1. Introduction

Optoelectronic devices such as solar cells, photodetectors, lasers, and light-emitting diodes (LEDs) have been widely used in many aspects of modern life such as telecommunication, energy, consumer electronics, and solid-state lighting over the past few decades. Recently, there is an ever-growing interest to extend the applications of optoelectronics to biological and biomedical research to offer new opportunities for biosensing, high-precision therapies, personized medicine, and human-machine interfaces, etc. [1–4]. Conventional high-performance optoelectronic devices rely on single-crystalline inorganic semiconductors, such as silicon or gallium arsenide, built on brittle and rigid substrates [5–7]. Those devices exhibit significant structure and property mismatches with soft and curvilinear biological tissues and organs. Therefore, innovations in new optoelectronic materials, design strategies, fabrication, and integration techniques are critical to achieve flexible and even stretchable optoelectronic systems for advanced wearable and implantable applications.

Organic semiconductors provide compelling alternatives for bio-integration compared to inorganic semiconductors. They offer a wide range of highly attractive features for bio-related applications, including mechanical flexibility, tailorable optoelectronic properties, low temperature solution processing, lightweight, and good biocompatibility [8,9]. Importantly, the performance of organic electronic devices such as organic thin-film transistors (OTFTs), organic LEDs (OLEDs), organic solar cells (OSCs), and organic photodetectors (OPDs) have been improved dramatically since the pioneering work of conjugated polymers by Heeger, MacDiarmid, and Shirakawa in 1977 and small molecule based OLEDs by Tang and Van Slyke in 1987 [10–12]. For example, power conversion efficiencies (PCEs) of OSCs have been increased from less than 5% in 2005 to over 17% in 2018 while OLEDs have been commercialized by Samsung and LG for use in flexible displays [13,14].

Herein, this review provides a detailed overview of recent progress in organic optoelectronic materials and devices for biointerfacing. We begin with a brief overview of the organic semiconductor materials. We then discuss strategies to design stretchable biointerfaces using organic materials. We provide examples of bio-integrated organic optoelectronic devices in different applications. A concluding section provides a summary with challenges that remain to be addressed and future research directions.

2. Materials, device structures, and design strategies

2.1 Materials and device structures

Organic semiconductor materials in general can be separated into two classes: conjugated polymers and small molecules. Polymers have high molecular weights that result in a slow aggregation speed during solvent volatilization, allowing fine-tuning of the morphology of organic photoactive layers with different solvents and solvent additives. Polymers are advantageous for large-area device fabrications. Small molecules have advantages in good batch-to-batch repeatability, high crystallinity, and excellent charge carrier mobilities due to strong intermolecular π-π stacking. Both polymers and small molecules are used to construct OSCs, OPDs, and OLEDs for bio-integration.

Figure 1 summarizes the most commonly adopted architectures of OSCs, OPDs and OLEDs, including a transparent anode, a hole transport layer (HTL), a photoactive layer, an electron transport layer (ETL), and a low work function metal cathode. Despite its brittle nature and the extensive ongoing research efforts in developing flexible transparent electrode materials [15,16], indium tin oxide (ITO) is still the major transparent electrode material reported in flexible bio-integrated organic optoelectronic devices to date. For polymer based OSCs/OPDs, the photoactive layer generally consists of a p-type conjugated polymer such as poly(3-hexylthiophene) (P3HT), and an n-type fullerene derivative such as [6,6]-phenyl-C₆₁-butyric acid methyl ester (PCBM) in a bulk heterojunction (BHJ) structure. For small molecule based OSCs/OPDs, a stacked bilayer structure with individual p-type layer such as pentacene, and n-type layer such as tris(8-hydroxyquinoline)aluminum is adopted. Depending on the applications of interest, the bandgaps of the organic semiconductors can be adjusted to tune the absorption spectra of the resulting devices. For example, the polymer thieno[3,4-b]thiophene/benzodithiophene (PTB7) with a bandgap at 1.8 eV is used to detect photons from 600 nm to 700 nm wavelength and the polymer polyindacenodithiophene-pyridyl [2,1,3] thiadiazole-cyclopentadithiophene with a narrower bandgap at 1.5 eV exhibits further near infrared (NIR) responsivity [17–19]. For OLEDs, the photoactive layer is an emissive layer consisting of fluorescent emitters for blue OLEDs and phosphorescent emitters for green, orange, and red OLEDs. Poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS), an optically transparent conductive polymer with superior electrical conductivity that can go up to 4600 S/cm [20], is widely used as the charge transport layers in OSCs, OPDs and OLEDs.

