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Although injuries, traumatic events, and surgical emergencies have been unlikely during the current space missions, their occurrence must be considered for future human space explorations. During long-lasting Space-missions, wound treatment and monitoring could become a fundamental problem, demanding more research in this area (Barratt and Pool, 2008; Cialdai et al., 2020). For example, first-degree burns can occur as a result of ultraviolet (UV) light exposure through unfiltered spacecraft windows (Barratt and Pool, 2008). Space represents a very special remote environment, and solutions developed for Space applications could inspire the creation of a better healthcare system in remote areas on-ground.

When a wound occurs, the first step is to assess its severity, which will form the basis for the following treatment. Clinical assessment of the wound is still the most common and cost-efficient method to assess wound severity. This method relies on a subjective evaluation of the wounds’ external features, such as size of the wound, wound edges, site of wound, wound bed (colour, amount of granulation tissue), presence of necrotic tissue, wound’s depth, level of exudate, and pain caused by the wound (Grey et al., 2006). The advantages of this method are that it is a rapid method and does not require specialized equipment. However, it is subjective. For instance, a nurse and clinician may have different evaluation and assessment criteria depending on their prior experiences (Nagle et al., 2020). Seeing that wound severity is usually diagnosed by a specialist, a problem appears when wound injury happens in places that lack wound care specialists, a situation common in remote areas. One way to face these problems is to develop diagnostic tools and utilities that are cost-efficient, easy-to-use, and that could support medical workers in wound diagnosis and evaluation of the healing by monitoring one or more biomarkers.

Among many clinical biomarkers, special attention has been given to pH (Shukla et al., 2007; Schreml et al., 2011; Jones et al., 2015; Power et al., 2017), oxygen (Schreml et al., 2010), and exudate composition, with the focus on matrix metalloproteinase analytes (MMPs) (Muller et al., 2008; Liu et al., 2009; Power et al., 2017). Also, temperature (Dini et al., 2015; Chanmugam et al., 2017; Jaspers et al., 2017; Power et al., 2017; Cwajda-Białasik et al., 2020) and moisture (Bishop et al., 2003; Milne et al., 2016) are the physical parameters that are frequently associated with the wound healing process. In particular, the temperature is considered as an informative parameter for all types of wounds, and it can provide information about infections of wounds, even before clinical signs appear (Power et al., 2017).

A common way to measure wound temperature is by means of Infrared (IR) thermography (Ring and Ammer, 2012). Thanks to the development of portable high-resolution affordable thermal cameras, interest in their use for wound assessment has increased. This is mainly due to their non-invasive and easily interpretable results obtained in a very short period. Various studies (Dini et al., 2015; Chanmugam et al., 2017; Jaspers et al., 2017; Martínez-Jiménez et al., 2018; Cwajda-Białasik et al., 2020; Ganon et al., 2020) have used the temperature difference between wound and healthy skin (ΔT), measured by using IR thermography, to characterize the wound status. An increase of ΔT can be related to hyperaemia, inflammation, or infection in venous leg ulcers (Dini et al., 2015; Cwajda-Białasik et al., 2020) or pressure ulcers and surgical wounds (Chanmugam et al., 2017). In burns, the temperature is often correlated to the burn depth and the burn healing time (Jaspers et al., 2017; Martínez-Jiménez et al., 2018; Ganon et al., 2020). Instead of just measuring the temperature locally, these studies demonstrated the importance of measuring the temperature distribution on the wound site and the surrounding healthy skin. Despite these promising results, IR thermography is not yet applied in daily clinical practice for wound assessment, mainly because standard guidelines and protocols have not been established yet (Shterenshis, 2017).

The field of flexible and wearable bioelectronics, capable of monitoring physiological information and assisting in proper treatment is growing exponentially (Rogers et al., 2018; Chen et al., 2020). Researchers and engineers are working on the development of new technologies for smart point-of-care systems (Zhang et al., 2016; Coppola et al., 2021). Therefore, the development of a point-of-care device for wound monitoring is more feasible. Such a device could reduce hospitalization times, the suffering of the patients, and costs (Mehmood et al., 2014; Derakhshandeh et al., 2018). Moreover, it could provide the ability to face emergency surgery, acute trauma, burns, and wounds, in remote and closed environments, such as Space. At present, wearable technology in wound care is limited to laboratory testing, and commercial wearable point-of-care systems are not broadly available. The reason could be found in the complexity of the wound healing process, the broad variety of wounds’ types, and the limited understanding of relevant wound biomarkers.

