan WL, Zhan JH. Porous SnO2 nanospheres as sensitive gas sensors for volatile organic compounds detection. Nanoscale. 2011;3:1646-1652
151.Wang CX, Yin LW, Zhang LY, Xiang D, Gao R. Metal oxide gas sensors: Sensitivity and influencing factors. Sesnors. 2010;10:2088-2106
152.Zhang Y, Li JP, An GM, He XL. Highly porous SnO2 fibers by electrospinning and oxygen plasma etching and its ethanol-sensing properties. Sensors and Actuators B. 2010;144:43-48
153.Jang JS, Choi JM, Kim SJ, Kim ID. Rational design of highly porous SnO2 nanotubes functionalized with biomimetic nanocatalysts for direct observation of simulated diabetes. Advanced Functional Materials. 2016;26:4740-4748
154.Khoang ND, Trung DD, Duy NV, Hoa ND, Hieu NV. Design of SnO2/ZnO hierarchical nanostructures for enhanced ethanol gas-sensing performance. Sensors and Actuators B. 2012;174:594-601
155.Huang JR, Wang LY, Gu CP, Wang ZJ, Sun YF, Shim JJ. Preparation of porous SnO2 microcubes and their enhanced gas-sensing property. Sensors and Actuators B. 2015;207:782-790
156.Sun P, Mei XD, Cai QY, Ma J, Sun JC, Liang XH, et al. Synthesis and gas sensing properties of hierarchical SnO2 nanostructures. Sensors and Actuators B. 2013;187:301-307
157.Huang JR, Xu XJ, Gu CP, Wang WZ, Geng BY, Sun YF, et al. Effective VOCs gas sensor based on porous SnO2 microcubes prepared via spontaneous phase segregation. Sensors and Actuators B. 2012;173:599-606
158.Tomer VK, Duhan S. Ordered mesoporous Ag-doped TiO/SnO2 nanocomposite based highly sensitive and selective VOCs sensor. Journal of Materials Chemistry A. 2016;4:1033-1043
159.Shankar P, Rayappa JBB. Gas sensing mechanism of metal oxides: The role of ambient atmosphere, type of semiconductor and gases – A review. Science Letters. 2015;4:1-18
160.Bai H, Shi GQ. Gas sensors based on conducting polymers. Sesnors. 2007;7:267-307
161.Kolmakov A, Moskovits M. Chemical sensing and catalysis by one-dimensional metal-oxide nanostructures. Annual Review of Materials Research. 2004;35:151-180
162.Askim JR, Mahmoudi M, Suslick KS. Optical sensor arrays for chemical sensing: The optoelectronic nose. Chemical Society Reviews. 2013;42:8649-8682
163.Han F, Inoue T, Wasai T, Kurauchi Y. K. Ohga, Odor sensor based on the information of colour. Journal of Engineering and Applied Sciences. 2006;1:1-6
164.Geelhood SJ, Horbett TA, Ward WK, Wood MD, Quinn MJ. Passivating protein coatings for implantable glucose sensors: Evaluation of protein retention. Journal of Biomedical Materials Research. Part B, Applied Biomaterials. 2007;81:251-260
165.Sokolov A, Hellerud BC, Lambris JD, Johannessen EA, Mollnes TE. Activation of polymorphonuclear leukocytes by candidate biomaterials for an implantable glucose sensor. Journal of Diabetes Science and Technology. 2011;5:1490-1498
Written By
Han Jin, Qinghui Jin and Jiawen Jian
Submitted: 02 March 2018Reviewed: 19 March 2018Published: 03 October 2018
Open access peer-reviewed chapter