 

Fig. 1. Schematic representations of the structures of an OSC/OPD (a) and a bottom emitting OLED (b).

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2.2 Design strategies for stretchable organic optoelectronic devices

Stretchable devices that can preserve designed optoelectronic properties under repeated tensile and compression strain are very attractive for bio-integrated applications involving motions. Below we discuss two common strategies to impart elasticity to organic optoelectronic devices.

2.2.1 Buckling strategy

The buckling strategy is a method used to build multilayer stacked devices on pre-stretched elastomer substrates followed by the release of the strain after device processing. The resulting devices can retain the initial performance when stretched up to the pre-stretched value. The buckling strategy was first applied to design stretchable OSCs in 2011 by spin coating a transparent PEDOT:PSS anode, an active layer consisting of P3HT and PCBM, and an eutectic gallium-indium based liquid metal cathode directly on a 20% pre-strained poly(dimethylsiloxane) (PDMS) substrate [21]. After releasing the pre-strain, the resulting film shows buckles perpendicular to the pre-strain direction. The photovoltaic characteristics of the OSCs are nearly independent of the tensile strain up to 27%. The first stretchable OLED was demonstrated in 2013 with a polymer based emissive layer and PEDOT:PSS anode [22]. The device is first fabricated on an ultrathin polyethylene terephthalate (PET) foil on glass substrate followed by transferring to a 100% pre-strained elastomer tape. Upon relaxing the elastomer, the OLED forms a random network of folds to accommodate the compression strain. The resulting device maintains an electrical performance comparable to its flat state. Sun et al. exploited a laser-programmable buckling process to realize highly stretchable OLEDs with high efficiency and mechanical robustness from small molecule based OLEDs [23]. The phosphorescent green OLEDs are first fabricated via thermal evaporation on a polymer/Si substrate. Then the OLED/polymer film is adhered to the top surface of a 120% pre-stretched elastomeric substrate with one-dimensional long-period grating grooves prepared by femtosecond laser ablation. After releasing the strain, ordered buckling features are only formed for OLED areas above the grooves. The stretchability depends on the width of the lines and grooves. The OLED displays a luminous efficiency of 70 cd/A under 70% strain, which is almost the same to that from the unstrained device (72.5 cd/A). Small fluctuations in performance are observed with tensile strain up to 100% over 15,000 cycles. The same group later fabricated two-dimensional stretchable OLEDs using a similar buckling strategy where the OLEDs are transferred to biaxially pre-stretched elastic substrates [24]. The resulting devices exhibit stable efficiency under 100 stretch-release cycles between 0 and 25% strain.

2.2.2 Intrinsically stretchable organic photoactive layer

Another straightforward approach to develop stretchable organic optoelectronic devices is to directly develop elastomeric components for every layer of the devices, including substrates, conductors, charge transport layers, and organic photoactive layer. Several detailed and excellent reviews regarding elastomer substrates and conductors have appeared in the past and the readers who are interested in those topics are encouraged to read these reviews [25,26]. Here we highlight recent progress in designing stretchable organic photoactive layer.

