flexible sensors

IEEE REVIEWS IN BIOMEDICAL ENGINEERING, VOL. 7, 2014 73

Flexible Sensors for Chronic Wound ManagementManuel Ochoa, Student Member, IEEE, Rahim Rahimi, and Babak Ziaie, Senior Member, IEEE

Methodological Review

Abstract—Chronic nonhealing wounds are a major source ofmorbidity and mortality in bed-ridden and diabetic patients. Mon-itoring of physical and chemical parameters important in woundhealing and remodeling process can be of immense benefit for op-timum management of such lesions. Low-cost flexible polymericand paper-based substrates are attractive platforms for fabrica-tion of such sensors. In this review, we discuss recent advancesin flexible physiochemical sensors for chronic wound monitoring.After a brief introduction to wound healing process and commer-cial wound dressings, we describe various flexible biocompatiblesubstrates that can be used as the base platform for integration ofwound monitoring sensors. We will then discuss several fabrica-tion methods that can be utilized to integrate physical and chemicalsensors onto such substrates. Finally, we will present physical andchemical sensors developed for monitoring wound microenviron-ment and outline future development venues.

Index Terms—BioFlex, biosensor, chronic wound, flexible,wound healing.

I. INTRODUCTION

CHRONIC wounds present a serious health burden and area source of staggering health care costs in the US and

developed countries. It is estimated that 6 million people inthe US are affected by chronic wounds, costing the health caresystem $20 billion annually [1]. These numbers will certainlyincrease in the coming decades with the aging population andthe increase in the prevalence of diabetes mellitus and its associ-ated cutaneous ulcers. A chronic wound is a type of wound thatdoes not heal in an orderly or timely manner; more specifically,wounds that do not heal within three months are consideredto be chronic. Wound healing process is tightly controlled bycytokines and typically progresses through distinct inflamma-tory, proliferative, and maturation phases (see Fig. 1). Over thepast few decades, there has been a growing understanding ofthe complex biology behind wound healing with major impli-cations for nonhealing wounds [2], [3]. However, therapies thatutilize a single physical or chemical parameter to enhance thehealing process (e.g., electrical stimulation, ultrasonic stimu-lation, oxygenation, cell, and growth factor therapy) have hadlimited success [4]–[10]. For example, hyperbaric oxygen ther-apies for foot ulcers are expensive, cumbersome, can result insystemic toxicity, and have shown marginal benefits. A major

Manuscript received September 12, 2013; revised December 16, 2013;accepted December 16, 2013. Date of publication December 18, 2013; dateof current version April 28, 2014. This work was supported by the NationalScience Foundation under Grant EFRI-BioFlex #1240443.

The authors are with the School of Electrical and Computer Engineering, Pur-due University, West Lafayette, IN 47907 USA (e-mail: [email protected];[email protected]; [email protected]).

Digital Object Identifier 10.1109/RBME.2013.2295817

Fig. 1. Sequence of molecular and cellular events in skin wound healing [28].Healing occurs in four phases: hemostasis (immediate), inflammation (days),repair (weeks), and remodeling (months). Suboptimal healing conditions inchronic wounds impede efficient progression of these phases, thus prolonging(or preventing) healing.

shortcoming associated with such methods is their open-loopnature, i.e., they do not use physical or chemical feedback inorder to modify or adjust the treatment based on wound physio-chemical microenvironment. Current treatment methods wouldgreatly benefit from incorporation of flexible physical and chem-ical sensors that can assess the local wound microenvironmentin order to create a therapeutic feedback loop.

Sensors and actuators fabricated on flexible substrates areparticularly attractive for wound monitoring and manipulationsince they can conformally cover the wound and do not applyexcessive force/stress to the healing area (most commerciallyavailable wound dressing are flexible and conformal). Initial ef-forts in flexible electronics started with the discovery of conduc-tive polymers [11] and their subsequent application in large areadisplays [12]–[14]. Although lower cost and room-temperatureprocessing conditions were the original motivations behind us-ing organic transistors for display applications, very soon it wasrealized that flexible displays can offer many attractive featuresunattainable with rigid substrates (conformability, compactness,foldability, etc.) [15]. With the introduction of soft lithographyin the 1990s, the field of flexible electronics entered a phaseof rapid growth. Polymeric substrates such as PDMS, poly-imide, and Parylene became widely utilized to make a varietyof sensors, actuators, active electronics, and microfluidic com-ponents [16]–[20]. In the biomedical area, flexible recordingand stimulating electrodes provided a less physically damag-ing interface to the nervous system and stretchable electrodesand interconnects allowed for a more dynamic interface to thetissue [21]–[25]. More recent efforts in this area have leveraged

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74 IEEE REVIEWS IN BIOMEDICAL ENGINEERING, VOL. 7, 2014

Fig. 2. Various types of commercially available wound dressings. (a) Commongauze pads/rolls [203], (b) transparent film [204], (c) hydrocolloid dressing[205], (d) foam dressing [206], (e) silver-loaded alginate dressing [207], and(f) soft silicone-coated dressing [208].

large scale integration of ultra-thin inorganic semiconductorswith flexible and biodegradable substrates to fabricate severalmulti-functional systems for biological applications [26], [27].