Ideally, a smart sensor for wound monitoring should have specific properties, such as 1) wearability/ability to adapt to the body shape, 2) biocompatibility, 3) high sensitivity, 4) easy-to-use, and 5) no external power. Due to these requirements, sensors based on the colourimetric approach appear as an ideal solution (Ajay et al., 2017; Isapour and Lattuada, 2018). In that respect, liquid crystals (LCs) have emerged as a promising technology for wound management sensing. The most interesting property of LCs is their structural colouration that can be manipulated by changing the external parameters, such as temperature. This makes them ideal candidates for the development of easy-to-use, label-free, and passive sensors, where the output signal is a change in colour detectable by the naked eye. LCs have a long history of being used as responsive materials in different technologies, thanks to their unique properties. Numerous studies have demonstrated the possibility to produce rapid diagnostic optical sensors for temperature (Gao et al., 2014), pH (long Chen et al., 2018), humidity (Saha et al., 2012; Zhao et al., 2019), gas (Esteves et al., 2020) and molecules (ang Niu et al., 2017; Zhang et al., 2018) based on LCs. The most widespread sensors are those for temperature, owing to the diversity and availability of thermotropic LCs.

LCs are a unique state of matter between crystalline solid and isotropic liquid. In thermotropic LCs, phase transitions from crystalline solid to smectic, cholesteric and, finally, isotropic liquid, are caused by temperature changes, and they are mainly composed of rod-like molecules. In cholesteric LCs, also known as chiral LCs, molecules are inherently chiral, and the average molecular orientation is twisted with a certain periodicity, leading to a helical structure. This structure is characterised by a helical pitch that refers to the distance over which the LC molecules undergo a whole 360° twist. The size of this pitch determines the wavelength of the reflected light. The pitch of a cholesteric LC can be of the order of magnitude that corresponds to the wavelength spectra of visible light, allowing structural colouration to occur. An increase in the temperature results in a decrease in the pitch, which causes a shift in the wavelength of the reflected light. This presents the basis of the sensing principle of cholesteric LCs (Mitov, 2012). Thermotropic LCs, intended to be used as colourimetric temperature sensors, are characterized by three parameters: the lower clearing point temperature, optical activation range, and the higher clearing point temperature (Abdullah et al., 2010). The lower clearing point temperature is the temperature at which LCs first reflect colour in the visible spectrum (red). The optical activation range is the temperature range at which thermotropic LCs actively reflect visible light. When the thermotropic LCs pass through their optical activation range, they reflect visible light from longer wavelengths (red) to shorter wavelengths (blue) as temperature increases until their higher clearing point temperatures are reached. The higher clearing point (further in the text referred as clearing point) temperature is the temperature at which thermotropic LCs stop to reflect colours in the visible spectrum. Beyond the clearing point temperature, thermotropic LCs are transparent again (Abdullah et al., 2010).

A study from 2017 (LeSar et al., 2017) investigated the possibility to use commercial thermotropic LC–coated fabric for the early detection of high-risk foot complications. Using direct visual analysis, they demonstrated that the fabric could accurately map temperatures on the surface of a hand, which supported the hypothesis that this approach can be used to develop a temperature-sensitive system to monitor complications high-risk foot. They used two different types of LC fabrics to expand the active range. The most recent prospective study (Hodorowicz-Zaniewska et al., 2020) used the Brast Tester–LC foil (Braster SA, Ozarów Mazowiecki, Poland) for the early detection of breast cancer. Despite the good performance, this contact LC thermography did not find its way into commercial use. Reasons for this include that the protocol still requires a medical specialist for its implementation, it lacks standardization, and it has been commonly replaced by contactless infrared thermography. Gao et al. (2014) were the first ones to propose a skin-like system that consists of thermotropic LCs patterned into large-scale, pixelated arrays on thin elastomeric substrates, demonstrating that such a system could be used as an epidermal temperature sensor.

This work compares the sensing ability of LCs with respect to IR thermography. To this aim, temperature mappings of surfaces of different topographies obtained using LCs sensing elements are compared against the ones obtained using IR thermography. The IR thermography is chosen as a reference since it is the most widespread technique used in wound temperature studies (Dini et al., 2015; Chanmugam et al., 2017; Jaspers et al., 2017; Martínez-Jiménez et al., 2018; Cwajda-Białasik et al., 2020; Ganon et al., 2020). This comparison will bring new insights into the possibility of using LCs-based systems for wound temperature monitoring. Some challenges in their application in clinical practice will be highlighted. Although, the final aim of the research here presented is temperature wound analysis, this paper is focused on preliminary construction and analysis of the measuring potential of the envisaged device.

As described in the introduction, IR thermography is commonly used to measure temperature distributions in wound studies. IR thermal cameras are sensitive to environmental conditions. Before imaging, IR cameras require input values including room temperature, humidity, distance from the object and its emissivity. It is not always possible to correctly determine these parameters, although they are directly influencing the reading output. The emissivity of human skin is considered to be between 0.97–0.99 (Boylan et al., 1992; Keenan et al., 2017), while different studies have found that the emissivity of wounds can be greater by 0.01–0.03. This could result in an underestimation of the ΔT value by 0.1–0.2°C (Boylan et al., 1992), and could prove significant in clinical evaluations of some studies. Depending on the location and size of the wound, the temperature of the healthy surrounding skin can vary for several degrees (Carriere et al., 2020). Hence directly influencing the ΔT value and the reading outcome. Moreover, the wound bandage should be removed for each temperature measurement using IR thermal cameras. This is a disadvantage because it prevents continuous temperature monitoring.

