**1. Introduction**

Lower limb compression garments are regularly employed in clinical settings [1] and, more recently, by athletes seeking to improve performance or decrease recovery time [2,3]. In medical settings, compression devices including passive compression socks and active intermittent pneumatic compression devices, are used in the treatment of hypertrophic scarring, venous ulcers, deep vein thrombosis and other diseases [1]. Some examples of physiological testing in [4,5] have shown a significant increase in the mean arterial pressure, cardiac output and cardiac stroke volume, with a significant reduction in the heart rate through applying active compression cyclically at frequencies between 1 to 2 Hz and at pressures of approximately 10 mmHg in the knee region. A major challenge for any compression device, regardless of application, is ensuring the consistency and repeatability of the applied pressure throughout its use [6]. The limited sizing options of most passive compression gear can vary the pressure experienced between subjects due to body shape differences. Further, the consistency of pressure applied by the garmen<sup>t</sup> on the same subject can be affected by the donning method [7]. A further challenge is that the shape, size and tissue compliance of the leg can change during compression application due to fluid redistribution, muscle contraction and movement of the garment/device.

Partsch et al. [8] made recommendations on the characteristics that an ideal sensor would need in order to measure the interface pressure between a compression garmen<sup>t</sup> and the lower leg skin. Specifically, they identify the need for sensors that are low cost, low hysteresis, have minimal creep, high sampling rate, high accuracy, flexibility, durability and small sensor surface area. To date, there has ye<sup>t</sup> to be a sensor developed that meets all these requirements.

Burke et al. [9] and Ferguson-Pell et al. [10] examined the use of Tekscan FlexiForce and Interlink Electronics FSR piezoresistive-based sensors for measuring interface pressure in compression garments. Piezoresistive sensors function by measuring the change in resistance under a mechanical load. They are thin and can provide reliable results during dynamic measurements. However, these sensors su ffer from significant hysteresis and drift, and thus are not dependable for long time-scale measurements [8]. Additionally, piezoresistive sensors are dramatically a ffected by base curvatures under 32 mm in radius [10] and require relatively large pressures for reliable measurements [11]. The total error exhibited in the sensors, accounting for errors in repeatability, hysteresis and linearity, is approximately +/− 10 mmHg for a pressure range 0 to 96 mmHg, [9], which limits the utility of these devices for pressure mapping of compression garments. Note, for the remainder of the document, unless otherwise noted, errors are described for the pressure range of interest between 0 and 100 mmHg.

Similar to piezoresistive sensors, strain gauges also respond to a change in resistance with mechanical strain [12,13]. Kraemer et al. [14] and Ghosh et al. [15] utilized strain gauges to measure the compression applied by compression garments/bandages on a subject's thigh and a mannequin leg, respectively. However, since strain gauges exhibit a better sensitivity along their largest dimension, these sensors tend to become thick for measuring interfacial pressures. Additionally, strain gauges tend to exhibit thermal and humidity dependence, and high hysteresis [12].

Pneumatic-based sensors are one of the most prevalent designs for measuring compression garmen<sup>t</sup> interfacial pressure [9]. These sensors utilize bladders or compartments that are filled with a small volume of air connected to a pressure transducer through a flexible hose [16]. As pressure is applied, the volume of the bladder decreases, thus increasing the internal pressure. At a constant temperature, it is assumed that the internal pressure measured by the transducer is equivalent to the external applied pressure [16]. Commercially available pneumatic sensors, such as the Salzmann MST MKIV [17], MediGroup Kikuhime [18] and Microlab PicoPress [19] are thin and flexible devices that can be used to measure interfacial pressures between the body and compression device. Furthermore, pneumatic pressure transducers have been found to exhibit repeatable results within +/− 3 mmHg of error [8,19]. A major shortfall of this technology is its temperature dependence [8]. Additionally, their high sensitivity to curvature can result in overestimates of pressure by up to 150% [9]. The sampling rate of the system is a ffected by the tubing length, which can be important in the evaluation of rapid inflation active compression systems. Additionally, pressure sensing with an array of pneumatic sensors presents multiple inconveniences. The sensors each require a dedicated pressure transducer and have a large surface area (a PicoPress sensor is approximately 50 mm in diameter), resulting in poor spatial resolution. These types of sensors need a grea<sup>t</sup> amount of tubing and rely on pressure transducers that are bulky and almost 10 times larger in footprint than the actual sensor, [19]. Sensor multiplexing can reduce the number of transducers in a system, which uses solenoid valves to isolate each bladder for pressure reading (see Figure 1). However, the addition of a manifold increases the system volume, lowering the sampling rate and the transducer's sensitivity to bladder volume change.

Capacitive sensors have also been employed in other textile applications, [20–22], such as muscle activity detection, [20], and pressure mapping for insoles, [21]. These sensors typically consist of a textile used as a dielectric material with conductive yarn woven as electrodes. They have the benefit of being small (good spatial resolution), flexible, and easily used in arrays for pressure distribution measurements [23,24]. Compared to other sensor technologies, they also possess relatively low temperature sensitivities [23,24]. Q. Guo et al. [25] developed a tactile floating electrode capacitive sensor which displayed good linearity up to 350 kPa and improved durability over parallel electrode designs. Capacitive measurements also present their own challenges, often requiring complex circuitry to filter measurement noise [23]. Signal hysteresis can also a ffect performance, as the dielectric material is often viscoelastic and exhibits nonlinear behaviour when stressed [26]. Additionally, the current-pressure relation is nonlinear [27] which makes it more challenging for proper calibration. Nano-fibers have been employed in the development of pressure sensing technologies for other applications [27,28]. However, due to the high temperature sensitivity of materials used, such as gold, they are not suitable for the application of interest.

**Figure 1.** Multiplexing example of four pneumatic sensors.

In this paper, a novel, capacitive-based, dielectric electroactive polymer (DEAP) pressure sensor is developed in collaboration with StretchSense Ltd. for in-situ pressure measurements under textile garments. Specific testing conditions are required when evaluating pressure sensors for compression garments, since the interface is neither flat nor solid. Thus, multiple testing methods are presented in Section 3, with results in Section 4. The sensor's pressure sensitivity, error and durability are characterized. The performance effects due to temperature, humidity, boundary conditions and hysteresis are also examined. This is the first sensor to use DEAP materials measuring low pressure ranges (0 to 100mmHg) for portable and wearable applications, including the fast sampling rate required for many physiological applications. In addition to the sensor fabrication and characterization, the small, portable wireless circuitry for the proposed sensor is one of the novel aspects of this work. The wireless sensing unit communicates data over Bluetooth with the actuation system.