This review discusses various flexible platform technolo-gies and sensors developed for wound healing applications us-ing modern microfabrication technologies. Section II discussescommercial wound dressings, the potential benefits resultingfrom integration of sensors onto such platforms, and sensorrequirements. Section III summarizes common biocompatiblepolymeric materials and their key properties. Section IV reviewsestablished and emerging methods for fabrication of sensors onpolymeric substrates. Section V elaborates on recent efforts to-wards fabrication of flexible sensors for wound healing. Finally,Section VI concludes with suggestions for further research anddevelopment in this area.

II. IMPROVING CHRONIC WOUND CARE WITH

FLEXIBLE SENSORS

A. Current Wound Dressings and Potential Enhancements

Modern protocols for the treatment of chronic wounds callfor attentive care and regular application and replacement ofwound dressings. Their primary purpose is to physically protectthe wound and maintain adequate moisture levels while allowingair exchange [28]–[31]. Since the exact optimal healing envi-ronment for a particular wound varies depending on its severityand healing state, various types of dressings are available fortreating a wide range of wounds (see Fig. 2) [28], [32]–[34].Each of these addresses a unique condition in the wound andeach follows a different replacement schedule (from twice perday to twice per week). The most common and least expen-sive is the standard gauze [see Fig. 2(a)], which is indicatedfor shallow wounds or as a base component in multi-layerdressings. For increased moisture retention, a porous transparentfilm dressing [see Fig. 2(b)] may be used instead or in addition

to the gauze. Such dressings permit gas exchange but preventwound-bed dry-out. If, on the other hand, the wound-bed is ex-cessively wet, one may opt for a more absorbent dressing suchas a gel or a hydrocolloid [see Fig. 2(c)]. These materials ab-sorb excess exudate but maintain the wound sufficiently moistto promote tissue regeneration. In more extreme cases of deeperor more exudate-prone wounds, a sponge-based dressing is thepreferred alternative [see Fig. 2(d)]; such dressings can moreeffectively fill a deep wound and absorb significant exudate. Inaddition to moisture management, many chronic wounds alsorequire infection control. These cases would benefit from theapplication of antiseptic-loaded carriers such as silver-ion algi-nate dressings [see Fig. 2(e)] [35]. This broad variety of moderndressings may seem sufficient for targeting many wound con-ditions; however, the high complexity of chronic wounds oftenresults in the simultaneous expression of multiple symptomsthat can be more efficiently addressed with improved woundmonitoring systems.

With wound dressings being the primary modality for chronicwound management, a promising approach to enhancing chronicwound care is to improve their effectiveness by using modernsensing and drug delivery technologies. This can be accom-plished by the incorporation of bio/chemical sensors (e.g., pH,oxygen) and drug delivery capabilities (e.g., via nanomedicine[36] or microfluidics [37]) into the dressings [38], [39] by ei-ther: 1) developing new wound dressing platforms; or 2) em-bedding novel flexible sensors into existing commercial ones.These smart systems can help optimize the healing process,decrease the healing time, and prevent infections. They canevaluate the local wound environment, release wound healingagents as needed, detect the optimal replacement time, andalert the patient/caregiver of any unusual phenomena (prefer-ably through a wireless link). Although the unit cost of suchsystems may surpass that of current dressings, their ultimate ef-ficacy can result in an overall reduced expenditure. As reportedby Kerstein et al. [40], while a pressure ulcer that is properlytreated shows satisfactory healing progress with either inexpen-sive gauze dressings or more expensive hydrocolloid ones, thetotal cost of the treatment is lower for the latter due to their lessfrequent replacement schedule and a reduced need for profes-sional medical attention.

B. Required Properties for Wound Dressings and Sensors

The effectiveness of new wound dressings depends on care-ful material and physical design. Despite the large variation instructural and physical properties of current wound dressings,certain elements remain core requirements for their functionalefficacy. First is flexibility, i.e., such dressings must be suf-ficiently flexible to conform to the wound and not limit thepatient’s mobility [41]. Second is gas permeability, which is es-sential for maintaining an adequate oxygen supply; alternatively,some dressings may supply oxygen at required levels [42]. Thirdis moisture control [43]; the dressing should keep the wound bedmoist but absorb excess exudate. Finally, the material in con-tact with the wound bed should be sufficiently soft to avoidcausing mechanical insults and interfering with the

OCHOA et al.: FLEXIBLE SENSORS FOR CHRONIC WOUND MANAGEMENT 75

epithelization process [a minimal adhesive is often preferredto reduce the mechanical load, e.g., Fig. 2(f)].

In addition to the above standard requirements, the emerg-ing sensor-equipped dressings must provide enhanced capa-bilities such as wound status reporting and (semi)automatedmicro-environment control. To achieve this goal, wound sen-sors require physical and operational specifications that differfrom those of typical industrial and process control sensors.For example, although sensitivity and specificity remain criticalfor wound sensors, other parameters such as working life andresponse time are less critical due to the disposable nature ofwound dressings and the relatively slow rate of biological woundhealing processes. In accordance with typical wound dressingreplacement schedules, most embedded sensors require a work-ing life of at most a few days to a week, after which they mustbe discarded. Hence, sensor deterioration and biofouling posea reduced threat to wound sensor design (unlike long-term im-plantable sensors). Additionally, response times in the range oftens of seconds or even a few minutes are often acceptable,depending on the sensing parameter. The level of oxygen, forexample, or pH in a wound is not expected to drastically changewithin such time windows. Similarly, their spatial density doesnot need micrometer resolution. As a result of such less strin-gent requirements, wound sensor fabrication can deviate fromtraditional MEMS or microfabrication techniques. Instead, theycan take advantage of emerging rapid prototyping technologies(as described in Section IV) that allow the processing of moreunusual materials while simultaneously enabling the personal-ization of wound dressing geometries and sensor distribution.Finally, the sensors must not impair the flexibility of the wounddressing; hence, it is favorable to fabricate such sensors on flex-ible substrates (instead of glass/silicon/metal). Such substratesmay include polymers (e.g., polyimide, Parylene) or nonwovenfiber mesh (e.g., paper, fabrics).