The mechanical and optoelectronic properties of organic semiconductor materials can be engineered by changing the conjugated backbone and side chains to control crystallinity, intermolecular interactions, solubility, molecular planarity, etc. In addition to molecular structures, the photoactive layer in OSCs, OPDs, and OLEDs is usually a heterojunction blend of two materials. As a result, the morphology of the blend has a significant influence on the elastic properties and device performance. Lipomi et al. studied the stretchablitity of OSCs based on two different conjugated polymers mixed with PCBM [27]. One polymer is P3HT and another is a donor-acceptor polymer with diketo pyrrolo-pyrrole moiety, thiophene, thienothiophene, and thiophene repeating units (DPPT-TT). P3HT is stiffened more than DPPT-TT when blended with PCBM. The crack formation in P3HT:PCBM film is much worse compared to that in DPPT-TT:PCBM film under the same 20% strain. Savagatrup et al. later found that the length of the alkyl side chains has a direct influence on the mechanical properties of poly(3-alkylthiophene) (P3AT) [28,29]. Longer side chains reduce the tensile modulus of both the pure P3AT film and P3AT:PCBM blend film due to the decreased crystallinity in the polymer films. The OSC device based on poly(3-dodecylthiophene):PCBM shows a stable PCE under a mechanical strain of 10% with no visible cracks in the film while the PCE of a P3HT:PCBM device decreases to 0 with visible cracks in the film under the same condition. Pei and co-workers performed a detailed study on the effect of small molecule additives on the elastic deformability of PTB7:PCBM blend [30]. The solvent additive creates free volume in the active layer that allows the nanocrystalline grains to slide in a reversible way and accommodate large external deformation. As a result, the morphology of the OSC device exhibits rubbery elasticity at room temperature and the device exhibits a moderate PCE even under 100% strain. Liang et al. reported the first intrinsically stretchable OLEDs with a luminescent layer based on an intrinsically stretchable white light-emitting polymer and 1,3-bis-[(4-tert-butylphenyl)-1,3,4-oxidiazolyl]phenylene blend [31]. Semi-transparent graphene oxide/Ag nanowire/polyurethane acrylate composite is used as both cathode and anode, PEDOT:PSS serves as the HTL, and polyethylenimine (PEI) works as the ETL. The brightness of the OLEDs decreases after 100 cycles between 0 and 40% strain while the current efficiency is improved, likely due to changes in thickness or morphology of the emissive layer after stretching. The resulting OLEDs could be stretched up to 130% with uniform light distribution across a 3 × 4 cm² device area.

3. Biological and biomedical applications

In this section, we discuss recent progress in applying organic optoelectronic devices to biointerfacing, including the development of bioelectronic power sources, epidermal pulse oximeters, optogenetic light sources, optical nongenetic cell stimulators, wearable displays, etc.

3.1 Bioelectronic power sources

Bioelectronic devices require sustainable power sources for stable operation. One possible solution to address this need is to combine sensors/stimulators with energy harvesting blocks. Solar cells is a promising renewable energy harvester for this purpose. Indeed, inorganic solar cells based on silicon and gallium arsenide have been successfully demonstrated as wireless power supplies for various biomedical devices such as LEDs [32,33], pacemakers [34], and temperature sensors [35].

OSCs are slowly gaining momentum as flexible and lightweight power supply alternatives. Arias and co-workers demonstrated a wearable energy bracelet by integrating a four-cell OSC module (Fig. 2(a)) with a stretchable and flexible silver-zinc battery where the OSC module directly charges the connected battery [36]. The photoactive layer consists of poly[N-9'-heptadecanyl-2,7-carbazole-alt-5,5-(4’,7'-di-2-thienyl-2’,1’,3'-benzothiadiazole)]:PCBM blend. The OSC shows a maximum power output at 5 mW under sunlight. The system shows a high energy conversion and storage efficiency even under compact fluorescent lighting and is especially suitable for wearable applications with high indoor illuminance (Fig. 2(b)). A recent report highlights a waterproof stretchable elastomer-coated OSC with an air-stable active layer including a blend of quaterthiophene and naphtho[1,2-c:5,6-c’]bis[1,2,5]thiadiazole and PCBM [37]. The device is washable by sandwiching the cell between two elastomer substrates and exhibits a PCE of 7.9% under a one-sun illumination (Fig. 2(c)). The PCE of the devices remains at 80% of the initial value even after 52% mechanical compression for 20 cycles with 100 min of water exposure, as shown in Fig. 2(d). The same group later reported self-powered ultra-flexible optoelectronic devices with OSC power sources and organic electrochemical transistor (OECT) sensors on ultrathin Parylene substrates (Fig. 2(e)) [38]. The photoactive layer and zinc oxide nanoparticle ETL of the OSC are patterned with one-dimensional grating structures to improve PCE and mechanical stretchability. This design results in a maximum PCE of 10.49% and a power-per-weight of 11.46 W g^-1 for the flexible OSCs. The PCE decreases to 7.33% after 900 repetitive compression cycles with 33% compression strain. When integrated with a flexible OECT into a conformal self-powered cardiac sensor, the sensor can measure electrocardiogram (ECG) signals of a rat heart with a high signal-to-noise ratio of 40.02 dB (Fig. 2(f)). In addition to self-powered sensors, OSCs can be integrated with neural stimulation electrodes to develop self-powered stimulators that can promote neurite outgrowth and manipulate the polarity of neuron-like cells [39]. Lee et al. designed a light-sensitive artificial synapse that uses a self-powered OSC based on P3HT:PCBM to generate voltage pulses under patterned optical illumination and produce a sufficient output to stimulate an organic synapse [40].

