The literature shows scarce efforts made for the development

The literature shows scarce efforts made for the development of diagnostic tools and devices with potential for effective monitoring of chronic wounds. Matzeu et al. [14] demonstrated an RFID-based skin temperature monitoring system obtaining a 0.2°C temperature measurement accuracy. A flexible platinum-based miniaturized temperature sensor is proposed by Moser and Martin [15] to operate within 0-400°C range. However, the sensor has not been used in wound dressings. McColl et al. [16] have developed an impedance sensor-based moisture monitoring system for wound dressings. Based on this principle, Ohmedics© has developed a commercial moisture monitoring device called WoundSense® [17]. However, the sensing system is not designed to stay within the dressing for continuous moisture measurement. To date, no device exists for continuous monitoring of sub-bandage pressure, though some sub-bandage pressure meters, such as Kikuhime®, are in practice, but they have very limited clinical application. Miniaturized pressure sensors have been designed and used for other applications such as intracranial pressure (Codman® [18] and Mejzlik [19]), intraocular pressure (Ning et al. [20]), spinal plates pressure (Sauser et al. [21]) and for general in-vivo applications (Clausen et al. [22], Willyan et al. [23] and Hill et al. [24]). For a wound monitoring system, calibration and characterization of sensors is important in order to sense accurate information beneath the dressing. Incorrect sensor measurements will negatively impact on clinical decisions about the wound under observation. In our recent review article [25], we have highlighted the gap in potential use of modern sensors and wireless technology for wound monitoring applications. In this paper, we present methods for calibration and characterization of temperature, moisture, and pressure sensors deemed suitable to be interfaced with a wireless telemetric sensing system. The sensors were carefully selected considering their small size, low power operation, flexibility and minimal invasiveness to the wounds. The calibrated sensors were interfaced with the developed prototype sensing system, and the performance of the system was tested by placing the system and sensors under a purchase 12-O-tetradecanoyl phorbol-13-acetate bandage on a mannequin leg.
Methods and materials
Experiments using a compression bandage For real-time validation of all the calibrated sensors working together, we designed a prototype flexible wireless transmitter (9.7cm×4.7cm) and a receiver (4cm×4cm), both operating at the radio frequency (RF) of 2.4GHz using the IEEE 802.15.4 ZigBee protocol. The sensors were interfaced with the transmitter circuit using additional interface circuits designed for impedance and voltage conditioning. Both the transmitter and the receiver use Atmel’s Atmega128RFA1 integrated circuit as transceiver, with all other required components for their proper operation. The temperature, moisture, and pressure sensors were mounted on the mannequin leg at proper positions, and then an elastic compression bandage (Coban™ 2) was wrapped over them (Fig. 8(a) and (b)). Distilled water was sprayed externally over the surface of the bandage. The data acquired by the sensors in real-time was transmitted to the nearby receiver at an interval of 5s. The receiver was attached to the computer’s serial port. On reception of the full data packet, the receiver first saved and then sent the captured data to the PC serial port (Fig. 8(c)). A graphical user interface (GUI) was designed in Visual Basic to display the real-time information on temperature, moisture, and pressure, along with capture-time values (Fig. 9).
Analysis and discussion The real-time performance of the selected sensors was tested extensively and monitored with the designed prototype RF transceiver system. The transmitted and received data was validated using commercial temperature, moisture, and pressure meters. The average errors in temperature, moisture and pressure measurements in the system were ±1.5°C, ±3 %RH and ±2mmHg, respectively. The measurement resolutions obtained for temperature, moisture, and sub-bandage pressure were 0.15°C, 0.85–5 %RH, and 0.05–0.56mmHg respectively using a 10-bit analog to digital converter (ADC) within the RF transceiver. The temperature resolution is more than sufficient to detect temperature variations in human skin, which range from 32°C to 37°C under normal conditions [14]. The moisture and pressure measurement resolutions are not uniform owing to nonlinear characteristics of respective sensors. In actual scenario, the temperature sensor will be placed on periwound skin and not directly on top of wound. The moisture sensor could be placed inside a foam dressing (e.g. Mepilex®, Allevyn™) over the wound, while the pressure sensor could be placed near ankle to monitor sub-bandage pressure. The measurements with the temperature and the moisture sensors were location independent, while the sub-bandage pressure measurements showed dependency on the sensor location on the mannequin limb, possibly due to changes in applied pressure on to uneven surface morphology around the limb. However, the optimum position of the pressure sensor can arguably be determined during clinical experiments.