III. BIOCOMPATIBLE SUBSTRATES FOR FLEXIBLE SENSORS

The fabrication of sensors for chronic wound monitoring re-quires substrates that are biocompatible and flexible to avoiddamaging the repairing tissue. Meanwhile, they should also besufficiently strong and robust. This section describes some ofthe most common materials that are currently used as substratesfor flexible sensors that can be adapted for wound management.Table I summarizes selected properties of some of the materialsdiscussed in the following subsections.

A. Silicone Elastomers

Silicone elastomers such as polydimethylsiloxane (PDMS)and EcoFlex are staple materials in the BioMEMS area andare the primary materials often associated with stretchable andflexible sensors [44]–[46]. PDMS is a transparent elastomerwith a Young’s modulus of about 100 MPa [47]; EcoFlex, incontrast, is only translucent but is much more flexible, with aYoung’s modulus of about 60 kPa [48]. Their silicone structurerenders them highly stable, both chemically and thermally (upto 250 ◦C).

Fig. 3. Various biocompatible substrates that have been used for fabricatingflexible sensors. These include: (a) PDMS [51], (b) Parylene-C [66], (c) poly-imide [71], (d) paper [131], (e) silk [96], and (f) Parylene-C/PDMS bilayer [60].

PDMS and EcoFlex are prepared by mixing a base pre-polymer and a crosslinker. The resulting mixture can then becast onto flat or patterned low-surface-energy substrates. Oncecrosslinked, sensors can be fabricated on the substrate beforereleasing it into a stand-alone flexible platform. They are excel-lent rapid prototyping materials for biomedical microsystem,microfluidic [45], [49], [50], and stretchable interconnects [seeFig. 3(a)] [51]. A major disadvantage of PDMS and EcoFlex istheir low adhesion to thin metallic films, creating difficulty infabricating electrical interconnects on them. However, one canovercome this to some extent by sandwiching the conductors inbetween two elastomeric layers.

B. Parylene

Parylene refers to a set of biocompatible, inert, and non-degradable thermoplastic polymers produced by chemical vapordeposition (CVD) of a dimer of p-xylylene [52]. While multipledimer variations (i.e., Parylene-C, N, D) are commercially avail-able, Parylene-C is the most commonly used due to its usefulcombination of electrical and physical properties including itsimpermeability to moisture and chemical inertness. Parylene-Chas been used for fabricating microchannels [53], [54], valves[55]–[57], and flexible sensors [22], [58]–[65].

Parylene deposition has the advantage of not requiring astrong vacuum or high temperature. It is typically depositedat pressures on the order of 0.1 torr, unlike other clean roomprocesses (e.g., metallization) that are conducted at much lowerpressures (below 10–5 torr). Although the solid dimer requires atemperature of about 150 ◦C to vaporize, once in the gas phase,the monomer enters a room-temperature deposition chamberwhere it is simultaneously adsorbed and polymerized on thesubstrate [52]. The thickness of a Parylene coat is controlledby the amount of dimer loaded into the Parylene chamber andits deposition time. The resulting Parylene coatings are com-pletely conformal, uniform in thickness, and pinhole-free [58][see Fig. 3(b)] [66].

C. Polyimide

Polyimides are a class of polymers with a stiff aromatic back-bone structure that results in high mechanical, chemical, and


TABLE IBIOCOMPATIBLE SUBSTRATES FOR FLEXIBLE SENSORS

thermal (up to 400 ◦C) stability. Such properties have madepolyimides ideal for use in the electronics industry as a dielectricmaterial in interconnects, and as a substrate for display appli-cations. More recently, they have been widely used as flexiblesubstrates in biomedical and chemical sensors [67]–[70].

Polyimide is available commercially (i.e., Kapton, DuPont)in sheet form, semi-cured sheets, and solutions suitable for spincoating. Many properties of the film such as its Young’s mod-ulus, thermal expansion coefficient, and dielectric constant canbe adjusted during the synthesis process (typically by the man-ufacturer), thus enabling the creation of polyimide with proper-ties that address specific needs. Recently, considerable attentionhas been focused on remote wound monitoring using wirelesspolyimide-embedded sensors that are incorporated into dress-ings. Other recent polymide-based sensors include flexible pHsensors [see Fig. 3(c)] [71], tactile sensors [72], ethanol gassensors [73], and piezoresistive sensor arrays [74].

1) Polyethylene Terephthalate: Polyethylene terephthalate(PET) is another commonly used flexible substrate that is com-mercially available in sheets (e.g., 3M PP2500). PET is a strongthermoplastic that is thermally stable up to about 250 ◦C. It isflexible enough to conform to typical body curvatures, but ad-equately rigid to process. PET is naturally hydrophobic, but itcan be made hydrophilic by exposure to plasma [75], [76]. Thisenables its use as a structural material that can be readily bondedto PDMS without additional adhesives [77]. All these proper-ties make PET an excellent candidate for fabricating flexiblesensors [78]–[80].