In general, this study showed that LCs have the potential to fabricate advanced temperature distribution sensors. This is not common to other sensing elements, such as resistance or impedance-based ones, that could measure only point-wise values. Another important advantage of this configuration is its simplicity in reading the output. When compared with infrared imaging, LCs are capable of reproducing the main features of the temperature fields surrounding a wound mimicking shape, even geometrically complex. However, the recognition of the patterns was somewhat less sharp for the LCs with respect to the IR camera. This might be caused by the contact between the patch and the to-be-measured surface.

The thermotropic LCs used in this research can detect even small temperature variations (high sensitivity) and have a good repeatability as well as a fast response time (less than 0.5 s). However, LCs used here have a narrow bandwidth, resulting in high sensitivity but a small effective temperature range. The temperature range can be extended by changing the type of thermotropic LCs or by changing the component ratio. According to (Ochoa et al., 2014), the preferred requirements for temperature wound monitoring sensors depend on the wound’s type. However, during the wound healing process, the dynamic temperature range is 25–41°C. The other formulations (see Figure 4), showed that it is possible to cover from 29–44°C. In this study only one of them was used in order to focus on the principle and show the feasibility of such an approach, but the same study can be performed on the other liquid crystals.

The ageing of LCs reduces their shelf life. Proper preparation and storage are necessary to minimize their ageing. It was suggested that pure thermotropic LCs should not be exposed to fats, greases, organic solvents and dust. They can also be susceptible to UV and IR radiation, and exposure to these sources reduces the shelf life (Abdullah et al., 2010). Another observation is that the hue–temperature dependence is non-linear. The decrease in sensitivity is particularly notable when entering in the blue region, where small changes in temperature do not cause significant changes in hue. The cause of hysteresis in cholesteric LC is complex. Previously it was suggested that hysteresis is strongly associated with chemical composition (Dixon and Scala, 1970) and that hysteresis depends on the sample thickness (Zink and Belyakov, 1997). In the case of a narrowband LC, the thicker sample showed no hysteresis, and the thinner sample showed up to about a 0.5°C bias toward lower temperatures when cooled (Zink and Belyakov, 1997). Moreover, the magnitude of the shift increases with increasing maximum temperature before cooling. For this particular application, hysteresis should be avoided since it can give false information about temperature at the wound during cooling. Hysteresis can be reduced significantly if the operating temperature is regulated. Most importantly, the operating and storage temperature should be kept below the clearing point temperature (Zink and Belyakov, 1997). However, this does not necessarily represent a problem if the calibration is done with small steps (0.1°C) and a computerised way of reading colour is included, such as the application used in this work.

Note that the technological readiness of LCs, although promising, does not directly translate into the possibility to include them in clinical trials. Several issues should be considered in that respect and they will be discussed below.

Wound contamination: the current status of the proposed patch requires that it should be in direct contact with the wound. In clinical practice, medical personnel are obliged to use wound dressings to avoid contamination, protect wounds, and promote wound healing. Nowadays, an ideal wound should fulfil the following characteristics: 1) creating a moist and warm environment around the wound, 2) allowing gas exchange, 3) protecting the wound from bacterial infections, 4) creating a mechanical protection, 5) controlling the exudate level, 6) being biocompatible, non-toxic, non-allergenic, 7) being easily removable, i.e., the material should be non-adhering to avoid removing of a newly formed tissue, 8) being able to stimulate healing, and 9) being costly acceptable (Sood et al., 2014). Applying the proposed LC system on top of the wound dressing could influence some of the aforementioned dressing’s characteristics, such as gas exchange, exudate and moist control properties. Moreover, the LCs sensing characteristics could be affected by the micro-environment of the wound and surroundings. Another option would be to use the LC-based sensor just during the bandage removal. In this case, continuous monitoring is not possible.

Size of the wounds: in literature it is possible to find an outstanding number of wound sensors, either based on LCs or not. Everyone who participated to a real clinical study on wound imaging knows that serious wounds, and especially burns, could cover large portions of the body, sometimes even the whole body. It is evident that the development of a contact sensor that would monitor such a wound is currently not possible.

Shape and 3D nature of the wounded tissues distributions: although maybe technically possible, it is very difficult to imagine a contact sensor when the shape and distribution of the wounds is not flat or even worse, changing “volumetrically” during the healing time. It would be possible to think of an intermediate layer hosting the sensing element. If this solution could potentially solve the flatness problem, it will also challenge the capability of this layer to transmit reliably biophysical signals from the wound to the sensing element.

Although it is beyond the scope of this paper to provide a complete overview of the clinical requirements for wound diagnosis and follow-up, it is possible to draw preliminary conclusions. LCs could represent a complementary element to support the medical decision and follow-up in the case of wounds, especially burns. However, the application of such kinds of sensors in clinical practice is far from being straightforward, and the plethora of sensors’ concepts presented in the literature are still at a low-TRL stage.