 

Fig. 2. OSC based power supplies for bioelectronics. (a) Structure of a four-cell OSC module. (b) Current-voltage characteristics of the module in (a) under different indoor lighting conditions. Reproduced with permission from [36]. Copyright 2017, The Authors. (c) Optical image of the washing process with the OSCs conformed to a dress shirt. (d) Normalized PCE as a function of the dipping time. Reproduced with permission from [37]. Copyright 2017, Springer Nature. (e) Schematic demonstration of an OECT integrated with a double-grating-patterned OSC. (f) Optical image of the self-powered integrated optoelectronic device attached to a rat heart for ECG measurement. Reproduced with permission from [38]. Copyright 2018, Springer Nature.

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3.2 Epidermal pulse oximeters

Oxygen selectively binds to hemoglobin and results in a mixture of oxyhemoglobin (HbO₂) and deoxyhemoglobin (Hb) in the blood. Pulse oximeters are optoelectronic systems utilized to measure the concentration of HbO₂ from pulsatile arterial blood and monitor the subject’s cardiovascular health. Epidermal pulse oximeters are ubiquitous because they measure the peripheral oxygenation (SpO₂) noninvasively. Conventional pulse oximeters are clipped on a fingertip with two quickly alternating LEDs operating at different wavelengths and a photodetector to record light signals. HbO₂ and Hb show different molar absorptivities with wavelength. As a result, the blood volume and concentration of HbO₂ determine the intensity of emitted light from the two LEDs transmitted through or reflected from the skin. In a transmission mode the photodetector is located at the opposite side of the skin to collect the transmitted light, while in a reflective mode the photodetector is placed in the same site as the LEDs to collect the reflected light.

OLEDs and OPDs are potential ideal candidates for epidermal pulse oximeters because they are low temperature solution processable and compatible with fabrication on ultrathin flexible substrates. Figure 3(a) shows a transmission mode all-organic pulse oximeter composed of solution processed red and green OLEDs and a printed OPD [41]. The active layer of the OPD is a BHJ blend of PTB7 and PCBM while the active layers of two OLEDs are fluorescent polymers with high stability and efficiency. Unlike commercially available pulse oximeters that use red and infrared wavelengths, green and red OLEDs are used due to the low efficiency and instability of solution-processable NIR OLEDs at that time. The emission peaks of the red and green OLEDs appear at 626 nm and 532 nm while PTB7 is sensitive to photons at both wavelengths. This pulse oximeter successfully measures the absolute pulse rate and arterial oxygen saturation with errors of 1% and 2%, respectively (Fig. 3(b)). More recently, the same group reported reflective mode pulse oximeters with blade coated red and green OLEDs and a Si photodetector placed on the same side of the wrist [42]. Blade coating technique allows simultaneous printing of two different emissive materials, which is beneficial in reducing the volume of ink used and improving the reproducibility of devices. Bansal et al. demonstrated an alternative reflective mode pulse oximeter with one OLED placed in between two OPDs, as shown in Fig. 3(c) [18]. The OLED is made of the conjugated polymer OC₁C₁₀-PPV with emission peak in the red and tail in the NIR regions. The two OPDs are based on PTB7:PCBM with optical filters to select different wavelengths at 610 nm and 700 nm. Figure 3(d) illustrates that the device can successfully detect changes in the concentration of HbO₂ upon induction and termination of ischemia in the arm when the OLED and OPDs are placed onto a subject’s forearm with 20 mm spacing. Lee et al. demonstrated a reflective patch-type monolithically integrated pulse oximeter based on OLEDs and OPDs with ultralow power consumption [43]. Both OLEDs and OPDs are composed of small molecule active layers on PET substrates. The device operates under an average low electrical power at 24 µW, which is far less than the inorganic counterparts.