D. Cellulose (Paper)

Paper is an old technology that can be repurposed with mod-ern methods to create complex devices, including flexible sen-

sors for wound monitoring. Paper consists of a mesh of cellulosewith many unique properties including biocompatibility, lowcost, and ubiquity. In addition, it is easy to vary its thickness,fiber size, porosity, and hydrophilicity. Cellulose is naturally hy-drophilic and insoluble in water and most organic solvents dueto a strong hydrogen bond between polymer chains. Of particu-lar importance to sensor fabrication are filter paper, wax paper,and parchment paper [81]. The first is known to have excellentwicking properties. The latter two are naturally hydrophobic,but their surface properties can be altered by a plasma treatmentor by laser ablation [82].

Recent research has taken advantage of these properties forthe fabrication of paper-based systems for biomedical applica-tion such as controlled drug delivery, tissue engineering, sutures,biodegradable vascular grafts [83], [84], low-cost microfluidics[see Fig. 3(d)] [85], and various flexible sensors [86]–[92]. Fur-ther integration of paper with emerging fabrication technologieswill surely promote the development of more complex flexiblesystems on paper.

E. Other Materials

Many other natural and biodegradable materials have alsogarnered significant attention due to their additional favor-able properties. Materials such as polylactic acid (PLA) andpoly(lactic-co-glycolic acid) (PLGA) can be cast into rigidtransparent films for use as substrates with programmable long-term degradability. Other materials such as polyvinyl alcohol(PVA), starch [93], [94], and silk can serve as structural materi-als that dissolve in water [see Fig. 3(e)]. Silk, in particular, hasbeen used to fabricate devices, including highly conformal brain


TABLE IITECHNOLOGIES FOR PROCESSING SENSORS ON FLEXIBLE SUBSTRATES

(a)[117]; (b)[16], [104]; (c)[127]; (d)[81], 210]; (e)[226], [227]; (f)[211], [228]; RT = room temperature.

electrodes [95] as well as completely biodegradable electroniccircuits [see Fig. 3(f)] [96].

In addition to homogeneous materials, in order to enhance theperformance, some researchers have turned to polymer compos-ite substrates [97]–[102]. By combining two or more materials,the resulting composite may feature many additional propertiesor it may exhibit new physical properties that are not manifestedin the original single-polymer systems [103]. For instance, agraphite-PDMS composite results in a material with electricalproperties of graphite and elastic properties of PDMS (such ma-terial has been used to create soft pads for a flexible temperaturesensor [101]). Similarly, gold nanoparticles can be embedded inpolyurethane to create a composite film that is stretchable whileremaining electrically conductive [97]. The improved propertiesof such materials can be exploited for the fabrication of novelwound healing sensors.

IV. FABRICATION METHODS FOR FLEXIBLE SENSORS

Most flexible sensors are fabricated using techniques thatoriginate from traditional microfabrication and MEMS pro-cesses such as photolithography, thin film deposition, andwet/dry etch [104]. Although these techniques usually requireflat, rigid substrates, they can be adapted to flexible sensor fab-rication by mounting the flexible material on a temporary rigidsubstrate for processing. After fabrication, the rigid substrate isremoved, releasing a completed device on a flexible substrate.Soft lithography using PDMS molded against patterned hardsurfaces in particular has been a popular approach for fabricat-ing flexible microsystems [19], [104] such as three-dimensional(3-D) microfluidic networks [105], micropumps [106], andbiomimetic surfaces and structures [107], [108].

Alternative approaches make use of rapid prototyping andother un-conventional techniques for developing low-cost flex-

ible sensors on an extended set of substrate materials, oftenwithout the need for a temporary rigid backing and photolithog-raphy. These include laser engravers, 3-D printers, inkjet print-ers, and cutter plotters [46], [109], [110]. These approachesenable direct processing of many flexible materials, includingpaper [111], [112], acrylic [113]–[115], and various other com-mon polymers [116].

This section summarizes various methods for the fabricationof flexible sensors. The first subsection covers more traditionalmicrofabrication techniques, while the second subsection de-scribes emerging approaches based on un-conventional meth-ods. These techniques are summarized in Table II.

A. Microfabrication Techniques for Flexible Sensors

The specific structure of any flexible sensor depends on itsintended function, but generally it is composed of a sensingmodule (e.g., electrodes or active materials) or array of mod-ules that is mounted on or encapsulated in a flexible material.In cases where the flexible material is rigid enough for han-dling during fabrication (e.g., cm-scale 100 μm PET films), thematerial may be used directly as a substrate to perform pho-tolithography and other processes. Less rigid materials, how-ever, can be mounted/bonded on a rigid low-surface-energysubstrate (e.g., a silanized silicon wafer) allowing for easy re-lease at the end. Alternatively, the material can be depositedon a hard substrate, processed, and released (e.g., Parylene-C vacuum deposited on silicon/glass or polyimide spin-coatedonto a silanized wafer) [19], [104], [117]. In such cases, thefilms are coated with photoresist and patterned using a mask.Micro patterns can then be defined by oxygen plasma etchingthe polymer [118]. This technique has been used to create con-formal brain electrodes on Parylene [58] and can be similarlyused for fabricating flexible sensors for wound monitoring.


