DEFINICIONES CONCEPTUALES

Piezoresistive Force Sensitive Resistors (FSR) are widely used in ergonomics studies (Jensen et al., 1991b; Williams et al., 2012b,b) to measure magnitude of applied forces by different parts of the human body. Low-cost per unit, simplicity in cal- ibration, ease of use and their availability in different shapes and sizes and close to linear behaviour in lower force ranges are some of the advantages of the FSR sensors. FSR sensors’ measurement range are highly affected by their circuitry in- terfacing method. Figure 5.6 shows the schematic suggested by the manufacturer for interfacing these sensors.

Figure 5.6: FSR sensors’ suggested circuitry schematic

The electrical resistance (RF SR) of the conductive material which is used

in these type of sensors decreases by increasing the applied forces on the sensor’s active area. This resistance is an infinite value in absence of forces and decays to

zero when the maximum measurable load is applied which varies between models (eg. 100 Newtons). A two-wire communication is suggested by the manufacturer: one to supply a reference voltage (Vref) (either 5 or 3.3 Volts) and, one to read the

output voltage (Vout). A measuring resistor (RM) is also used in the sensor’s output

circuit to limit the current and to enhance the sensitivity range. The output voltage is computed from equation 5.1.

Vout=Vref ×

RM RF SR+RM

(5.1)

The voltage-to-force conversion equation (rarely provided), uses this value to estimate the force. However, since the manufacturer’s equation is highly influenced by laboratory test constraints, a more accurate calibration method is proposed to estimate the applied force values from the collected samples.

A digital force gauge measurement instrument (Sauter F H −500) with a dynamic measurement range of 0 to 500 Newtons was used to map the sensors’ digital output during sampling process to an accurate numerical force value (Sauter GmbH, 2015). The device was mounted on a test stand with an adjustable lever to change the applied forced on the sensors detection area by mechanical displacement of the measurement instrument. Figure 5.7 shows the calibration interfaces which are used in this phase.

Figure 5.7: Calibration Interface -Sauter F H−500 for actual force value, ParsGlove and DigiScale for digital force value

Round-shaped FSR sensors (see figure 5.9) were used in three variety of sizes (small, medium, and large) with respect to the mounting position on the human hand. These locations were identified with guidance of the medical experts in R&D stage of our previous experiment. Five sensors were randomly chosen from small and medium size categories for the calibration phase to monitor the sensors’ behaviour by changing the applied forces on their detection area. The large sensors where not included in calibration stage since they were only used to indicate the presence or absence of the force exertion on palmar surface of the human hand. Medical students are continuously advised to avoid leaning on the patient’s body with this area during palpation procedure to avoid any discomfort. Figure 5.8 illustrates the thenar and hypothenar eminences on the palmar surface of the human hand.

The sensors’ output voltages (0 to 5 Volts) were sampled into digital numer- ical values (0−1023 as one byte of data) by on-board Analogue-to-Digital (ADC block in figure 5.6) converting modules. The actual force magnitudes (in New- tons) were simultaneously recorded from the force sensors by a calibration test tool (Sauter GmbH, 2015) in each force application step. The sensors’ digital output was increased by 50 arbitrary units in each calibration step until small changes on the digital output shows significant force readings (550 units for small and 750 units for medium sensors).

Figure 5.9: IEE round-shaped FSR sensors; small (6mm), medium (12mm), large (24.8mm)

Despite the noted advantages of the FSR sensors their force measurement reliability is highly dependant on application time. Two undesired behaviours are well described when time is considered in a calibration experiment (Florez et al., 2010). The first phenomenon is known asCreepwhich is caused by the reduction in sensors’ electrical resistance after long term application of static forces (approxim- ately 2 Newtons higher than the actual force value after 10 minutes). The second is

Hysteresis reported as the loading and unloading curves (voltage-time plot) were not overlapped. The use of an epoxy resin dome on the sensors’ active area is ad- dressed to increase its pressure sensitivity. However, the force measurement range was significantly different when a dome is used particularly when sensors are moun- ted on the human hand. This may also reduce the human hands’ pressure sensitivity and flexibility of the sensors’ active area. The loading curve was monitored in this study when participants were asked to reach a given target force, hence, the occur- rence of hysteresis was not considered. Also, the total force application duration

was identified as 10 seconds; this helps avoid the behaviour of creep.

The force was loaded for a short period of time (approximately 5 seconds based on our previously captured data from medical experiment) to mimic a hand press-release action (see figure 5.10) and was released from the sensor’s detection area soon after each calibration step to avoid appearance of creep and hysteresis effects.

Figure 5.10: Press-release actions during abdominal palpation examination

Finally, the sensors’ calibration data were plotted (see figure 5.11) to il- lustrate the best force estimation (red asterisks) in each step and the potential variations among identical sensors (whiskers). Figure 5.11 shows the calibration outcomes for each category of sensors.

A 5th degree polynomial equation (see equation 5.2) was calculated from the calibration data by curve fitting technique to accurately estimate the actual forces

F(V) applied by the medical users from the sensors’ output voltages V that are digitally sampled by the ADC module (0 to 1023).

F(V) =c5V5+c4V4+c3V3+c2V2+c1V +c0 (5.2)

5.4

Summary

A brief overview of the ParsMoCap interface, which was designed and prototyped with help of the medical collaborators, and used as a crucial part of this quantifica- tion research, is proposed in this chapter. In addition, a robust method for calibra-

(a) small size sensors’ calibration results

(b) medium size sensors’ calibration results

Figure 5.11: Calibration results for five randomly tested force sensors per type

tion of force sensors has been proposed to provide a guide for future researchers to validate data reliability. The results in the calibration stage revealed the extreme variations (more than 10 Newtons) between identical sensors when heavy loads are applied on the sensors’ surfaces. The reliability of these sensors could also drop when a load is applied on them for long period of time. However, as stated in chapter 3 the durations of the press-release action in the target abdominal palpation examina- tion tasks are below this threshold. In the next chapter, the proposed measurement technique is used to create a ground truth model for the target abdominal palpation tasks with help of the medical experts. Medical students’ learning progress were evaluated with help of this technique in chapter 7 as compared to the developed gold standard model.