 

Fig. 3. Organic pulse oximeters. (a) A transmission mode pulse oximeter with red and green OLEDs and one OPD placed on subject’s finger. (b) Pulsating photoplethysmogram signals recorded from the pulse oximeter in (a). Reproduced with permission from [41]. Copyright 2014, Springer Nature. (c) Schematic of a pulse oximeter consisted of an OLED and two OPDs. (d) Optical responses at two wavelengths from forearm skeletal muscle of a subject during ischemia and subsequent recovery. Reproduced with permission from [18]. Copyright 2014 The Authors, published by, Wiley-VCH. (e) Photos of a flexible reflectance oximeter array. Reproduced with permission from [44]. Copyright 2018 the Author(s). Published by PNAS. (f) Smart organic optoelectronic e-skin comprising pulse oximeters and displays. (g) Output signals from a pulse oximeter with 90% of oxygenation of blood. Reproduced with permission from [46]. Copyright 2016, The Authors.

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In addition to single point measurement, Arias group reported a flexible reflectance oximeter array (ROA) containing four red and four NIR OLEDs and eight OPDs for two-dimensional mapping of oxygenation of an area (Fig. 3(e)) [44]. The OLED and OPD arrays are fabricated separately using blade coating and screen printing on flexible polyethylene naphthalate substrates and later assembled together. The use of printing techniques makes the ROA comfortable to wear and efficient at high-quality signal extraction. The ROA sensor exhibits a 1.1% mean error when laminated on the forehead to measure oxygen saturation and creates two-dimensional oxygenation maps of forearms under pressure-cuff-induced ischemia.

Organic optoelectronic devices suffer from their instability in air due to reactions of the organic semiconductor materials with oxygen and moisture [45]. To address this challenge, Yokota et al. developed a high quality flexible and ultrathin encapsulation strategy combining repeated passivation layers of inorganic (SiON) and organic (Parylene) layers [46]. The encapsulation layers show a low water vapor transmission rate of 5.0 × 10⁴ g/m² per day and an oxygen delivery rate of 0.1 cc/m² per day. This thin geometry allows the OLEDs to operate stably after 1000 cycles with 60% stretching or under compression from 0 to 67% and the OPDs to sustain functioning after 300 stretching cycles with no degradation. The pulsating photoplethysmogram signals measured from the pulse oximeter remain constant after four days. The pulse oximeter is further connected to a skin mounted OLED display where the recorded signals could be displayed in real time, which demonstrates the possibility of implementing artificial electronic skins consisting of a pulse oximeter and a wearable display, as shown in Figs. 3(f) and 3(g).

3.3 Optogenetic light sources

Optogenetics is a seminal technology that allows fast optical control of targeted specific subtypes of cells or neurons in biological systems [47]. Microscale inorganic LEDs are widely used as minimally invasive implantable light sources for optogenetic research [48–51]. OLEDs offer several major advantages over inorganic LEDs for optogenetics. For example, the Young’s modulus of organic semiconductor materials in OLEDs is ususally much smaller than that of the inorganic semiconductors in inorganic LEDs, which helps to minimize foreign body response under chronic conditions. Tuning the emission properties of OLEDs could be made easily via chemical synthesis. More importantly, OLEDs can be structured to high density arrays with sub-cellular length scales (e.g. 10 µm) for optogenetic modulation of single cells/neurons by micro- and nano-fabrications. Gather group first demonstrated that OLED microarrays borrowed from microdisplay industry can be used for lens-free control of local motion of live cells [52,53]. Figure 4 describes the experimental set up. The OLED array has an active area of about 20 mm² and 230,000 pixels, each defined by a 6 × 9 µm² sized aluminum anode. Alternating layers of Al₂O₃ and an organic Barix polymer encapsulate the devices and enable close contact with live cells. Recently, the same group reported millimeter-scale high-brightness OLEDs with sufficient illumination intensity (0.3 mW/mm²) to achieve optical activation of channelrhodopsin-2 (ChR2) in vivo with low operating voltages at 5 V [54]. Blue, green, and orange OLEDs are fabricated via thermal deposition. The ChR2 expressed motor neurons show strongest behavioral response to blue OLED illumination while orange OLED illumination doesn’t trigger any cell activity due to the spectral mismatch.

 

Fig. 4. Schematic of an OLED microarray with cells adhered on top. Reproduced with permission from [53]. Copyright 2016, The Authors.