B. BioMEMS Techniques for Flexible Sensors

BioMEMS fabrication builds on traditional MEMS tech-niques with additional emphasis on pattern transfer onto elas-tomeric and biological substrates [104], [119]. A central tech-nique in the field is the casting and processing of PDMS sub-strates for replica micromolding. The technique is often usedto create reservoirs and channels [16]. With micromolding, amaster is first fabricated in a hard material (e.g., an etched sili-con wafer). The master is then used to make replicas on PDMSby pouring the prepolymer onto the silanized master and allow-ing it to crosslink. The PDMS is then peeled off to reveal thedesired pattern. The master can be reused many times to createadditional replicas, which can be bonded onto other flexible sub-strates to form closed channels and reservoirs. This techniquehas been used to form microchannels that are subsequently filledwith a room temperature liquid metal alloy to create stretch-able interconnects [51], [120]; alternate applications include thefabrication of flexible pressure sensors [121], [122] and actua-tors [123]–[126], all of which can be incorporated into woundmonitoring systems.

C. Emerging Rapid Prototyping Techniques

Modern rapid prototyping techniques have attracted signifi-cant interest in the past few years due to their ability to processa broader range of materials, including those that are heat sen-sitive or require dry processing conditions (e.g., PVDF, paper).These techniques have the additional advantage of being muchmore economical (at low production volume scale) than tradi-tional cleanroom-based methods while still offering adequateprinting resolution for many biomedical applications.

1) Cutter Plotters: A tabletop cutter-plotter such asCAMEO from Silhouette America (Orem, UT) is a low-costfabrication equipment capable of machining thin (up to 500 μm)polymer films as well as fabrics and thin metal foils, with reso-lutions of up to 100 μm [127]. The film is fed through the cutter-plotter and the machine drives a sharp cutting blade along thefilm, transferring patterns from a CAD drawing onto substrate.These systems operate at room temperature, enabling rapid ma-chining of low-critical-temperature materials such as PVDF.

2) Laser Engravers: Commercial laser engraving systemsoffer good resolution, control, and processing speed [seeFig. 4(a)]. These systems consist of a laser module connectedto a machining enclosure that contains a working stage and asoftware-controlled lens. The substrate is placed on the stageand the lens module guides the laser beam on the surface of thesubstrate to cut or ablate regions as defined in a CAD drawing.Most commercial systems use either a 10.6 μm CO2 laser (typ-ical powers of up to 150 W) suitable for cutting polymers andwood or a 1.06 μm fiber laser (typical powers of up to 40 W)that can mark metals and cut thin foils [128]–[130]. These sys-tems have a linear scanning speed of a few meters per secondand the output power and laser spot size/focus can be adjustedby software. The laser spot size can be as small as 25 μm forcurrent systems.

Laser systems do heat up the substrate at the focal pointand are not compatible with heat-sensitive materials such as

Fig. 4. Commercial rapid prototyping tools are increasingly used in biomedi-cal microdevices development, including (a) hydrophilic patterns on hydropho-bic paper created using laser surface treatment [82], and (b) 3-D-printed needlearrays [209]. The inset of (a) shows a commercial laser engraver [210] and thatof (b) depicts an example of a modern 3-D printing system [211].

PVDF; however, they are very attractive for rapid machining ofmany other emerging microsystem materials such as polymersand paper. Their high degree of control and ease of operationenables a variety of capabilities such as fabrication of shadowmasks for screen printing. With a laser system, such masks canbe readily machined out of a polyimide tape and then transferredonto the target flexible substrate (e.g., paper, PET) followed bydeposition of the ink via squeegee printing or spin coating.

Another laser-enabled fabrication method is the direct writ-ing of hydrophilic traces on a hydrophobic paper. This techniqueuses commercial or custom-fabricated hydrophobic paper (e.g.,wax paper or parchment paper) as the flexible substrate [131].Regions of the substrate that are ablated become hydrophilic,while the rest of the substrate remains hydrophobic. The result-ing traces can be used as hydrophilic microchannels or openreservoirs for paper-based microfluidics and chemical sensorsthat can be incorporated into wound monitoring systems.

3) Inkjet Printers: Another technology for flexible sensorfabrication is high resolution inkjet printing [132]. These sys-tems consist of a print head that moves relative to a flat stageand deposits droplets of liquid compounds or functional suspen-sions onto a substrate. The liquid “inks” are thus used to transferCAD drawings onto the substrate with a minimum droplet sizeof about 25 pL [80], [133] without the need for shadow masks.Inkjet printing is compatible with various substrates, both rigidand soft, as long as they are maintained flat on the stage. The ma-jor concern is possible inadequate adhesion of the ink onto thesubstrate. With some inks, such as metallic ones, a sintering stepis necessary to ensure trace conductivity and improve adhesion.In such cases, the possible substrates are limited to those that canwithstand sintering temperatures (typically about 150 ◦C). Thislimitation may be overcome by employing a localized sinteringtechnique (e.g., local laser heating).