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3.4 Optical nongenetic cell stimulators

Despite the impressive progress achieved with optogenetics, it requires genetic modifications, making it challenging to serve as a treatment in humans. Nongenetic approaches to optically stimulate neurons or cells are highly desired for in vivo applications in humans, such as central or peripheral nervous therapeutics and implants for retinal prostheses [55]. Silicon based devices were extensively studied for neuronal stimulation applications [56,57]. However, these devices are rigid and not ideal for biointerfacing. Moreover, silicon is an indirect bandgap semiconductor, requiring a thickness of tens to hundreds of micrometers to efficiently harvest photons, and therefore limits the overall device thickness.

Meanwhile, organic semiconductor materials have shown the ability to photostimulate neurons. P3HT:PCBM blend, when integrated into an electrolytic photocapacitor with ITO anode, a reference electrode, and an ionic electrolyte, could transduce light pulses into localized displacement currents to electrically stimulate the neuronal cells cultured at the polymer/electrolyte interface, as shown in Fig. 5(a) [58]. The photostimulation follows a desired pure capacitive mechanism without inducing cell degradation and electrode damage. The capacitive charging of the polymer/electrolyte interface requires the presence of the back electrode such as ITO to collect photogenerated electrons in the photoactive materials (Fig. 5(b)) [59,60]. Band engineering by including an intermediate layer between the photoactive layer and back electrode enables controlling of capacitive and Faradaic processes [61]. Moreover, an OSC photosensor based on a small molecule anilino squaraine donor (2,4-bis[4-(N,N-diisobutylamino)-2,6-dihydroxyphenyl] squaraine) and PCBM blend is used to investigate the activation mechanisms of ion channels of single neuroblastoma cells [62]. Although photothermal effects are negligible at low illumination intensities [60], upon high intensity pulsed illumination, a slow photothermal nonelectrical activation of the potassium ion channels is observed. The electrolytic photocapacitor could be built on top of microelectrode arrays to enable simultaneous photoexcitation and recording [63]. Besides BHJ donor/acceptor blends, conjugated polymers alone can be directly used to stimulate cell activity. Antognazza and co-workers reported that P3HT film could activate ionic channels via the release of thermal energy from excited P3HT and the variation of the local ionic concentration mediated by photoexcitation [64]. However, not all conjugated polymers are suitable for nongenetic optical stimulation applications and their electrochemical stability, biocompatibility, and cell photostimulation efficacy vary a lot and should be evaluated case by case [65]. For example, the high bandgap polyfluorene derivative polymer poly[9,9-dioctylfluorenyl-2,7-diyl] showed a poor biocompatibility and failed to establish functional coupling with cells.

 

Fig. 5. Schemes of the photosensing interfaces based on (a) polymer bulk heterojunction layer. Reproduced with permission from [58]. Copyright 2011, Springer Nature. (b) small molecule bilayers. Reproduced with permission from [60]. Copyright © 2019 The Authors, some rights reserved; exclusive licensee American Association for the Advancement of Science.

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3.5 Wearable displays

Wearable displays that can act as communication channels between sensors and end users have attracted increased attention. OLEDs are suitable for wearable display applications because they are flexible and ultrathin. Kim and co-workers demonstrated a wearable health monitoring display that could visualize ECG signals on skin in real time [66]. The ECG sensor consists of an ultrathin stretchable electrode and a carbon nanotube amplifier (Fig. 6(a)). The OLED display is composed of ITO and Al electrodes, thermally evaporated blue and red small molecule emitting layers, an exciton blocking layer (EBL) for color tuning, and both Parylene-C substrate and encapsulation layer. By controlling the material and thickness of the EBL, the emitted colors of the wearable OLEDs exhibit an dependence on the recorded ECG signals and change from dark red, to pale red, to white, to sky blue, and finally to deep blue, as shown in Fig. 6(b). The ultrathin nature of the integrated device allows the display to conform to the curvilinear surface of human skin with high resilience against extreme bending and folding conditions.

 

Fig. 6. Wearable OLED displays. (a) Schematic illustration of a wearable cardiac-monitoring system. (b) Optical images showing color tuned emission of a wearable OLED display according to the measured real time ECG signal shapes. Reproduced with permission from [66]. Copyright 2017, American Chemical Society. (c) Demonstration of fiber OLEDs in clothes and textiles. Reproduced with permission from [68]. Copyright 2018, American Chemical Society.

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In addition to planar substrate-based OLED displays, OLED displays on fiber-type or planar textile substrates are attractive because they allow display functions to be incorporated into hierarchically woven clothes. Kim et al. demonstrated a fabric-type OLED fabricated on polyester-fiber-based woven fabric substrates with a thickness of 100 µm [67]. A 20-40 µm thick polyurethane layer serves as the planarization layer for the OLED devices. Multilayer barrier layers composed of Al₂O₃ and poly(vinyl alcohol) are inserted on the top and bottom sides of the OLEDs to protect the devices from moisture and oxygen. The small molecule based OLEDs are fabricated by vacuum deposition processes. With the multilayer barrier films, the fabric-type OLED devices exhibit performance comparable to those fabricated on glass substrates, and particularly, the OLED devices have lifetimes over 1000 h. Choi and co-workers reported weavable and highly efficient fiber-based OLEDs based on a simple low-temperature solution process, where the fiber OLEDs also show similar efficiency and lifetime to the conventional glass-based counterparts [68]. The structure of the OLED is PET fiber/PEDOT:PSS cathode/ZnO and PEI HTL/Super Yellow emissive layer/MoO₃ and Al anode/Al₂O₃ encapsulation. All layers except for the metal anode and encapsulation layers are fabricated via dip coating in a source solution while MoO₃/Al are prepared by thermal deposition and Al₂O₃ is deposited via atomic layer deposition. The OLEDs could withstand tensile strain up to 4.3% at a radius of 3.5 mm and are verified to be weavable into knitted clothes and textiles (Fig. 6(c)).

3.6 Other biomedical applications

In addition to applications in pulse oximeters, OPDs have been very recently applied to optical imaging of fluorescent calcium signals from neural cells [69]. The OPD consists a simple structure of ITO/PEDOT:PSS/active layer/Al. The active layer contains an extra-large bandgap polymer:PCBM blend that shows a high absorption over the wavelength range of 320-670 nm. The OPD exhibits an ultralow noise current density and linear behavior at sub-µA/cm² illumination intensities. Importantly, the response time and cut-off frequency of the OPD is faster than the fluorescence rising time of genetically encoded calcium indicators and higher than the fastest neural firing in the brain, which are crucial for optical imaging. Millimeter-scale blue OLEDs have been integrated with calcium imaging microscopes to measure the responses of neurons expressed with adeno-associated virus vectors [70]. Neuronal responses are recorded when the optical power density from the blue OLEDs reaches 60 µW/mm² at various frequencies.

There have been several reports that LEDs could be used in light therapy to improve the wound healing process [71,72]. OLEDs are advantageous towards this application because they can be integrated into wearable and portable formats. Indeed, it is shown that wearable red emitting OLED photobiomodulation patches could effectively stimulate fibroblast proliferation (over 58% of control) and enhance migration of fibroblast (over 46% control) under various illumination conditions [73].

4. Summary

Recent progress in organic semiconductor materials and optoelectronic devices have made it possible to fabricate flexible and stretchable organic optoelectronic biointerfaces for multiple wearable and implantable applications. This review covers various types of organic optoelectronic materials, devices, integrated systems, and highlights the current status of this exciting emerging field. Meanwhile, several scientific and engineering challenges need to be addressed before we can fully exploit the benefits of organic optoelectronic biointerfaces in practical biological applications. The unsatisfying long-term environmental stability of the organic optoelectronic devices must be improved to not affect their chronic functioning when in direct contact with air after laminated onto the skin or biofluids after implanted in the tissue. This requires developing innovative organic optoelectronic materials that are stable when exposed to air and biofluids for sensing/stimulation purposes and encapsulation strategies to protect the other parts of the devices. As a result, we will need a better understanding of molecular design rules to develop new organic semiconductor materials that exhibit superior optoelectronic and mechanical properties and good biocompatibility. Currently, there is a lack of mechanistic models to fully understand the organic device-biological interface. Other important future research directions include integration with advanced electronic modules for wireless communication or functional components such as microelectromechanical subsystems and microfluidic channels for multifunctional sensing/stimulation. Synergistic efforts from chemists, engineers, device physicists, and biologists are crucial to making further progress in this highly interdisciplinary area. In summary, organic optoelectronics based biointerfaces offer a fertile playground for more exciting basic science discoveries and clinical applications to be made.