4) 3-D Printers: 3-D printers have garnered significant at-tention in the past few years due to their increased resolution andaffordability [see Fig. 4(b)]. These systems enable direct manu-facturing of 3-D CAD drawings in a layer-by-layer fashion. Theprinter may either deposit a polymer resin that cures with timeand heat, or it may deposit layers of a polymer that is selectivelycross-linked with a laser. The structural materials used by the3-D printers are available in biocompatible formulations for usein biomedical applications. One major drawback is that printing


TABLE IIIPREFERRED REQUIREMENTS FOR WOUND MONITORING SENSORS

Dynamic range Normal values Resolution

pH (a) pH 4−8.6 pH 4−6 0.1 pH

Oxygen (b) 5−100 mmHg 30−50 mmHg 1 mmHg

Moisture (c) 0−100 %RH Varies by wound

dressing 1 %RH

Temperature (d) 25−41 ºC 29−35 ºC 0.1 ºC

(a)[139], [140], [229]; (b)[157], [230]; (c)measuring dressing moisture level [231], [232]; (d)[183]--[185].

is limited to a single material at a time, which does not permitthe manufacture of complex (e.g., conducting and nonconduct-ing) systems. Nevertheless, current 3-D printers have alreadyshown to be invaluable in manufacturing biomedical structuresfrom tissue scaffolds to bionic ears [134]–[138].

V. PHYSICAL AND CHEMICAL SENSORS FOR CHRONIC

WOUND MONITORING

This section describes the importance of monitoring variousphysical and chemical parameters in chronic wounds and listsrecent flexible sensors that are relevant to such applications.Various parameter values for wound healing sensors are sum-marized in Table III.

A. Flexible pH Sensors

Chronic wound pH is a key indicative parameter for identi-fying problematic regions. The primary reason is the elevatedpH caused by bacterial infections. Unlike healthy skin or heal-ing acute wounds that have a slightly acidic pH (5.5−6.5),chronic wounds often contain regions with pH values higherthan 7.4 due to the alkaline byproducts of bacterial colony pro-liferation [139]–[141]. To promptly identify and treat infectedregions, it is important to measure the pH at various locationsacross the wound. Commercial pH meters are adequate for eval-uating acute wounds; since healing in these wounds occurs rel-atively uniformly throughout the lesion. Chronic wounds, how-ever, present a challenge due to the large pH variations within thewound bed. Such cases require multiple measurements through-out the wound with adequate spatial resolution not achievable bycommercial probe. A more practical alternative would be a flex-ible array of miniaturized pH sensors that can cover the woundand generate a high-density map of pH levels, thus pinpointingthe precise location and concentration of infection. This is theapproach that recent research has adopted, i.e., electrochemicalpH sensors fabricated on a flexible substrate using conventionalmicrofabrication technology [142]–[144] as well as rapid pro-totyping techniques (e.g., drop casting, inkjet printing, screenprinting) [145]–[149].

Traditional electrochemical pH meters (i.e., glass probe) con-sist of a reference (e.g., Ag/AgCl in a stabilizing KCl electrolyte)electrode and a sensing electrode, typically platinum or anothernoble metal. The sensing electrode is enclosed in a glass tubefilled with an electrolyte. When the probe is placed in a so-lution, H+ ions create a potential across the glass, which can

be measured with the sensing electrode with respect to the ref-erence. The measured potential is proportional to the H+ ionconcentration. These pH meters require regular calibration andmust be stored in an electrolyte solution to prevent the probeelectrolyte from drying out. Additionally, the rigidity of suchsensors renders them impractical for incorporation into confor-mal structures for wound dressing applications.

More recent pH sensors rely on a solid-state design that re-places the glass probe with a conductive electrode that is coatedwith a pH-sensitive film, which can be patterned on a flexiblesubstrate. In most cases, the film is either a metal oxide (e.g.,IrOx , RuO2 , SnO2) or a conductive polymer (e.g., polyaniline,polypyrrole). Metal oxide films can be deposited by variousestablished processes, including sputtering, electrodeposition,and sol-gel [145]. The operation of these sensors is based onvalency changes in the oxygen atoms of the metal oxide causedby absorption of hydrogen ions form the test solution, whichresults in a generated potential relative to the reference elec-trode [150]. Such sensors can feature high sensitivities (e.g.,∼−77 mV/pH [151]), biocompatibility, fast response time (sec-onds), and thermal stability; however, they often suffer fromdrift [152] and require expensive fabrication materials. Poly-mer film sensors offer a similarly straightforward fabricationprocess, typically achieved by drop-casting or electropolymer-ization. Their operation is based on the protonation and de-protonation of nitrogen atoms in the polymer in response to thehydrogen ion concentration, which creates a potential across theelectrodes that is proportional to pH. These sensors generallyoffer a similar response time to metal oxide sensors, but theyexhibit higher sensitivities (∼−1300 mV/pH [153]), increasedstability, and lower cost. Both, metal oxide and conductivepolymer film architectures offer high flexibility and sensitivity,eliminate the need for frequent calibrations, and are compatiblewith the flexible substrates processing techniques discussed inSection IV.

Chemomechanical sensing methods have also been in-vestigated for pH transducers using pH-sensitive hydrogels[154]–[156]. A recent publication features one such pH sensorfabricated on a gold wire coated with a pH-sensitive hydrogel,which is then wrapped with a second wire and mechanicallystabilized with a PVA/PAA coat [156] [see Fig. 5(a)]. pH ismeasured via impedances across the two wires, which changesas the hydrogel swells/shrinks in response to pH values between6 and 9.1. An array of such thread-form pH sensors can be read-ily incorporated into wound dressings.

B. Flexible Oxygen Sensors

Suboptimal oxygenation is one of the major healing inhibitorsin chronic wounds [29], [42], [157], [158]. Unlike acute in-juries that receive sufficient oxygen via a working blood vesselnetwork, the irregular vasculature structure of chronic woundsis incapable of providing sufficient oxygen for tissue growth andremodeling. Modern treatments [159], [160] often expose largeareas of the body to unnecessarily elevated oxygen concentra-tions that can damage healthy tissue [4]. Emerging technolo-gies promise to more precisely regulate oxygen administration


Fig. 5. Examples of various flexible sensors for measuring (a) pH [156],(b) oxygen [85], (c) oxygen [167], (d) humidity [102], (e) temperature [101],and (f) mechanical stress [196].

via oxygen-trapping [161], [162] or oxygen-generating [163]wound dressings. Their oxygen control can be further improvedby the incorporation of flexible oxygen sensors into such sys-tems for closed-loop oxygen sensing/delivery.

A common type of flexible oxygen sensor is a Clark cell type.It consists of a gold or platinum cathode (sensing electrode), aAg/AgCl anode (reference electrode), and an electrolyte (e.g.,KOH). Such sensors are typically fabricated on an oxygen-permeable substrate, most commonly Teflon [164]–[166]. Whenthe cathode is maintained at 0.7 V with respect to the anode, oxy-gen diffusing through the membrane is reduced at the cathode,thus generating a measurable electric current that is proportionalto the concentration of oxygen. This type of sensors offers a reli-able measurement technique but requires a power source duringmeasurements.

An alternate approach is to use a galvanic oxygen sensor,which uses two dissimilar metals (i.e., a Zn or Pb anode and asilver or gold cathode) in an electrolyte enclosed by an oxygen-permeable membrane. In the presence of oxygen, the anode isoxidized, producing electrons that reduce the oxygen at the cath-ode. The generated current is directly proportional to the oxygenlevel, and no power source is required while taking measure-ments. A recent example developed in our group consists ofa silver cathode and a zinc anode patterned on a parchmentpaper substrate by screen printing [85]. An electrolyte and aPDMS membrane seal the two electrodes in a single chamber[see Fig. 5(b)]. The paper behaves as the gas-permeable mem-brane, allowing oxygen to be reduced at the cathode. For oxygenconcentrations between 0% and 45%, this sensor produces anelectric current that changes linearly from 2 to 129 μA (sensitiv-ity of 2.6 μA/%O2) with a time response of 17 s. The dynamicrange and response time are acceptable for assessing hypoxia inchronic wounds, and the structure can be readily embedded intotraditional wound dressings.

Another recent type of oxygen sensor is based on ZnOnanowires [167] and uses the piezotronic effect for sensing.Metal connections are made to a single ZnO nanowire on aPET substrate, creating a metal–semiconductor–metal structure[see Fig. 5(c)]. Local variations in oxygen concentration alter

the carrier density in the ZnO, thus changing the current for agiven bias. This sensor has the unique advantage of having atunable sensitivity that can be adjusted via its piezo charges.Additionally, its nanoscale dimensions enable the fabrication ofsensors much smaller than other flexible oxygen sensors; hence,an array of these would allow high-resolution measurements inthe wound without compromising flexibility.

C. Flexible Moisture Sensors

Clinical studies have shown the need for a delicate balanceof moisture for optimum wound healing, but wound moistureassessment remains largely qualitative [168]–[171]. As men-tioned in Section II, moisture control is typically achieved bycovering the wound with a dressing that is appropriate forthe moisture status of the wound and replacing it as needed.Although commercial dressings (e.g., sponges, gas-permeablefilms, hydrocolloid patches) offer various levels of humiditycontrol [28], [43], they lack a means for quantitatively assessingmoisture levels. This need can be addressed by the developmentof flexible sensors that can be incorporated into commercialwound dressings. Most flexible moisture sensors measure elec-trical impedance [172]–[180] or capacitance [102], [181], [182]between two adjacent (typically interdigitated) electrodes ona humidity-sensitive film (e.g., polyimide, benzocyclobutene,conductive oxides). When the film absorbs moisture from theenvironment, its electrical properties (i.e., resistance or permit-tivity) are altered, thus changing the impedance or capacitanceacross the electrodes. Recent examples include an impedance-based humidity sensor on a polyimide substrate with its humiditysensitive film fabricated by screen printing [177], as well as a ca-pacitive sensor that uses in situ polymerization of a nanoscaledpolypyrrole (PPy) film on a cellulose substrate [102] to form aconducting humidity sensitive composite [see Fig. 5(d)].

D. Flexible Temperature Sensors

Although skin temperature is often neglected during chronicwound assessment, recent studies suggest that differential tem-perature measurements (i.e., between two symmetrically locatedbody regions) may provide useful information regarding thecondition of a wound [183]–[185]. In particular, since a clas-sic symptom of infection is an increase in temperature, it maybe beneficial to incorporate temperature sensors with routinewound assessment to detect possible emerging infections. Aswith any other wound dressing, a sensor that is flexible wouldbe most beneficial for such applications.

Most flexible temperature sensors consist of a resistive ele-ment (e.g., Pt) mounted on a flexible substrate such as poly-imide [186]–[189]. The measured resistance is directly propor-tional to the temperature of the resistive element and dependson the temperature coefficient of resistance (TCR) of its mate-rial composition. Such sensors are stable, reliable, and linear.Alternative approaches incorporate composite materials to im-part additional functionality to the sensor [101], [190]. Onerecent example is the use of a PDMS–graphite composite tocreate an array of soft resistive elements [101] [see Fig. 5(e)].Its elastomeric nature improves contact with sensing surfaces.


Additionally, the TCR of the composite is larger than that oftraditional Pt elements, allowing for increased temperature sen-sitivity. With a dynamic range of 30−110 ◦C, such sensors maybe applicable for wound monitoring applications.

E. Flexible Mechanical Sensors

Applied mechanical forces are known to affect the rate andquality of wound healing. Animal studies have shown that me-chanical tensile stress on a wound site promotes acceleratedneovascularization and cellular proliferation in addition to amechanically stronger and more organized tissue [191]–[193].Too much stress, however, can be detrimental, as it can tearhealing tissue. Aside from qualitative reports, the effect of ap-plied stress/strain remains difficult to quantify due to a lack ofstandard assessment techniques. One way to address this issueis to incorporate flexible mechanical sensors into wound dress-ings to periodically (or continuously) and uniformly measuremechanical forces in the wound.

Some of the most readily adaptable sensors for woundstress/strain assessment are those that have been developed fortactile sensing [72], [194]–[199]. Such sensors feature soft, de-flectable membranes that translate strain into changes in resis-tance [195], [196] or capacitance [197] [see Fig. 5(f)]. These areparticularly suited for wound applications since their soft com-position reduces the risk of sensor-induced mechanical insultsto the healing tissue. Additionally, their overall flexible natureallows integration with commercial products.

VI. CONCLUSION

This review summarized the current and emerging materi-als and processes that can be used to fabricate flexible physi-cal/chemical sensors for use in chronic wound monitoring. Thehealing process in the wound bed depends on many parameters,including oxygen level, pH, moisture, temperature, and mechan-ical forces. Some of these can also serve as indicators of woundhealing status. Ideally these conditions would be monitored bya sensor-loaded wound dressing to increase measurement fre-quency while reducing the need for medical professional at-tention and the overall therapeutic expenditures. In order todevelop clinically relevant sensor platforms, it is important tounderstand the various sensing modalities as well as materi-als and processes that are available for fabricating appropriatedevices. The numerous publications cited in this review offersolutions for creating sensors that are flexible and inexpensive,as well as readily adaptable for incorporation into commercialwound dressings.

A common theme among many recent publications is theuse of rapid prototyping technologies for fabrication on uncon-ventional materials. This trend is an important one that offersa fresh alternative to the more expensive MEMS and batch-scale microfabrication methods. The emerging new materialsand processing techniques with their focus on flexibility andshort design cycle is paving the way for health monitoring flex-ible electronics that can integrate seamlessly into many bodilyregions. Such techniques, however, are more suited for smallvolume production in academic labs and niche applications

and require scale-up transition for large volume applicationssuch as wound dressing. The modular nature of laser systems,inkjet printers, laminators, and screen-printers allows them tobe integrated into an assembly line configuration for roll-to-rollmanufacturing [200]–[202]. Such mass production setups maypresent their own set of challenges, but current investigationswill help to facilitate this transition.

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Manuel Ochoa (S’08–GSM’09) received the B.S.degree in electrical engineering from the CaliforniaInstitute of Technology, Pasadena, CA, USA, in 2009,and the M.S. degree in electrical and computer engi-neering from Purdue University, West Lafayette, IN,USA, in 2012. He is currently working toword thePh.D. degree in electrical and computer engineeringat Purdue University.

Since 2009, he has been a Research Assistantwith the Ziaie Biomedical Microdevices Laboratory,Purdue University. His research interests include the

integration of common materials (e.g. paper, tape, yeast) for the developmentlow-cost multifunctional microsystem platforms for biomedical applications,including transdermal drug delivery and dermal wound healing.

Rahim Rahimi received the B.S. degree in electri-cal engineering from the Iran University of Scienceand Technology, Tehran, Iran, in 2008. He is cur-rently working toward the Ph.D. degree in electricaland computer engineering at Purdue University, WestLafayette, IN, USA.

He has been a Research Assistant with the Zi-aie Biomedical Microdevices Laboratory, PurdueUniversity, since 2012. His research interests in-clude drug delivery, microfluidics, implantable de-vices, flexible electronics, electrochemical sensors,

biomedical sensors, and biomedical applications of microelectromechanicalsystems.

Babak Ziaie (M’95–SM’07) received the Ph.D. de-gree in electrical engineering from the University ofMichigan, Ann Arbor, MI, USA, in 1994.

From 1995 to 1999, he was a Postdoctoral Fel-low and an Assistant Research Scientist at the Centerfor Integrated Microsystems, University of Michi-gan. From 1999 to 2004, he was an Assistant Pro-fessor with the Electrical and Computer EngineeringDepartment, University of Minnesota, Minneapolis,MN, USA. Since January 2005, he has been with theSchool of Electrical and Computer Engineering, Pur-

due University, West Lafayette, IN, USA, where he is currently a Professor.His research interests include biomedical applications of microelectromechan-ical systems and microsystems, including implantable wireless microsystems,smart polymers for physiological sensing and control, micromachined interfacesfor the central nervous system, and biomimetic sensors and actuators.